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LASERS IN OPTHALMOLOGY BASIC, DIAGNOSTIC AND SURGICAL ASPECTS A REVIEW
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Author and Author
In memoriam: you all died too soon Verena Fankhauser Simon Kugler Didier Riguin Pascal Rol
Cogito, ergo sum
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Title
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LASERS IN OPHTHALMOLOGY BASIC, DIAGNOSTIC AND SURGICAL ASPECTS A REVIEW
edited by Franz Fankhauser and Sylwia Kwasniewska
Kugler Publications/The Hague/The Netherlands
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Author and Author
ISBN 90 6299 189 0
Distributors: For the U.S.A. and Canada: Pathway Book Service 4 White Brook Road Gilsum, NH 03448 U.S.A. Telefax (603) 357 2073 For all other countries: Kugler Publications P.O. Box 97747 2509 GC The Hague, The Netherlands Telefax (+31.70) 3300254 E-mail:
[email protected] website: kuglerpublications.com
© Copyright 2003 Kugler Publications All rights reserved. No part of this book may be translated or reproduced in any form by print, photoprint, microfilm, or any other means without prior written permission from the publisher. Kugler Publications is an imprint of SPB Academic Publishing bv, P.O. Box 97747, 2509 GC The Hague, The Netherlands
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Table of contents
v
Table of contents Foreword R. Ritch
ix
Preface F. Fankhauser and S. Kwasniewska
x
Ophthalmic Laser Safety D.H. Sliney
1
The Purposes of Surgery G.L. Spaeth
11
Contact Lenses for Ophthalmic Laser Treatment E. Stefánsson and F. Fankhauser
15
Fundamentals of Optical Fibers M.J. Poulain
27
On the Application of Optical Fibers in Ophthalmology P.F. Niederer
33
Laser Speckle T. Halldórsson
43
Laser Doppler Techniques in Ophthalmology. Principles and Applications Ch.E. Riva and B.L. Petrig
51
Principles of Optical Coherence Tomography C.K. Hitzenberger
61
From Physical Energy to Biological Effect: How Retinal Laser Treatment Affects Diabetic Retinopathy E. Stefánsson
73
High-Resolution Multiphon Imaging and Nanosurgery of the Cornea Using Femtosecond Laser Pulses K. König
79
Selective Absorption by Melanin Granules and Selective Cell Targeting C.P. Lin
91
Mechanisms of Short-Pulsed Plasma-Mediated Laser Ablation and Disruption A. Vogel
99
The First Clinical Application of the Laser C.J. Koester and C.J. Campbell
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Selective Retinal Pigment Epithelium Laser Treatment J. Roider, R. Brinkmann and R. Birngruber
119
Confocal Microscopy of the Eye C.J. Koester
131
Imaging in Ophthalmology J.S. Schuman, Z.Y. Williams, J.G. Fujimoto and L.A. Paunescu
143
Different Methods of Refractive Surgery. The Advantages and Risks, and Their Relationship to Professional Ethics and Morals B.M. Tengroth
153
Corneal Laser Surgery for Refractive Corrections M. Mrochen, M. Bueeler and T. Seiler
159
Selective Laser Trabeculoplasty M.A. Latina and D.H. Gosiengfiao
171
Photocoagulation, Transpupillary Thermotherapy and Photodynamic Therapy for Choroidal Neovascularization R.S.B. Newsom, A.H. Rogers and E. Reichel
175
Photodynamic Therapy: Basic Principles and Mechanisms H. Van den Bergh and J.-P. Ballini
183
The Concept and Experimental Validation of Photodynamic Therapy in Neovascular Structures in the Eye R. Birngruber
197
Photodynamic Therapy: Clinical Status U. Schmidt-Erfurth and S. Michels
205
Controversial Aspects of Photodynamic Therapy K. Mori, D.M. Moshfeghi, G.A. Peyman and S. Yoneya
217
Lasers in Diabetes R.A. Stolz and A.J. Brucker
229
Retinal Photocoagulation with Diode Lasers R. Brancato, P.G. Gobbi, R. Lattanzio
241
Central Serous Chorioretinopathy A.P. Ciardella, S.J. Huang, D.L.L. Costa, I.M. Donsoff and L.A. Yannuzzi
255
Scanning Laser Polarimetry of the Retinal Nerve Fiber Layer in the Detection and Monitoring of Glaucoma C. Bowd, L.M. Zangwill and R.N. Weinreb
277
The Glaucomatous Optic Nerve Staging System with Confocal Tomography R. Sampaolesi and J.R. Sampaolesi
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Table of contents
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Principles of Photodisruption J.M. Krauss
303
Ultrastructual Effects of Laser Irradiation at the Anterior Chamber Angle E. Van der Zypen
315
Erbium:YAG Laser Trabecular Ablation T.S. Dietlein and G.K. Krieglstein
333
Laser Cyclodestructive Procedures of the Ciliary Body G.P. Schwartz, L.W. Schwartz and G.L. Spaeth
341
Laser Uveoscleroplasty: Basic Mechanisms and Clinical Experience S. Okisaka, K. Miyazaki, K. Morimoto, A. Mizukawa and Y. Sai
353
Transpupillary Laser Phototherapy for Retinal and Choroidal Tumors. A Rational Approach P. Rol(†)
363
Lasers in Intraocular Tumors G. Anastassiou and N. Bornfeld
377
Erbium:YAG Laser Vitrectomy M. Mrochen and T. Seiler
387
Lasers in Small-Incision Cataract Surgery J.M. Dodick and I.A. Pahlavi
395
Some Applications of the Neodymium:YAG Laser Operating in the Thermal and Photodisruptive Modes. Vitreolysis S. Kwasniewska
403
The Neodymium:YAG Laser in Strabismus and Plastic Surgery of the Face. Wound Repair F. Fankhauser
415
Hemostasis, Hemodynamics, Photodynamic Therapy, Transpupillary Thermotherapy: Controversial Aspects F. Fankhauser and S. Kwasniewska
429
Lasers in Lacrimal Surgery K. Müllner, T. Hofmann, G. Lackner and G. Wolf
441
Index
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Foreword It is impossible to imagine ophthalmology today without lasers, so ubiquitously and thoroughly do they dominate the field. Just 25 years ago, virtually all of what we rely upon today was in its infancy. Not only were lasers new on the scene, but also computers, intraocular lenses, and microsurgical instrumentation. Intracapsular cataract extraction was the standard, and ocular imaging devices, gene therapy, and stem cells had not yet even been conceptualized. A rapid explosion of argon laser techniques occurred in the late 1970s and early 1980s, and during this time, laser iridotomy, peripheral iridoplasty, and trabeculoplasty brought revolutionary changes to the approach of both angleclosure and open-angle glaucoma, while panretinal photocoagulation did the same for diabetic retinopathy. Neodymium:YAG laser capsulotomy and iridotomy were developed in the early 1980s. In the 1990s, another explosion occurred in the treatment of posterior segment disorders, including macular degeneration and intraocular tumors. The development of lasers for plastic surgery, cataract extraction, and ocular imaging is progressing rapidly and is expected to find much greater use and usefulness in the coming years. Professors Fankhauser and Kwasniewska have brought together many leading experts from different subspecialties in ophthalmology and laser physics to provide a comprehensive overview of the status of the broad range of laser applications at the present time. Professor Fankhauser has investigated the uses of laser in ophthalmology for over 30 years. Professor Kwasniewska joined him as a collaborator in the 1980s. After a brief history of laser applications, the first section of this book deals with the fundamentals of laser optics and principles of various imaging devices. The theory and clinical applications of lasers in corneal surgery, glaucoma, tumors, and vitreoretinal disease receive extensive coverage. Newer uses of the laser for cataract surgery, strabismus, plastic surgery, and lacrimal surgery are discussed in the final chapters. Ten years ago, the fundamentals of many of the chapters in this book had not yet been formulated. It will be interesting to see where we will stand ten years hence.
Robert Ritch, MD
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Preface
Preface One of the advantages of a long life is that it allows you to look at the research work of more than one generation. As far back as the 1950s, research in ophthalmology using intense light sources was mainly based on the radiation emitted by the high pressure xenon arc lamp of the Meyer-Schwickerath apparatus. Despite the many restrictions, many insights into light-tissue interaction phenomena were obtained. In the first half of the 1960s, the ruby laser made its debut and was applied to ophthalmic research by Charles Campbell, M. Catherine Rittler and Charles J. Koester. Due to the appearance of an ever-increasing number of laser energy sources, research in ophthalmology, with regard to basic, diagnostic, and surgical aspects, has been gaining momentum ever since. Lasers in Ophthalmology, Basic, Diagnostic and Surgical Aspects: A Review bears witness to the value of collaboration between physical scientists and medical researchers. Thanks to the ingenuity of a great number of dedicated research workers, basic research into the biology of and the application of the laser in ophthalmology has assumed an ever more important role, starting from the 1970s and continuing to run at breathtaking velocity ever since. George L. Spaeth writes (this volume) that there is only one appropriate ultimate goal of surgery: specifically, the restoration, maintenance, and enhancement of the health of the patient. This book pursues these ideals and its various chapters are basically oriented towards providing a solid foundation for restoring, maintaining, and enhancing health. We would like to thank our many friends and co-workers who have given their very best in making this progress possible. Without them, our research efforts would surely have failed. Sadly, not all of them are still among us today, but they will never be forgotten. Our sincere thanks go to Hans Bebie, George Benedek, Ulrich Dürr, Hans Giger, Jay M. Enoch, Martin Frenz, Hans Goldmann, Pierre-David Henchoz, Willi Hess, Alfred Jenni, Hans König, Pierre Koch, Klaus Meyer, Peter Niederer, Jean-Marcel Piffaretti, Jaroslav Ricka, Didier Riquin, Pascal Rol, Philippe Roussel, Eugen Van der Zypen, Heinz Weber, Martin Zulauf, and many others. Our special thanks go to Alice Gerber, who generously supported our research efforts, and to Dr Hans Pratisto whose adminstration of the electronic data pool was an enormous help. We are also immensely grateful to Peter Bakker and his crew, Ineke Ris, Simon Bakker, and Gay Wylie, who never shied away from any hard work or expense in order to bring this project to a successful conclusion. Franz Fankhauser Sylwia Kwasniewska
Lasers in Ophthalmology – Surgical and Diagnostic Aspects, pp. 000–000 edited by F. Fankhauser and S. Fankhauser-Kwasniewska © 2002 Kugler Publications, The Hague, The Netherlands
Ophthalmic laser safety
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Ophthalmic laser safety David H. Sliney Laser/Optical Radiation Program, US Army Center for Health Promotion and Preventive Medicine, Aberdeen Proving Ground, MD, USA
Keywords: safety, ocular hazard, protection
Introduction A variety of laser systems is used in ophthalmic applications. In each application, there are potential laser hazards to both the patient and the clinician. In most applications these potential hazards are minimal, but it is well to remember that, if a laser is used for altering tissue by photocoagulation, photodisruption or laser ablation, the exposures, if misdirected, are almost by definition potentially hazardous to other biological tissues as well. In the special terminology of laser safety, virtually every surgical laser is known as ‘Class 4’ and will be considered by safety specialists as being very hazardous. Therefore, it is important for ophthalmologists who use lasers to be familiar with laser safety guidance and terminology, in order to have a balanced view of what might be perceived as a very hazardous laser and what procedures require serious attention in order to take appropriate precautions. Safety standards worldwide group laser products into at least four different safety categories of risk, known as ‘hazard classes’. These range from Class 1 products, which pose no hazards, to Class 2, which are visible-wavelength lasers that are no more dangerous than a bright lamp, Class 3, which are significant hazards to the eye, and Class 4, which are hazardous to both skin and eye and are readily capable of cutting or photocoagulating biological tissue. From a physical standpoint, lasers owe their particular usefulness in most applications to their extraordinarily high brightness, and this factor, known technically as ‘radiance’ also leads
to their significant hazard (Fig. 1). A laser light source is millions of times brighter than an ordinary incandescent lamp, and even brighter than an arc lamp or the sun. Although, laser light is generally monochromatic and has spatial and temporal coherence, quite unlike any light from a conventional light source, these physical characteristics contribute little to the hazard. Coherence and monochromaticity can be important in some diagnostic applications, but are not very significant in terms of the surgical value or laser hazards. Laser parameters As noted above, it is the high radiance of lasers that leads to their significant ocular hazards, but this radiometric quantity is seldom actually specified, and is only indirectly used in laser safety assessments. Instead, the more familiar laser output parameters that determine the safety or hazard classification are: the wavelength, pulse duration and energy, if pulsed; and wavelength and power, if continuous wave (cw). The accessible emission limits (AELs) that determine the laser safety classes also vary with spectral band, e.g., ultraviolet (100400 nm), visible (400-780 nm) and infrared (780 nm-1 mm), since the biological risk varies with wavelength. Laser wavelength determines how effectively light is absorbed in the target tissue and how effectively light penetrates overlying media to reach a tissue target (transmission).1,2 Pulse durations can range from tens of femtoseconds (10-15 seconds) to milliseconds (msec). From
The opinions or assertions herein are those of the author and should not be construed as official policies of the US Department of the Army or Department of Defense. Address for correspondence: David H. Sliney, PhD, Laser/Optical Radiation Program, US Army Center for Health Promotion and Preventive Medicine, Aberdeen Proving Ground, MD 21010-5422, USA. e-mail:
[email protected] Lasers in Ophthalmology – Basic, Diagnostic and Surgical Aspects, pp. 1–10 edited by F. Fankhauser and S. Kwasniewska © 2003 Kugler Publications, The Hague, The Netherlands
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Fig. 1. Radiance. The great value of a laser source in most surgical applications is its very high ‘brightness’ or radiance, which permits laser light to be focused on an exquisitely small spot at very high concentrations of power-per-unit-area (irradiances). Only in photodynamic therapy can a low radiance source be used (but for longer-duration exposures).
Fig. 2. Radiometric terms. Several radiometric terms employed in photobiology are illustrated. Note that fluence and radiant exposure, as well as fluence rate and irradiance, have the same radiometric units of J/m² and W/m², respectively, but these are not equivalent concepts. Fluence and fluence rate incorporate backscatter, and do not incorporate a cosine function in the formal definition.
the point-of-view of laser safety, lasers that emit for any period greater than 0.25 seconds are considered to be cw.3 Power (in watts) describes the rate at which energy (in joules) is emitted from a laser or delivered to a target tissue (power = energy/time, in watts = joules/sec). Irradiance is the incident laser power per unit area delivered to a
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target surface (irradiance = power/area, in watts/ cm²). If the backscattered irradiance is added and a cosine factor ignored, this is referred to as ‘fluence rate’ in watts/cm² (Fig. 2). Irradiance is sometimes referred to as ‘power density’ at the target surface (but this is technically incorrect). And, for pulsed lasers, it is often useful to describe laser exposures
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Ophthalmic laser safety in terms of radiant exposure rather than irradiance. Radiant exposure is the energy divided by the exposed surface area (radiant exposure = energy/area, in joules/cm²). Fluence is widely used as a synonym for radiant exposure, but technically it should be reserved for situations in which an exposed surface encounters both forward and back-scattered light.3 The concepts of fluence and fluence rate are important in dosimetry for photodynamic therapy, where multiple scattering and diffusion in the target tissue are of great importance for redistributing and homogenizing the incident light. Ocular hazards Normally, the eye is well adapted to protecting itself against optical radiation (ultraviolet, visible, and infrared radiant energy) from the natural environment. The brow ridge and upper lids greatly shield the eyes from overhead solar radiation,4 and the aversion response with blink reflex and eye movements limit direct viewing of the sun and welding arcs to a fraction of a second, and therefore preclude photoretinitis.5-7 However, when a patient is anesthetized, the normal aversion responses to bright light or heat may be greatly reduced – or absent – in the cornea and skin. The normal aversion response to a heat sensation of exposed skin will normally limit potentially hazardous thermal exposure of the skin if the injury threshold is not reached within 0.2-0.3 seconds.8 Furthermore, some body movements (e.g., eye movement) will not limit exposure duration, and injury that would normally not occur can result to the tissues. Vascularized tissues can carry away excess heat and limit the possibility of thermal injury of tissue (e.g., the choroid), but this is a limited capability that can be readily overtaxed by laser exposure or if vascular flow is impaired. During ophthalmic examination, a cooperative patient may willingly be exposed to extremely discomforting light in order to cooperate with the examining clinician.9,10 It is generally well accepted that the eye is more susceptible to optical hazards than the skin. Although both skin and eye are susceptible to injury from lasers and other intense optical sources, protection of the eye is central because of the potential for loss of vision. Therefore, it is important to recognize any potential hazards of intense light sources used in diagnosis and surgery. Optical radiation hazards There are at least nine separate types of hazards to the eye and skin from lasers and other intense optical radiation sources, and protective measures must be chosen with an understanding of each of these. One or more of the following effects may pose a potential hazard, depending upon the laser wave-
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3 length and the temporal and geometrical characteristics of the exposure:5,11,12 (a) ultraviolet photokeratoconjunctivitis (also known as ‘welder’s flash’ or simply ‘photokeratitis’, one aspect of ‘snow-blindness’ from wavelengths of ~180-400 nm);13 (b) ultraviolet cataract (~295-325 nm – and perhaps to 400 nm). This is expected only from chronic exposure under normal conditions;13-15 (c) ultraviolet erythema (~200-400 nm); (d) skin cancers arising from chronic exposure to ultraviolet radiation, particularly from UV-B (280-315 nm), but also demonstrated for UVA (315-400 nm);16,17 (e) thermal injury to the retina (400-1400 nm). Normally this type of injury is only possible from lasers, a focused, very intense xenon-arc source, or the nuclear fireball. A local burn of the retina results in a blind spot (scotoma). Because of heat conduction from the retinal image, a very intense exposure delivered within seconds normally is required to cause a retinal coagulation, otherwise surrounding tissue conducts the heat away from the retinal image. This leads to an image-size dependence of retinal thermal injury;2,5 (f) blue-light photochemical injury to the retina (principally 400-550 nm blue light);18 ‘blue light’ photoretinitis, e.g., solar retinitis and welder’s maculopathy, which may lead to a permanent scotoma. Prior to conclusive theoretical work19 and animal experiments only two decades ago,20 solar retinitis was thought to be a thermal injury mechanism. Unlike thermal injury, there is no image-size dependence as with (e) above; (g) near-infrared thermal hazards to the lens (approximately 800-3000 nm) with a potential for industrial heat cataract.21,22 The average corneal exposure from infrared radiation in sunlight is of the order of 10 W/m². In comparison, glass and steel workers exposed to infrared irradiances of the order of 0.8-3 kW/ m² daily for ten to 15 years have reportedly developed lenticular opacities.5 These spectral bands include IR-A (700-1400 nm) and IR-B (1.4-3.0 µm); (h) thermal injury of the cornea and conjunctiva (approximately 1400 nm to 1 mm). This type of injury is almost exclusively limited to pulsed, or very brief, laser radiation exposure;5 (i) thermal injury of the skin (approximately 400 nm to 1 mm). This type of injury rarely occurs from conventional optical sources, since the aversion to thermal pain (occurring at a temperature of 45°C or greater) will normally limit exposure to a few seconds. Laser-induced thermal injury is possible from most Class 4 lasers.
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Table 1. Selected occupational exposure limits (MPEs) for some lasers* Type of laser
Principal wavelength(s)
MPE (eye)
Argon-fluoride Xenon-chloride Argon ion Copper vapor Helium-neon Gold vapor Krypton ion Neodymium:YAG (primary λ)
193 nm** 308 nm 488, 514.5 nm 510, 578 nm 632.8 nm 628 nm 568, 647 nm 1064 nm
3.0 mJ/cm² over 8 hours 40 mJ/cm² over 8 hours 3.2 mW/cm² for 0.1 seconds 2.5 mW/cm² for 0.25 seconds 1.8 mW/cm² for 1.0 seconds 1.0 mW/cm² for 10 seconds
Neodymium:YAG laser (secondary λ) Pulsed Nd:YAG (1.44 µm) Pulsed holmium cw holmium
1334 nm 1.44 µm 2.1 µm 2.1 µm
cw carbon monoxide
~5 µm
Carbon dioxide
10.6 µm
5.0 µJ/cm² for 1 nsec to 50 µsec no MPE for t < 1 nsec 5 mW/cm² for t >10 seconds 40 µJ/cm² for 1 nsec to 50 µsec 40 mW/cm² for > 10 seconds 0.1 J/cm² for 1 nsec to 1 msec
100 mW/cm² for 10 seconds to 8 hours, limited area 10 mW/cm² for t > 10 seconds for most of body (skin surface) same as 2.1 and 5 µm wavelengths above
*All standards/guidelines have MPEs at other wavelengths and exposure durations **Sources: ICNIRP (2000); IEC 60825-1.2-2001; ANSI Z136.1-2000 Note: to convert MPEs in mW/cm² to mJ/cm², multiply by exposure time t in seconds, e.g., the He-Ne or argon MPE at 0.1 seconds is 0.32 mJ/cm².
The importance of wavelength and time of exposure Thermal injuries (e) and (h) above are generally limited to very brief exposure durations, and eye protection is designed to prevent these acute injuries; the laser-tissue interaction mechanism is put to use in (e) retinal photocoagulation. However, photochemical injuries such as (a) and (c) are possible from low dose rates spread over minutes or even hours as a result of the Bunsen-Roscoe law. Most photochemical effects are limited to a narrow range of wavelengths known as an ‘action spectrum’, whereas a thermal effect can occur at any wavelength in the spectrum. Still other photochemical interaction mechanisms besides (f) exist that can produce retinal injury, but these are generally thought to be only theoretical in the case of human exposure. Noell42, Lawwill43, Kremers and van Norren44, and others described retinal damage from light produced by relatively low-brightness fluorescent lamps, and these mechanisms were only detectable for exposure durations extending beyond two hours and repeated for several days. These effects are clearly related to excessive over-stimulation of the photoreceptors. Safety standards Laser safety standards for medical applications exist in many countries (including Australia, Great Britain, Germany, and the USA). The basic guidance is very similar in all these standards. In addition, the IEC has a technical note (IEC 608258-1999-11) on the safe use of medical lasers. In the USA, there are two safety standards which apply to
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medical lasers: the American National Standard for the Safe Use of Lasers in the Health Care Environment, ANSI Z136.3-1996, which is a voluntary, consensus standard for users,23 and a Federal Regulation that applies to laser manufacturers (21 CFR1040), issued by the Food and Drug Administration (FDA).24 The latter regulatory standard imposes certain labelling, informational, and design performance requirements upon the manufacturer and all laser products (medical or non-medical) must meet these. Examples of the FDA performance requirements are the laser emission indicator and the key switch. Although many occupational safety standards and guidelines (e.g., the ANSI Standard in the USA) are ‘voluntary’, the guidelines can have a legal impact if labor, insurance or health-service regulations refer to guidelines when questions arise as to the adequacy of the safety in a facility. If an accident were to occur, these types of standards would surely be referred to in any litigation. For an ophthalmic laser facility, most standards basically require that one person be given safety oversight responsibility, and that person is usually termed the ‘laser safety officer’, or simply, the LSO; in the UK, this is the ‘laser safety advisor’. The guidelines also recommend that the laser operator be trained in the safe use of the laser device, that a laser safety warning sign be placed at the entrance to the laser treatment room, and that persons assisting wear laser eye protection within the range of potentially hazardous exposure. This region – where diffuse reflections and stray beams could be hazardous – is termed the ‘nominal ocular hazard area’ (NOHA) in some international standards, or the ‘nominal hazard zone’ (NHZ) in the USA.23 A
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Ophthalmic laser safety key element in all laser standards is the risk assignment of the laser to a ‘hazard class’. Class 1 products may be thought of as ‘eye-safe’ lasers; Class 2 is a 1-mW (or less) visible laser (e.g., an aiming beam); Class 3 is a significant eye hazard (e.g., an Nd:YAG photodisruptor); and Class 4 is a skin hazard as well (e.g., all photocoagulators capable of exceeding 500 mW total average power). Hazard control measures are assigned to each class by the ANSI standard.3,5,23 Occupational exposure limits Guidelines for limiting human exposure (both eye and skin) to laser radiation have been issued by the International Commission on Non-Ionizing Ra-
5 diation Protection (INCIRP) and by other national and international standardizing groups following the ICNIRP guidelines.1,2,25-29 Table 1 lists permissible occupational maximum permissible exposure (MPE) limits for some of the commonly used surgical lasers. Reflections and probability of exposure An examination of laser accident records indicates that the source of accidental ocular exposure is most frequently a reflected beam. Figure 3 illustrates the types of mirror-like (specular) laser beam reflections that can occur from the flat or curved surfaces, which are characteristic of contact lenses
Fig. 3. Reflections from contact lenses. The diagram shows the reflected beams diverging from a contact lens during laser photodisruption or trabeculoplasty. The hazard distance is normally less than 1-2 m from the lens.
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Fig. 4. Tissue interaction.
or of the metallic instruments used in some surgical procedures. Skin injury of the hand holding an instrument is also possible. Normally, the collimated beam is considered the most hazardous type of reflection, but, at very close range, a diverging beam may pose a greater likelihood of striking the eye.3,8,9 A number of steps can be taken to minimize the potential hazards to both the patient and the surgical staff. Preventive measures will depend upon the type of laser. Since laser wavelengths in the ultraviolet and infrared spectral regions are invisible, the presence of hazardous secondary beams could go unnoticed. This added hazard resulting from an infrared laser beam’s lack of visibility is common to the 2.1 µm holmium or the 1064 nm Nd:YAG laser. In contrast, the argon and the second-harmonic Nd:YAG (sometimes referred to as the ‘KTP’) lasers emit highly visible, blue-green (488, 514.5, and 532 nm) beams and, in some respects, pose a lesser potential hazard. Most current surgical lasers, such as the Nd: YAG, holmium, diode or argon, are cw, or nearly so. In contrast, the single-pulse ophthalmic laser photodisruptors or some excimer ablative lasers emit very short pulses. The biological effects and potential hazards from high-peak, power-pulsed lasers are quite different from those of cw lasers. This is particularly true of lasers operating in the retinal hazard region of the visible (400-760 nm) and near-infrared spectrum (IR-A: 760-780 to 1400 nm), as shown in Figure 4. The severity of retinal lesions from a visible or near-infrared (IR-A) cw laser is normally considered to be far less than from a Q-switched laser. Another major factor that influences the potential hazard is the degree of beam collimation. Almost all surgical lasers are focused,
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thereby limiting the hazardous area (referred to as the ‘nominal hazard zone’ in IEC 60825-1 and ANSI Z136.1-2000.1 An exception is the highly collimated beam from many lasers with articulated aims, which may remain hazardous at some distance from the instrument.8 Reflections are most serious from flat mirror-like (specular) surfaces – characteristic of many metallic surgical instruments. Many surgical instruments now have black anodized or sandblasted, roughened surfaces to reduce (but not eliminate) potentially hazardous reflections. The strong curvature and surface roughening spread the reflected energy and greatly reduce the reflection hazard. The surface roughening is generally more effective than the black (ebonized) surface, since the beam is diffused. However, in some cases, combining a special black surface with roughening provides increased protection, and adding a black polymer finish, has been shown by experiment measurements to offer the greatest protection at the CO2 wavelength – despite initial scepticism by investigators.9 However, other groups argue against blackening the surface, since the instrument will become hotter than without for visible wavelengths. Therefore, the use of the special blackened surfaces must be approached with caution for each application. It should be noted that both the surface finish and reflectance seen in the visible spectrum do not indicate those qualities in the invisible, far-infrared spectrum. In fact, a roughened surface that appears to be quite dull and diffuse at a shorter, visible, or IR-A wavelengths, will always be more specular at far-infrared wavelengths (e.g., at 10.6 µm). This results from the fact that the relative size of the microscopic structure of the surface relative to the incident wavelength determines whether the beam
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Ophthalmic laser safety
Fig. 5. The beam irradiance decreases rapidly with distance from a bare fiber.
is reflected as a specular or diffuse reflection.3,9,30 A specularly reflected beam with only 1% of the initial beam’s power can still be quite hazardous. Hence, the rougher the surface of an instrument likely to intercept the beam, the safer the reflection. For example, even a 1% reflection of a 40-W laser beam is 400 mW! It is somewhat surprising that there have been few cases reported of eye injuries to residents and other persons observing Nd:YAG laser surgery without eye protectors. Hazardous specular reflections from a laser beam emerging from an endoscopic optical fiber are limited in extent because the beam rapidly diverges as shown in Figure 5. Most invisible beam surgical lasers and pulsed lasers have a visible alignment beam. Infrared lasers usually make use of a low-power coaxial HeNe (632.8 nm) or diode (e.g., 635 nm) red laser. Where feasible, it is desirable for this alignment beam to be 1 mW or less, since the maximum cw, visible laser beam power that can safely enter the eye within the aversion response (i.e., within the blink reflex, etc., of 0.25 seconds) is 1 mW. This type of laser is then classified as Class 2, and poses a very low risk to the user.
7 tion of heat. Details of accidents are often not published because of litigation, but anecdotal reports indicate that misfiring of a laser when not in use, or undetected breakage of optical fibers, has led to fires from the ignition of surgical drapes, with serious injury to patients. Procedural methods, such as the use of the standby switch or proper placement of the laser foot-switch, can reduce the number of such accidents, but never completely eliminate them. Preparations for extinguishing fires or the moistening of drapes must always be part of the OR safety standing operating procedure (SOP). Accidental injury to the eye is of particular concern when lasers are used for non-ophthalmic procedures near the eyes. Where exposure of the eye itself is not intended, special eye shields are available for patient protection, such as that shown in Figure 6. Safety of the surgeon Normally, the surgeon views the target issue through the optics of an endoscope, operating microscope, colposcope, slit-lamp biomicroscope, etc., and the reflections are safely attenuated within the optics. Under such indirect viewing conditions, the surgeon or laser operator is not normally highly susceptible to injury, due to the proper design of the laser instrument. However, if the laser is accidentally actuated when the surgeon is not looking through the viewing optics, he or she will be just as much at risk as any other person in the room. Additionally, with hand-held laser delivery systems, it should be remembered that the surgeon’s hand is the closest to the laser target and therefore it is closest to potentially hazardous reflections from adjacent surgical instruments (e.g., metal retractors).
Patient safety
Safety of the surgical staff
Most laser safety regulations do not apply to the exposure of the patient at the target site for surgery. However, accidental exposure to the patient from misdirection of the laser beam should be of concern, and can result in injury of eye and skin.3 Ignition of drapes can be particularly hazardous to the patient, who is under anesthesia and unable to warn the operating-room (OR) staff of the sensa-
Nurses, surgical assistants, and other assisting staff are potentially exposed to misdirected laser beams. Lasers have been accidentally initiated when the beam delivery system was directed other than at the patient, a foot switch was accidentally pressed, or similar errors have occurred, and the beam directed at a person. In panretinal photocoagulation, an unexpected eye movement has led to an unin-
Fig. 6. Examples of commercial eye protectors used in laser surgical procedure near the eye. (Photographs courtesy of OculoPlastik, Inc., Montreal, Canada.)
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tentional exposure of the macula. Accidental firing of a laser has also occurred because of confusion created by having more than one foot switch positioned nearby. Some safety guidelines recommend that the laser foot switch be covered and clearly identified. Assistants are potentially exposed to secondary reflections from contact lenses (and in some special procedures, from reflections from surgical instruments), whereas the surgeon’s eyes are protected by filtration in the viewing optics. Reflections from the cornea or from the contact lens used in ophthalmic surgery have been shown to be potentially hazardous to assistants or bystanders in line of view of the contact lens to a distance of 1-2 m. The operating microscope used in laser microsurgery by a number of specialties would protect the eyes of the surgeon if properly designed, whereas, assistants, and bystanders may be exposed to potentially hazardous reflections from surgical instruments inserted into the beam path. Safety of other bystanders Bystanders in the surgical facility or outpatient laser facility who are present to observe or to calm the patient (e.g., a relative) may be susceptible to exposure from reflected laser beams in the same manner as a surgical assistant or nurse. In addition, because of lack of training or knowledge about the laser surgical procedure, bystanders may be at greater risk by inadvertently placing themselves in a dangerous position. Those individuals should always be provided with laser eye protectors. Service personnel Service personnel are particularly susceptible to laser injury since they often gain access to collimated laser beams from the laser cavity itself or from opening up the beam delivery optics to gain access to collimated laser beams prior to the beam focusing optics or fiber-optic beam delivery system. Once the laser beam leaves the delivery system and comes rapidly into a focus, it then diverges again, or if emerging from a fiber, it also rapidly diverges. The zone where the beam is concentrated to a level sufficient to pose severe hazard to the eyes or skin (the NHZ) is normally a limited zone of 1-2 m near the beam focal point. However, a collimated laser beam, as the raw beam for most laser cavities, or a specular reflection from a turning mirror or Brewster window in the laser console may be emitted from the laser cabinet (protective housing) when the service person gains access. Several serious eye injuries have occurred to service personnel exposed to secondary, collimated, invisible 1063 nm Nd:YAG laser beams when the service personnel gained access to the laser cavity.
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Protection Photocoagulators Laser photocoagulators with slit-lamp delivery systems provide a fixed beam with limited directional movement. Ideally, to minimize the chance of anyone intercepting the primary or reflected beam, the laser delivery system is directed at a wall to terminate the beam in case a patient is not in position, reflections back from the contact lens are terminated at the opposite wall and not directed toward the door (Fig. 3). Cw argon, krypton, and diode laser photocoagulators may have a collimated beam emerging from the laser (as with an articulated-arm delivery system), and the open beam can in theory be hazardous for some distance greater than the clinical treatment room. However, with most devices, the beam emerges as a diverging beam and the hazardous distance may be only 2-3 m when a patient is not in the beam. The reflection from any contact lens will produce potentially hazardous secondary, specularly reflected beams as shown in Figure 3 to a distance of 0.3-2 m, depending upon the laser system. Therefore, it is important that persons observing or holding the patient be given laser eye protectors.5 The operator of the laser is protected by a laser safety filter that is permanently built into the viewing optics. This may be fixed as in the case of the diode laser, or part of a shutter system as in the case of argon and krypton lasers. There have been some instances in the past of filter shutter failure, but modern photocoagulators have effectively eliminated this potential. The visible reflection of the aiming beam between therapeutic exposures has been measured, and is below permissible exposure limits.31 Of particular note is the potential hazard of attaching auxiliary viewing optics, unless laser safety filters are installed. The laser photocoagulator system employing an indirect ophthalmoscopic delivery system poses more problems than the slit-lamp delivery systems. The beam can be directed anywhere, and a momentary misfiring can be directed away from the patient. Hence, laser eye protectors are mandatory for any ancillary personnel or visitors in the treatment room. A warning light or sign should be displayed during laser use. Photodisruptors Nd:YAG photodisruptors have very large convergence angles, such that the hazard distance along the primary beam is only 1-2 m from the focus if the patient were not in the beam.32 With the patient in position during treatment, the specular reflections from the contact lens are potentially hazardous, at least out to a distance of 0.3-1 m, as shown in Figure 3.32 Therefore, it is important that persons observing or holding the patient be given laser eye protectors. The operator of the laser is pro-
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Ophthalmic laser safety tected by a laser safety filter built into the viewing optics. As with the photocoagulator, the photodisruptor should be positioned so that neither the direct beam nor the reflections from the contact lens can be directed toward the doorway to the treatment room. Excimer laser photoablators As with slit-lamp delivery systems, the fixed delivery system in ArF excimer lasers used in corneal ablative procedures has a stable beam path, and the nominal ocular hazard area is very limited. In fact, there is virtually no reflection hazard during the procedure and the need for eye protection for persons in the treatment room can be reasonably questioned. The ocular hazard in the room is only theoretical, but from a legal standpoint, it may be advisable to offer visitors clear plastic visitor’s goggles. Generally, the greatest safety concerns with excimer lasers relate to the safe handling of the excimer laser gasses, such as fluorine, and not the direct laser beam.33 Fortunately, today, ArF excimer lasers are well designed and only premixed gasses are used. Illuminators Not all potential hazards stem from laser radiant energy. The illumination from operating microscopes can pose a potential hazard to the retina of the patient. Studies show that the tungsten filament of an operating microscope imaged on the retina can lead to photoretinitis.34-37 This risk is significant only in cataract surgery if the illuminators’ image is fixed on the macular area for a period exceeding at least 15 minutes. Other ophthalmic diagnostic lights are far less hazardous.38-41 Conclusions Ophthalmic lasers pose potential hazards which are well understood by the ophthalmologist. Fortunately, the laser beams emitted by most ophthalmic lasers are stable and fixed, so apart from the placement of specular (mirror-like) surfaces, such as contact lenses, in the beam path, the nominal ocular hazard area is extremely limited. The greatest safety problem is introduced by a non-fixed beam delivery system such as the indirect ophthalmoscope. References 1. Boettner EA, Wolter JR: Transmission of the ocular media. Invest Ophthalmol Vis Sci 1:776-783, 1962 2. Mainster MA: Wavelength selection in macular photocoagulation: tissue optics, thermal effects and laser systems. Ophthalmology 93:952-958, 1986 3. Sliney DH, Trokel SL: Medical Lasers and Their Safe
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9 Use. New York, NY: Springer Verlag 1993 4. Sliney DH: Eye protective techniques for bright light. Ophthalmology 90(8):937-944, 1983 5. Sliney DH. Wolbarsht ML: Safety with Lasers and Other Optical Sources. New York, NY: Plenum Publ Corp 1980 6. Gerathewohl SJ, Strughold H: Motoric response of the eyes when exposed to light flashes of high intensities and short durations. J Aviat Med 24:200-207, 1953 7. Sliney DH: Laser effects on vision and ocular exposure limits. Appl Occup Environ Hyg 11(4):313-319, 1996 8. Fisher KJ, Gehly J, Sliney DH: Surgical lighting: a review of ocular safety. J Illum Eng Soc (North Am) 20(1):28-31, 1991 9. Mainster MA, Ham WT, Delori FC: Potential retinal hazards: instrument and environmental light sources. Ophthalmology 90(8):927-931, 1983 10. Sliney DH, Wolbarsht ML: Safety standards and measurement techniques for high intensity light sources. Vision Res 20(12):1133-1141, 1980 11. World Health Organization (WHO): Environmental Health Criteria No 160, Ultraviolet Radiation. Joint publication of the United Nations Environmental Program, the International Radiation Protection Association and the World Health Organization. Geneva: WHO 1994 12. World Health Organization (WHO): Environmental Health Criteria No 23, Lasers and Optical Radiation. Joint publication of the United Nations Environmental Program, the International Radiation Protection Association and the World Health Organization. Geneva: WHO 1982 13. Pitts DG, Cullen AP, Hacker PD: Ocular effects of ultraviolet radiation from 295-365 nm. Invest Ophthalmol Vis Sci 16(10):932-939, 1977 14. Sliney DH: Physical factors in cataractogenesis: ambient ultraviolet radiation and temperature. Invest Ophthalmol Vis Sci 27(5):781-790, 1986 15. Taylor HR, West SK, Rosenthal FS, Munoz B, Newland HS, Abbey H, Emmett EA: Effect of ultraviolet radiation on cataract formation. New Engl J Med 319:1429-1433, 1988 16. Forbes PD, Davies PD: Factors that influence photocarcinogenesis. In: Parrish JA, Kripke ML, Morison WL (eds) Photoimmunology, Ch 7. New York, NY: Plenum Publ Corp 1982 17. Sterenborg HJCM, Van der Leun JC: Action spectra for tumorigenesis by ultraviolet radiation. In: Passchier WF, Bosnjakovic BFM (eds) Human Exposure to Ultraviolet Radiation: Risks and Regulations, pp 173-191. New York, NY: Excerpta Medica Division, Elsevier Science Publ 1987 18. Ham WT Jr, Mueller HA: The photopathology and nature of the blue-light and near-UV retinal lesion produced by lasers and other optical sources. In: Wolbarsht ML (ed) Laser Applications in Medicine and Biology, pp 191-246. New York, NY: Plenum Publ Corp 1989 19. White TJ, Mainster MA, Wilson PW, Tips JH: Chorioretinal temperature increases from solar observation. Bull Math Biophys 33:1-17, 1971 20. Ham WT Jr, Mueller HA, Sliney DH: Retinal sensitivity to damage from short wavelength light. Nature 260(5547): 153-155, 1976 21. Lydahl E: Infrared cataract. Acta Ophthalmol (Kbh) Suppl 166:1-63 (with six appendices), 1984 22. Pitts DG, Cullen AP: Determination of infrared radiation levels for acute ocular cataractogenesis. Graefe’s Arch Klin Exp Ophthalmol 217:285-297, 1981 23. American National Standards Institute (ANSI): Safe Use of Lasers in Health Care Facilities ANSI Standard Z136.31996. New York, NY: ANSI 1996 24. US Food and Drug Administration (FDA): Laser Performance Standard, Title 21, Code Federal Regulations, Part
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25.
26. 27. 28. 29.
30. 31.
32. 33.
34.
35.
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D.H. Sliney 1040 (21CFR1040). Washington, DC: Government Printing Office 1986 Mainster MA, Sliney DH, Belcher CD III, Buzney SM: Laser photodisruptors: damage mechanisms, instrument design and safety. Ophthalmology 90:973-991, 1983 Overheim RD, Wagner DL: Light and Color. New York, NY: Wiley 1982 Minnaert M: Light and Color in the Outdoors. New York, NY: Springer-Verlag 1993 Meyer-Arendt JR: Introduction to Classical and Modern Optics, 2nd edn. Englewood Cliffs, NJ: Prentice-Hall 1984 Mainster MA: Ophthalmic applications of infrared lasers: thermal considerations. Invest Ophthalmol Vis Sci 18:414420, 1979 Bursell SE, Mainster MA, Sliney DH: Spectral properties of ocular pigments. (Unpublished data) Sliney DH, Mainster MA: Potential laser hazards to the clinician during photocoagulation. Am J Ophthalmol 103:758-760, 1987 Wood RL, Sliney DH, Basye RA: Laser reflections from surgical instruments. Lasers Surg Med 12:675-678, 1992 Sliney DH: YAG laser safety. In: Trokel SL (ed) YAG Laser Ophthalmic Microsurgery, pp 67-84. Norwalk, NJ: Appleton-Century-Crofts 1983 Sliney DH: Safety of ophthalmic excimer lasers with an emphasis on compressed gasses. Refract Corneal Surg 7:308314, 1991 Byrnes GA, Antoszyk AN, Mazur DO, Kao TC, Miller SA: Photic maculopathy after extracapsular cataract sur-
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37. 38.
39. 40.
41.
42.
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44.
gery: a prospective study. Ophthalmology 99(5):731-737, 1992 Calkins JL, Hochheimer BF, D’Anna SA: Potential hazards from specific ophthalmic devices. Vision Res 20:10391053, 1980 Marshall J: Light damage and ageing in the human macula. Research and Clinical Forums 7(3):27-43, 1986 Sliney DH, Armstrong BC: Radiometric evaluation of surgical microscope lights for hazards analyses. Appl Opt 25(12):1882-1889, 1986 Sliney DH, Campbell CE: Ophthalmic instrument safety standards. Laser Light Ophthalmol 6(4):207-215, 1994 James RH, Bostrom RG, Remark D, Sliney DH: Handheld ophthalmoscopes for hazards analysis: an evaluation. Appl Opt 27:5072-5076, 1988 Young RSL, Goldberg MF, Fishman GA: The minimum retinal irradiance required for viewing human fundi in indirect ophthalmoscopy. Invest Ophthalmol Vis Sci 20:701704, 1981 Noell WK: Effects of environmental lighting and dietary vitamin A on the vulnerability of the retina to light damage. Photochem Photobiol 24(4):717-723, 1979 Lawwill T, Crockett S, Currier G: Retinal damage secondary to chronic light exposure. Documenta Opthalmologica 44(2):379-402, 1977 Kremers JJM, van Norren D: Retinal damage in macaque after white light exposures lasting ten minutes to twelve hours. Invest Opthal Vis Sci 30(6):1032-1040, 1989
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The purposes of surgery
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The purposes of surgery George L. Spaeth Wills Eye Hospital, Philadelphia, PA, USA
Keywords: medical ethics, morality of surgery, Hippocratic oath, megalomania in surgery
Before any surgical procedure is undertaken, the surgeon must be clear in his or her mind regarding why the surgery is being undertaken. So also must the patient be clear. That is, the doctor and the patient must know what the surgery is trying to accomplish. And they must be in agreement regarding the ultimate goal. This comment is so fundamental that it may seem condescending even to say it. Of course the purpose, the desired outcome, must be clear. But frequently the surgeon is so involved in the technical aspects necessary to achieve the ultimate goal that the goal itself fades from view. This chapter is intended to remind us all, but especially surgeons, that there is only one appropriate ultimate goal of surgery, specifically, restoration, maintenance, or enhancement of the health of the patient. Even patients themselves sometimes forget that. For example, they may want a drooping right upper lid lifted because they are dissatisfied with their appearance. But the dissatisfaction may have little to do with their actual appearance. The surgery that restores the right upper lid to its ‘proper’ position, then, no matter how perfect technically, may not help them at all, that is, may not restore, maintain, or enhance health. Patients with cataracts may be unhealthy because they are limited by the reduced vision. They feel incomplete, unhealthy. The strategy used to restore health in such a patient is to restore good vision. The mechanism by which this can be done is by removing the cataract and providing a tolerable refractive correction – such as an intraocular lens. The surgeon knows that merely removing the cataract will not make the patient happy. The problem is not the cataract, but the poor vision. But the patient does not know this. It is no wonder that, in years past, patients having had ‘successful’ intracapsular cata-
ract extractions were so often dissatisfied, whereas today patients whose hazy lenses have been replaced by clear intraocular lenses of the right strength to allow them to see without glasses are so often ecstatic following surgery. The major purpose of this chapter is to consider how other purposes of surgery get layered onto the basic purpose, that of restoring, maintaining, or enhancing health. Specifically, surgery is also undertaken: (1) for the purpose of making a living for the surgeon; (2) to teach others how to perform surgery; (3) to discover new knowledge; (4) to allow the surgeon become famous; (5) to have fun; (6) for the purpose of becoming rich; and (7) to achieve a specific goal, perhaps not directly related to improving the health of the patient. 1. Altruistic surgeons who perform surgery solely for the joy of helping people will not be able to help people for long unless they get some type of recompense that allows them to survive. It is entirely appropriate for the surgeon to expect and to receive some type of compensation from the person undergoing surgery, or from the society in which that person lives and that recognizes that a healthy, functioning populace is its best security. It is not appropriate, however, for surgeons to use patients’ vulnerability as a way of exploiting them for monetary gain. For care to be available, there must be caregivers and facilities, and these must be supported or else there will soon be no care. Someone has to pay. The person benefitting is the most appropriate person to pay. And the society benefitting is the next most appropriate source of funds to cover the inevitable costs of care. Individuals and societies need to consider how much they value health, and then provide the funds needed to ac-
Address for correspondence: George L. Spaeth, MD, Wills Eye Hospital, 900 Walnut Street, Philadelphia, PA 19107, USA. e-mail:
[email protected] Lasers in Ophthalmology – Basic, Diagnostic and Surgical Aspects, pp. 11–14 edited by F. Fankhauser and S. Kwasniewska © 2003 Kugler Publications, The Hague, The Netherlands
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complish that level. But surgeons need not be apologetic for wanting compensation for their work. Making a living is an appropriate purpose, but never the primary purpose. It is when placing the fee motivates the decision to undertake surgery that the surgeon is acting improperly. 2. Just as there would be no surgical care without surgeons, there would soon be none without teaching. But teaching takes time, and accordingly decreases the remuneration to the surgeon. Some schools or societies pay the teachers. Others think that teachers earn enough anyway and should teach for free. But more problematic is the effect on patients’ health. The problem is that no matter how carefully students are supervised, they cannot perform as well as master surgeons. They take longer, do not react as knowledgeably, and make more mistakes. The good teacher knows that the students must be allowed to proceed on their own, and will only intervene when concerned that the ‘mistake’ just made, or about to be made, will have a significant harmful effect on the patient. But sometimes the teacher errs, and the problem becomes a significant one for the patient. Student cases take longer, and the longer the case, the greater the likelihood of infection. The chance of infection is so low, however, that there is no measurable increase in the rate of infection with student-performed surgery. But, and here is the issue, there is an unavoidable conflict of purposes: (a) helping the patient undergoing surgery to obtain as good a result as possible; and (b) helping society by ensuring that there will be well-trained surgeons in the future. The two purposes are not mutually exclusive, but they are not the same. Some training programs are particularly sought out by students because they know that the program gives its students great leeway, that they will make more mistakes, but, in so doing, will learn more than they would in a program where complications are fewer. Without having seen a complication, it is difficult to know how to recognize the premonitory sign that led up to it, and, when recognized, can prevent it from developing. Students are correct in believing that they will end up as better surgeons when they train in a program where they ‘get to do everything’! The moral is: teaching is a proper purpose of performing surgery. However, the primary purpose is the well-being of the patient undergoing surgery. Not to make that clear to the student being taught conveys the message that the patient is ‘teaching material’, thereby devaluing the patient, and vitiating the basic principle underlying great medical care, which is that every person is deserving of the best care that can be provided within the context in which the care is given. 3. A third purpose of surgery is the discovery of new knowledge. Without some attempt to improve
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the cognitive knowledge about what makes patients do well or poorly with surgical procedures, and without an effort to improve the procedural skill involved in the technique of surgery, the methodology of surgery would either not improve at all, or improve only slowly. The consequence of that lack of improvement would be a lack of improvement for both individual patients and society as well. A person who develops a mature cataract becomes incapacitated. Not only can such affected individuals not support themselves, but also they cannot support society. The cost in terms of individual suffering and societal difficulty from cataract has, throughout history, been immense. We have overcome that suffering and difficulty to a large extent by developing surgical methods that remove the cataract with a morbidity of less than 5%, and restore a vision that is far better than that prior to the cataract extraction in well over 90%, and better than the vision before the cataract developed in a significant percentage. Clearly, these methods have made a major contribution to the welfare of individuals and society. What was required for cataract surgery to become so dramatically successful? The answer, of course, is the discovery of new knowledge. Advancement in surgical knowledge has its costs. When Harvey Cushing, at Yale, first started to remove pituitary tumors, the first 30 or so patients on whom he performed the surgery died from it. The first ten patients in whom Harold Ridley inserted the first intraocular lenses did not have a result that was as good as was readily available with intracapsular cataract extraction with use of a contact lens. But the comparison between these two events deserves careful consideration. There was no alternative treatment for a pituitary tumor when Cushing started his surgical exploration. There was a very satisfactory alternative with regard to cataract surgery. The difference is important. While it was incumbent upon Cushing to indicate to his patients that the surgery he was about to perform was unproved and could result in the patient’s death, Ridley had responsibilities that went beyond that. He also needed to tell his patients that a highly successful procedure was already in existence, and that the likelihood was that the patients undergoing his procedure would not do as well as with the older surgery. But the potential benefit of intraocular lenses was enormous, and visionaries such as Ridley realized that. Nevertheless, for such visionaries to act properly, it was essential for them to recall the primary purpose of surgery, which is to help the individual patient. The part of the surgical event that is concerned with the development of new knowledge must be subsidiary to the part that is intended to benefit the patient. When this is not the case, the surgeon is acting improperly. He or she is not performing as a physician, but rather as an investigator. If that surgeon presents himself or herself to
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The purposes of surgery the patient as an investigator who is not primarily interested in the well-being of the patient, if that investigator is aware of the patient’s extreme vulnerability, then in some situations it may be appropriate to proceed with procedures that are primarily for the purpose of developing new knowledge rather than for helping the patient. However, such settings are extremely rare, if they exist at all. Having said that, however, physicians who do not participate in efforts that are designed to provide new knowledge without introducing a significant risk to the patient under consideration are as unethical as those surgeons who present themselves as physicians, and yet are acting as investigators. There is an ethical responsibility to develop new knowledge, so that all patients may benefit in the future. The physician’s primary purpose is to maintain or improve the well-being of his or her patient, but that is not the physician’s sole purpose. To ignore the subsidiary purposes of surgery and the subsidiary responsibility of being a physician is to be at least partially negligent. 4. One of the motivations for performing surgery is the desire to become famous. That is, for some surgeons, performing a surgery becomes driven by a purpose which is, to become famous, to win the Nobel Prize, to be the most revered surgeon, etc. These motivations can lead to great accomplishments. But we have to wonder what percentage of the great contributions in medical history are the result of a desire on the part of the contributor for fame in the sense just described. Nevertheless, as valuable as they are, when those contributions become the primary purpose of the action, they pervert the primary purpose of being a surgeon. In such a situation, the patient is not used as ‘teaching material’ or ‘research material’, but rather as ‘building material’, used by the surgeon to build up a reputation that he or she hopes will bring about his or her desired fame. The consequences of such actions are that developing physicians see such behavior rewarded and therefore may wish to emulate it. Additionally, such behavior rarely results in patients obtaining the best care they can. Again, as with the discovery of new knowledge, such behavior can only be justified if the patient is fully aware of the primary purpose of the surgeon. Performing surgery in order to win an award has its ‘good’ side, as does performing surgery in order to teach and to discover new knowledge. However, that good side does not justify the bad side of it. The good side of it can coexist with the primary purpose, which is to help individual patients, so long as that primary purpose is kept primary in the surgeon’s mind. 5. Surgery is challenging. It is exciting. For some surgeons, it is fun. It can become a game. And the seductive power of games has become especially apparent with the development of computer games.
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13 Millions of individuals across the world now spend millions of hours playing games that will bring them no reward other than the fun involved in the playing of the game itself. The value of having fun is great. Play is one of the wonderful aspects of existence, whether that play be squirrels cumbering on the lawn, seagulls soaring, or computer nerds figuring out the details of a new game. Fun is conducive to health. One of the characteristics of Mozart’s music that makes it one of the towering accomplishments of mankind is the great delight that the music evokes in listeners. It is appropriate for surgeons to enjoy themselves while performing surgery. It is not appropriate for surgeons to enjoy themselves at the expense of those on whom they are performing that surgery. If the fun part of surgery results in the diminishment of the surgeon’s ability to accomplish the primary goal, i.e., helping individual patients, then the fun part is inappropriate. However, as with the other purposes discussed, this need not be the case. It is a matter of remaining clear in one’s mind and actions as to what is a primary and what is a subsidiary purpose. 6. For some individuals, surgery becomes a way to amass wealth. It is easy to say that the desire to become hugely rich is less worthy than the desire to ‘make a living’. However, this is probably not a defensible position. Becoming vastly wealthy allows the vastly wealthy individual to do things that are not possible for an individual who has less means at his or her disposal. A significant portion of the great contributions to human society has been made by individuals who amassed great wealth. Universities have been founded, churches built, centers for the homeless established, great works of art commissioned, students supported, and on and on, by wealthy people. Creation of wealth is one of the essential components of a successful society. There is, then, nothing inherently wrong in wanting to become rich by performing surgery, although the initial response of most to such a purpose will be that, becoming rich at the expense of the sick, is wrong to the point of being obscene. But this is not so, as long as the primary purpose continues to be primary. If a surgeon is able to provide care which is superb and to make a great deal of money at the same time, that surgeon should be congratulated and encouraged by the medical profession and society at large. 7. The final purpose for performing surgery includes a miscellany of purposes. For example, a surgeon may want to discover a better way to perform cataract extraction, because his parents became blind from cataract surgery. Or a surgeon may want to improve the technique of treating compound fractures of the femur because his brother was incapacitated by such a fracture. These purposes, too, are appropriate, because they may motivate, encourage, and sustain surgeons through
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difficult times. However, when these purposes become all-consuming, so that what the surgeon does is done at the cost of individual patients, then the purpose is improper. But, as with all the other purposes discussed in this chapter, they are inappropriate only to the extent that they interfere with the effort of the surgeon to provide the best care possible for the patients under his or her care within the context in which he or she finds him- or herself.
Conclusions Medical ethics is an essential part of medical actions, and is a subgroup of bioethics, which is de-
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rived from applied ethics and philosophical ethics. Applied ethics is based on the principle that each allows the other the same amount of freedom that he requests for himself. When this is not the case, but when a person acts according to the principle, “most for me, what’s left over for the others”, justice has been violated (or: a problem regarding justice arises). A number of purposes are included in the aims of surgery, i.e., surgery is undertaken: (1) for the purpose of repairing; (2) to teach others how to perform surgery; (3) to discover new knowledge; (4) to allow the surgeon to become famous; (5) to have fun; (6) for the purpose of becoming rich; and (7) to achieve a specific goal, perhaps not directly related to improving the health of the patient.
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Contact lenses for ophthalmic laser treatment
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Contact lenses for ophthalmic laser treatment Einar Stefánsson and Franz Fankhauser University of Iceland, Reykjavik, Iceland
Keywords: photocoagulation, laser contact lens, positive/negative lens, complications of laser therapy, parfocal systems, defocus system, Goldmann lens, aphakic eye, phakic eye
Abstract Contact lenses are frequently used for the diagnostic examination and laser treatment of the fundus and anterior segment of the eye. The lenses differ in magnification, laser spot size, and field of view. This chapter deals with the optical characteristics of various contact lenses, and how they apply to laser treatment.
Introduction This chapter presents a short survey of a number of optical systems used in laser therapy in ophthalmology. It is restricted to the most common optical systems used today, and is predominantly oriented toward the application of the laser at the posterior segment. When visible or near infrared laser light is used for treatment of the fundus of the eye, the light must pass through the transparent ocular media, to reach the fundus of the eye. The ocular media must be considered as part of the optical system, together with the laser instrument and its delivery optics. A thorough review of the optics of ocular laser application has been written by Fankhauser et al., and this chapter relies substantially on that work.1 Contact lenses Contact lenses are commonly used to direct laser energy to various sites in the eye. They were introduced in 1967 by Fankhauser and Lotmar as coupling elements for photocoagulation of the retina.2,3 The use of the contact lens for this purpose has now become universally accepted. The more recent clinical use of low-power pulsed lasers for photocoagula-
tion in the human eye places even greater demands upon the optical performance of contact lenses as auxiliary coupling elements. Some general and well-known advantages of using contact lenses for laser treatment, as well as for examination, may be briefly summarized as follows: the cornea is kept moist, the eye is stabilized, and the lid is kept out of the way. Although handling a contact lens implies additional effort for the physician, given the considerable advantages, this minor inconvenience is worthwhile. For a more thorough understanding of the advantages of using certain contact lenses, it may be helpful to review some basic principles. The laser treatments considered in this chapter will be divided into ‘photocoagulation’ and ‘disruption’. Photocoagulation Photocoagulation, which followed the pioneering work of Meyer-Schwickerath and the introduction of the laser, is based in the first approximation upon the ‘linear’ absorption of light for at least as long as the evaporation of water is not reached. The absorbed light energy is converted into thermal energy, which leads to denaturing, coagulation, evaporation and/or carbonization of the tissue, depending upon the interaction time and the amount of energy deposited. All these effects can be seen as ‘thermal effects’, and the laser working in the thermal domain may be defined as a low power laser, although the pulse energy may be high. Typical tissues treated in this way include retina/choroid, trabecular meshwork, iris, ciliary body, and blood vessels. The most common lasers used are the continuous wave (cw) models, such as the argon ion, krypton, dye, diode, and Nd:YAG.
Address for correspondence: E. Stefánsson, MD, PhD, The National University Hospital of Iceland, Department of Ophthalmology, Reykjavik, Iceland. e-mail:
[email protected] Lasers in Ophthalmology – Basic, Diagnostic and Surgical Aspects, pp. 15–25 edited by F. Fankhauser and S. Kwasniewska © 2003 Kugler Publications, The Hague, The Netherlands
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The demands made on low-power lasers and on their handling by the physician are less than those involved in the application of high-power lasers. A focal spot diameter of less than 50 µm is not required, whereas with high-power lasers, the focal spot diameter is often reduced to a tenth of this value. Good reproducibility of the laser parameters is important in order to avoid excessive and inadequate effects. In addition to stable laser emission, the focal diameter on the target tissue should be kept constant once the laser parameters have been set for the desired therapeutic effect. Photodisruption The application of g-switched lasers, which usually work at high power levels, is based on an interaction mechanism between light and tissue, which is completely different from the thermal effects mentioned above. Instead, very high laser intensities are used to cause optical breakdown in the media through which the laser ray passes. The intense shock waves produced in the region of the breakdown cause the cutting or disruptive effects. In order to obtain the high intensities required for this nonlinear effect, the laser energy must be delivered in a very short time interval, i.e., within the ns or ps range (10-9-10-12 seconds), and in a very small volume, which requires a small focal spot. The confinement of laser energy to a small focal spot is another feature that distinguishes the highpower from the low-power laser. As a consequence of the Lagrange invariant, the size of the focal spot of a laser beam is inversely proportional to the beam convergence anterior and divergence posterior to the focal spot (i.e., the cone angle of the beam is greater for smaller focal spots, and small focal spots are required to get therapeutic effects at small pulse energies). The diameter Df of the focal spot of a fundamental (TEM00) Gaussian laser beam is in good approximation proportional to the wavelength λo in vacuum in the medium, and inversely proportional to the refractive index of the medium n as well as to its cone angle α, i.e., the full angle of the laser beam:
Contact lenses: principles and safety The power and energy density of the laser at its focus are proportional to the inverse size of the laser spot. This is true in photocoagulation as well as in photodis-ruption. The high intensities required to achieve optical breakdown depend upon the cone angle of the beam, and this influences the cutting efficiency, the volume of the focus, and the energy deposited in the media in front of and behind the focus. The cone angle of both the observation and the aiming beams is equally important. Increasing the cone angle of the observation beam increases the magnification of the observed image and the optical resolution, although it also results in a reduced depth of focus and therefore enhanced focusing accuracy. According to the above, two interdependent quantities, namely the size of the focal spot and the cone angle of the laser beam, must be considered when evaluating contact lens performance. The former is responsible for the cutting efficiency and dictates the volume of the tissue affected by laser irradiation, and the latter is responsible for the amount of tissue damaged beyond the focal spot and, hence, for the safety of the pre- and postfocal media and structures. Therefore, the reduction in pulse energy and the increase in cone angle of the beam act in the same way with regard to cutting efficiency. When working with a photodisruptive laser, intensities equal to or greater than a certain threshold must be generated in order to achieve optical breakdown. This threshold intensity is a function of the optical and chemical properties of the medium in which the breakdown is generated. For instance, the threshold intensity for bulk water has been measured to be about 1012 W/cm2.8 The focal spot should be kept as small as is necessary to safely meet the energy density requirement of the microsurgical task at hand. By so doing, the intensity and energy density in the prefocal media are minimized, and safety is maximized. The same applies to the postfocal media if we assume that the laws of geometric optics hold for laser light passing through the disturbed media after optical breakdown has been induced.
Df = 4λ0 Df
π n0α
This means that, by focusing a laser beam from air into a medium with a refractive index greater than one, even without changing the cone angle α, the focal spot diameter is reduced because the refractive index is enlarged. A further reduction of the spot diameter can be obtained by enlarging the cone angle. (The increase in the cone angle, however, is limited mainly by optical aberrations, which tend to increase with increasing cone angle.) This problem can be solved by using appropriately designed optical components such as contact lenses,4-6 and by aiming the laser beam correctly.1,7
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Optics of the eye with a contact lens The imaging properties of optical systems such as the eye with a contact lens satisfy the paraxial Gaussian approximation. This means that the rays of an incident light bundle are all focused on the same point, as long as they are relatively near the optical axis. It the rays are not close to the optical axis, they are not longer focused on one single point, and the resulting aberrations, known as optical aberrations, increase with increasing cone angle (numerical aperture). However, there are two exceptions, and it is upon these exceptions that all contact lenses for photodisruptive purposes should be based. The first
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Contact lenses for ophthalmic laser treatment exception applies to light bundles which are focused towards the center of curvature of the front refracting surface of the contact lens system. Here, all the rays are normal to the optical surface, and consequently they are not refracted. This means that, by focusing a beam towards the center of curvature of a surface, no spherical aberration occurs. This also means that the cone angle of the incident beam remains constant and is not changed by a spherical surface. There are other aberration free points: a refracting surface has two distinct conjugated points, A and A’, which are completely free from any optical aberration. Therefore, all the rays from a light bundle focused towards the point A are perfectly focused into point A’ after passing through the surface, regardless of their angle to the optical axis. These points are known as aplanatic points, and are defined by the Young-Weierstrass theorem.9 By rotating a converging beam around the center of rotation of a refracting spherical surface, aplanatic surfaces are realized. A beam imaged on such a surface suffers no spherical aberration. One such surface is displayed in Figure 1. For each photodisruptive task (iridotomy, capsulotomy, dissection of vitreous membranes), a specific working distance (or focal length) is required, and appropriate contact lenses must be used in order not to compromise safety.5-7,10
17
Fig. 1. Definition of aplanatic surfaces. These are realized by rotating the converging beam around C. There are two aplanatic surfaces, S and S’, where S’ is the image of S. A is the focal spot in air, A’ the focal spot in a medium with index n’. All points of S’ are aplanatic. (Reproduced from Born and Wolf 9 by courtesy of the publisher.)
Contact lenses for photocoagulation Laser light for photocoagulation may be applied anywhere in the fundus, including off-axis regions, and not necessarily near the paraxial space. Contact
Table 1. Compilation of a number of representative fundus contact lenses (see Dewey32) Lens
Magnification
Mainster high magnification Volk 60 D Volk centralis direct NF Volk centralis direct Volk area centralis Mainster standard Ocular high magnification (=78 D) Goldmann three mirror (direct) Volk 78 D Volk 90 D Osher pan-fundus Ocular standard (=90 D) Panfundoscope (Rodenstock) Volk transequator Mainster wide field Kreiger (direct) Volk Quadraspheric Haag-Streit 901 9 fundus Haag-Streit 903 three mirror Haag-Streit 908 3+1 mirror Haag-Streit 630 three mirror
Distance to
fundus image
laser spot
lens front
retina
lens focal length (mm)
field of view (full angle)
1.25 1.14 1.12 1.11 0.96 0.95 0.93 0.93 0.92 0.76 0.76 0.75 0.71 0.69 0.67 0.65 0.50 0.96 0.91 0.91 0.91
0.80 0.88 0.89 0.90 1.04 1.05 1.07 1.08 1.09 1.32 1.32 1.34 1.41 1.44 1.50 1.53 2.01 1.04 1.1 1.1 1.1
23.35 14.04 -20.30 -20.21 13.85 14.89 10.43 -32.52 0.77 9.40 9.80 9.52 -9.42 10.81 6.37 -11.96 6.05 -17.35 -30.3 -30.3 -24.95
64.55 66.14 6.00 6.09 50.82 67.34 57.13 17.56 57.22 50.20 51.50 49.72 45.47 53.65 53.57 14.58 43.03 8.35 16.1 16.1 13.5
22.41 19.07 -18.57 -18.29 16.58 16.46 15.68 -15.03 15.38 12.66 12.70 12.55 11.81 11.79 11.33 -10.89 8.48 -15.9 -15.05 -15.05 -15.05
68° 67° 31° 31° 90° 90° 84° 36° 73° 69° 98° 94° 120° 110° 125° 42° 130° 37° 37° 37° 37°
-14.25
37°
-14.85 -14.25 -15.05
37° 37° 37° 64° 64° 65°
Haag-Streit 906 three mirror (pediatric) Haag-Streit 907 three mirror (children) Haag-Streit 1110 laser fundus 0.9 Haag-Streit 1140 laser two mirror 0.86 Meridian/Haag-Streit CGR 3 0.75 Laser fundus 0.75 HS 900 (indirect) 0.72
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1.1 1.1 1.4 1.4 1.38
Depending on the size of the eye -16.2 9.5 -31.15 18.25
12.3
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lenses for coagulation tasks (Table 1) are not restricted to the Gaussian (aberration-free) space, nor to an aplanatic point in the eye. On the contrary, they may sometimes be directed towards eccentric sites in the eye. The price to be paid may be an aberration-dependent enlargement of the focal spot, which often results in a considerable increase in the laser energy required for the therapeutic effects, according to the square of the focal spot size. Some such lenses are realized by internal mirrors (such as the classical Goldmann three-mirror lens) for the irradiation of peripheral retinal areas. The characteristic feature of mirror contact lenses is their off-axis, often quite oblique beam path produced in the eye. This oblique path results in aberrations – mainly at the anterior surfaces of the cornea and at the crystalline lens or implant – such as coma and astigmatism, in addition to the previously-mentioned spherical aberration. This has been shown to be an important performance-degrading factor for laser irradiation of the peripheral retina. For a more intimate understanding of the effect of image degradation on both observation and laser irradiation, the reader is referred to the data derived from schematic eye modelling.11-14 Aberrations in the periphery of the fundus oculi increase with visual angle. They impair both observation and coagulation efficiency. These difficulties can be overcome, to a large extent, by the use of a plano-concave contact lens.11,15 Other attempts to combat peripheral image degradation have not been very successful.16,17 Another problem that arises in biomicroscopy of the eye is the determination of the absolute dimensions of objects and structures at the periphery of the retina. Little is known about the effective focal length of the eye in oblique view, and therefore about its magnification, as a function of the visual angle. Auxiliary lenses Illuminated emmetropic eyes emit parallel rays from the fundus. Therefore, the fundus cannot be exam-
Fig. 2. Ray path and symbols as used in the eye-contact lens system with mirrors: α: visual angle without contact lens; γ: angle of acceptance at the retina; dashed line: angle of incidence of the ray with the contact lens. (Reproduced from Fankhauser and Lotmar45 by courtesy of the publisher.)
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ined with a microscope which has only a short object distance. In order to use a microscope to view the fundus, an intermediate image of the fundus must be made at the objective plane. This can be achieved using either a concave or a convex auxiliary lens. Concave (negative) lenses create an upright, virtual image of the fundus or vitreous, whereas convex (positive) lenses create an inverted, real image. Prototype concave lenses include the Goldmann contact lens18,19 and the Hruby preset lens,20 and there are a large number of convex preset and contact lenses (Table 1). No essential difference exists between indirect ophthalmoscopy and slit-lamp biomicroscopy using a preset convex lens, except that, in indirect ophthalmoscopy, a real, inverted image in air is inspected by the naked eye instead of having to use a microscope. Figure 3 shows the intermediate image created by some commonly used contact lenses.21 The Goldmann three-mirror contact lens, which is used as a universal lens, is a contact lens with excellent optics. Its field of view is limited to about 30° (full angle) in each direction of observation. Its outstanding optics make it ideal for examination of the vitreous, and the fact that it is a contact lens makes for a somewhat wider field of view. Finally, magnification of the fundus image with the Goldmann lens is not so dependent on the refractive power of the patient’s eye as are other negative lenses. The images seen via the mirrors are laterally reversed, i.e., they are mirror images compared to direct viewing. All negative lenses suffer from the fact that the greater the patient’s myopia, the shorter the working distance, although in practice this is rarely a serious problem. For examination of a highly myopic eye in which a larger working distance is required, positive lens systems are preferred.22 Since positive lens systems tend to magnify the entrance pupil of the eye under examination, the iris in first approximation, in contrast to negative lenses, no longer acts as a field stop; a favorable field size is achieved which can even enable examination of the ora serrata. Curvature of field, once considered a problem in early positive lenses, such as the El Bayadi lens22 or the panfundoscope,23 is corrected in modern wide-angle lenses, although peripheral image degradation obviously remains a problem. Illumination of the peripheral fundus is virtually impossible with most slit lamps, and fiberoptic transillumination has been suggested. A greater working distance than with negative lenses is required when positive lens systems are used, and this may be beyond the working range of some slit lamps. Examination and laser treatment of the peripheral retina Observation of the lateral periphery can be difficult; the reduced effective aperture of the pupil in oblique
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a.
Fig. 4. Images of the microscopic objectives and illuminating slit together with the laser beam in the pupil of the eye to be examined and irradiated, with negative (top) and positive (bottom) contact lenses. (Reproduced from Fankhauser et al.1 by courtesy of the publisher.)
b.
c.
d. Fig. 3. a. The Goldmann lens has a flat anterior surface and produces an erect, virtual ophthalmoscopic image located near the posterior surface of the crystalline lens. b. The Krieger lens has a concave anterior surface and produces an erect, virtual ophthalmoscopic image located in the anterior vitreous humor. c. The Panfundoscope lens has a biconvex, spherical anterior lens element, and produces an inverted, real image inside the biconvex lens. d. The Mainster lens has a biconvex, aspherical anterior lens element, and produces and inverted, real image anterior to the biconvex lens. (Reproduced from Mainster et al.21 by courtesy of the publisher.)
projection makes it difficult to assemble both objectives, the illumination and the laser beam, into the narrow space available. Figure 4 clearly shows the advantages when using positive lenses or lens systems.1 However, it should not be overlooked that narrowing of the beam associated with positive lenses results in an increase of intensity and radiant exposure in prefocal regions, proportional to the square of the reduction in beam diameter. This should be taken into consideration, particularly when photocoagulation is performed in the presence of opaque media, which have enhanced absorption, since otherwise damage to the cornea, iris and crystalline lens
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could result.24 In particular, vignetting of the laser beam by the iris, when using positive optics, may have serious consequences, whereas damage when using negative lenses, due to the larger beam diameter at the pupil and therefore to reduced intensity and fluence, is negligible. Similar problems have been observed in laser irradiation of the fundus by means of indirect ophthalmoloscopy,25-30 which is identical to laser irradiation with positive contact lenses, except that in the former a non-contact method is used and observation is performed with the naked eye. Contact lens-assisted irradiation using positive lenses is much safer than when using non-contact irradiation such as that in indirect ophthalmoscopy, because with rigid slit-lamp delivery systems, movements of the surgeon’s and/or the patient’s head will only slightly affect aiming accuracy. In the best circumstances, using indirect ophthalmoscopy-related retinal irradiation, targeting has been estimated to be approximately ±200 µm,26 although under unstable conditions, this may be significantly greater. Laser treatment using a contact lens While generally helpful, the contact lens-eye optical system used for imaging the laser beam may also be a possible threat to safety, if mismanipulations or improper calibration lead to inadvertent exposure of the cornea, iris, and crystalline lens to high-intensity laser radiation. As shown in Figure 5, with a wide-angle positive lens (Volk Quadraspheric for the same spot setting), the size of the retinal spot is twice as great as with the Goldmann lens and Volk 78 D lenses. As a consequence, for the same threshold effect, four times as much power is required with the Volk Quadraspheric lens as with the Goldmann and the Volk 78 D lenses. Since the beam diameter at the cristalline lens is about twice as small, the energy density and power density are 16 times as much with Volk Qua-
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spot setting on the slit lamp (microns) beam diameter in air at the observation plane
Fig. 5. Laser spot diameter on the retina (mm) as a function of the laser spot diameter on the observation plane in air (µm), as indicated on the scale. The laser spot diameter is twice as large with the Volk Quadraspheric contact lens as with the Goldmann three-minor and the Volk 78 D contact lenses. When working with the Volk Quadraspheric lens, four times as much power is necessary. Because the beam waist at the crystalline lens is about twice as small as with the Goldmann and Volk 78 D lenses, the intensity of the beam at the crystalline lens is 16 times as much with the Volk Quadraspheric lens. (Reproduced from Fankhauser et al.1 by courtesy of the publisher.)
draspheric lens compared to the other two auxiliary lenses (Fig. 5). Two optical systems for controlling laser spot size are currently in use, namely, parfocal (generic parfocal) and defocusing (generic defocus) systems, which differ in their basic optical principles.1,31 Parfocal systems With the parfocal system, the core of the laser delivery fiber is imaged on to the slit-lamp focal plane
at varying magnifications. This produces a clean, round image with relatively homogenous laser intensity over the entire spot. The parfocal approach changes the beam divergence as the laser spot size is changed, using a zoom, and this results in a large depth of field, i.e., a long cylindrical beam waist at large spot sizes (Fig. 6). The large depth of focus associated with large laser focal spot sizes ensures that the correct focal spot size is delivered to the retina, even if small aiming errors occur. As the focal spot size at the retina is increased, the beam diameter at the cornea, iris and crystalline lens decreases. With all positive fundus lenses, the situation can arise where the beam diameter is larger at the retina than it is at the pupil of the eye. Indeed, for each combination of laser and lens, a critical point can be found where the corneal or lenticular beam diameter becomes smaller than the retinal beam diameter. This cross-over point has been suggested to be the limit of the largest retinal spot size that is recommended for use with particular combinations of fundus lenses and laser delivery systems. In all laser-lens combinations in which the corneal or lenticular beam diameter becomes smaller than the retinal beam diameter, the fluence at the cross-over point is about 2 J/mm2.32 Therefore, the limit of 2 J/mm2 has been used for the calculation of safe beam diameter limits (Table 2). This limit only applies to transparent media; if opacities, which absorb energy more strongly, are present in the eye, lower limits are required if safety is not to be compromised. The critical point must be computed or must be indicated by the manufacturer of both the auxiliary lens and the biomicroscope.
Fig. 6. Two imaging systems are shown. Left: the generic parfocal system. In this configuration, the focus location does not change. The focal spot diameter (∅ is increased by reducing the cone angle α. Advantage: sharp image of fiber and ‘top hat’ intensity profile at the focus. Disadvantages: (1) to obtain a spot size of as small as 50 µm, a large cone angle is necessary, as a 50-µm diameter optical fiber was used for beam delivery. In most commercial instruments, the laser beam is clipped at the slit lamp, resulting in a lower maximum power available to the physician; (2) a true parfocal system with spot diameters of up to 1 mm can be dangerous because the reduction of the cone angle results in a small beam diameter at all structures through which the beam passes. Right: generic defocus. In this configuration, the focus location is moved, but the cone angle is not changed. The focal spot diameter ∅ is constant and the beam diameter at the target (T) is adjusted by the amount of defocus. Advantage: no increased power density anywhere in the laser beam. Disadvantages: (1) intensity profile at the target has a Gaussian form which is less suitable for treatment of discrete retinal areas; (2) with a true defocus system, the laser beam diameter continues to converge to a 50-µm focus beyond the target. Hence, if the medium is slightly transparent beyond the target the power density behind the target may be much higher than on the target, thus endangering the medium. (Reproduced from Fankhauser et al.1 by courtesy of the publisher.)
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Table 2. Spot diameter for a number of contact lenses at which the corneal energy density equals 2 J/mm2. The magnification (%) which is required to focus the laser beam in order to obtain the desired spot size at the retina is shown, i.e., the laser spot size without contact lens. For example, the Volk Quadraspheric lens with a beam magnification of 201% gives a maximum usable spot setting of 200 µm. System Lens
M
Volk 60 D indirect actual spot size (µm) Volk area contralis actual spot size (µm) Mainster actual spot size (µm) Goldmann three-mirror actual spot size (µm) Volk 78 D indirect actual spot size (µm) Volk 90 D indirect actual spot size (µm) Panfundoscope actual spot size (µm) Krieger actual spot size (µm) Volk Quadraspheric actual spot size (µm)
88.0% 104.0% 105.0% 108.0% 109.0% 132.0% 141.0% 153.0% 201.0%
VariSpot P
°D
450 396 400 416 370 389 500 540 350 382 300 396 270 381 350 536 200 402
1000 880 570 593 1000 1050 1000 1080 940 1025 570 752 500 705 1000 1530 200 402
920
LDS-20
Parfocal
Keeler
1000 1369 540 700 880 1400 1000 1680 820 1300 580 1000 520 900 1000 2380 200 533
1000 937 760 900 1000 1118 1000 11150 1000 1161 800 1170 700 1120 1000 1629 260 700
450 396 400 416 370 389 1000 1080 350 382 300 396 270 381 350 536 200 402
500 440 500 520 500 525 500 540 500 545 500 660 500 705 500 765 430 864
M: magnification; P and D: two different realizations by Coherent Medical; 920, LDS-20, Parfocal, Keeler: system names (see Dewey32)
Defocus systems G
In the defocus system, a smaller laser spot (typically 50-200 γm) is shifted to a point beyond the slit-lamp focal plane (Fig. 6). This avoids the major drawback of the parfocal systems, in which the beam diameter at the pupil of the eye is relatively constant. The price paid for the larger beam diameter at the pupil with the defocus system is that the nice sharp image of the fiber core, typical of parfocal systems, gradually fades from high power to low power at the edge (Fig. 7). The Goldmann contact lens/Varispot32 combination (parfocal/defocus) may be regarded as the one providing maximum safety for the preretinal media (Fig. 8). Here, the retinal irradiance and fluence will always be greater than that in the preretinal media. The Goldmann contact lens/parfocal combination provides the maximum safety and quality performance for laser spot sizes up to about 300 µm, and the Goldmann lens/defocus combination matches this performance even with larger spot sizes (Fig. 8). In contrast, when using the combination Volk Quadraspheric lens/VariSpot projection system (or any other lens of the same power as the Volk Quadraspheric lens), the energy density and power density with a focal spot size of about 30 µm are almost the same at the retina as at the crystalline lens, when utilizing the parfocal mode (Figs. 8 and 9). With larger retinal spot sizes, both in the defocus and in the parfocal regime, the energy density in the pre-
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Fig. 7. Schematic parfocal (generic parfocal) and defocus (generic defocus) beam profiles with a 500-µm diameter spot size. Perfect ‘top hat’ and Gaussian profiles are plotted for reference. (Reproduced from Fankhauser et al.1 by courtesy of the publisher.)
retinal media far exceeds the energy density at the target (Fig. 8). In most modern laser projection systems, mixed parfocal/defocus systems are used, which combine the advantages of both principles. Changes in refraction by silicone oil or gas in vitreous surgery As stated at the beginning of this chapter, the contact lens and the eye should be seen as one optical system. Changing the optical characteristics of the
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Fig. 8. Beam diameter at the crystalline lens as a function of the selected spot setting for Goldmann three-mirror, Volk 78 D, and Volk Quadraspheric contact lenses, imaged in the generic defocus mode. Spot setting: beam diameter in air at the observation plane. See text. (Reproduced from Fankhauser et al.1 by courtesy of the publisher.)
spot setting on the slit lamp (microns) beam diameter in air at the observation plane
Fig. 9. Beam diameter at crystalline lens as a function of the selected spot setting for Goldmann three-mirror, Volk 78 D, and Volk Quadraspheric contact lenses, imaged in the generic parfocal mode. See text. (Reproduced from Fankhauser et al.1 by courtesy of the publisher.)
ocular media will affect the optics of the eye. Filling the vitreous cavity with liquid silicone, air or gases, as is frequently done in vitreous surgery, has a profound effect on the total refractive power of the eye, on the visibility of the fundus, and on the refractive state of the eye during and following surgery.33-36 More seriously, targeting of the laser may be affected by the modified optical parameters of the eye, and this may alter both the diameter of the laser focus spot and its intensity. Stefánsson and Tiedeman examined this optical situation, and discussed fundus image location and magnification using water, air, or silicone oil as vitreous substitutes in phakic and aphakic eyes, and for various auxiliary contact lenses.37 They also computed the fundus image location and magnification in the eye when using both planoconvex posterior and anterior chamber lens implants (the former positioned 5 mm and the latter 3 mm behind the front surface of the cornea) and water, air or silicone oil in the eye. In addition, they determined when the fundus was visible with the above combinations using
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either the operating microscope or slit-lamp biomicroscope. They stated that exact determination of the location of the fundus image requires complex optical calculations, although it is possible to simplify this issue by concentrating on a few principles. When air or silicone oil is placed in the vitreous cavity, the refractive power of the posterior surface of the crystalline lens is changed. This surface is normally a low-power positive lens, as the refractive index of the lens cortex, 1.386, is higher than the refractive index of the vitreous gel, 1.336. As the refractive index of air is only 1.000, when the cavity is filled with air, the power of the posterior crystalline lens surface is dramatically increased. A highpower negative lens, such as the biconcave corneal lens, is then needed to counteract the effects of this high-power positive lens (Fig. 10). While the refractive index of vitreous gel is lower than that of the crystalline lens, the refractive index of silicone oil is slightly higher. Thus, when the vitreous cavity is filled with silicone oil, the posterior surface of the crystalline lens changes from a low-power positive to a low-power negative lens. While this change is too small to be a problem during vitreous surgery, it must be corrected postoperatively with spectacle glasses for hyperopia. If the aphakic eye is filled with silicone oil, the posterior surface of the cornea changes from a lowpower negative lens to a low-power positive lens, as the refracting index of silicone oil is higher than that of the cornea. This conversion of the cornea to a positive lens reduces the hyperopic correction that the aphakic eye would otherwise need. The convex silicone ball in the aphakic eye has a positive refractive power and, therefore, induces a myopic shift, whereas the concave silicone ball in the phakic eye induces a negative lens and, therefore, makes the phakic eye slightly hyperopic (Fig. 11). In the aphakic eye, the posterior corneal surface is important. When the aphakic eye is filled with air, the posterior surface of the cornea changes from being a low-power to a high-power negative lens that neutralizes the refractive power of the cornea. The latter is reduced to nearly zero, and thus the fundus can be visualized without any additional optical aids (Fig. 12). In the pseudophakic eye, where the posterior surface of the intraocular lens is flat, there is no change in the refractive power at this surface when vitreous substitution is performed, since a flat surface has no refractive power, regardless of the refractive index of the substances on either side. This has important implications in vitreous surgery, in that it is no longer necessary to use a biconcave corneal lens in a pseudophakic eye which has an air-filled vitreous cavity. It is possible to observe and treat the fundus in these eyes using a flat-faced contact lens or a prism lens. The complexity of optical changes with vitreous substitution is huge, and solutions in specific cases
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Fig. 10. Schematic drawing of a phakic eye showing an object arrow on the fundus and an image arrow located appropriately. Each row of figures corresponds to a particular intraocular optical situation, i.e., normal eye, air-filled vitreous cavity, air-filled eye, airfilled anterior chamber, and silicone oil filled vitreous cavity, as indicated. The columns correspond to the corneal contact lens used, i.e., no lens, flat or planoconcave lens, biconcave lens with anterior surface powers -63 D, and biconcave lens with -93 D anterior surface power. Each situation, where the fundus can be visualized through the operating microscope or slit-lamp microscope, is marked with a black dot. (Reproduced from Stefánsson and Tiedeman37 by courtesy of the publisher.)
may not always be possible, as mentioned by Launay et al.38 and Docchio et al.39-41 In in vitro experiments, Launay et al. determined the variation in shape of a silicone ball as a function of the volume of fluid injected into the vitreous cavity in emmetropic, phakic or aphakic eyes for different positions of the head. Obviously, the volume, shape, and location of the injected silicone will all strongly and unpredictably influence the optical properties of the eye. A very ambitious approach was undertaken by Docchio et al.39-41 and Azzolini et al.42 These authors used a ray-tracing model to investigate the refractive properties of the interfaces between different ocular media and vitreous substitutes, with regard to transpupillary laser beam delivery during photocoagulative procedures. The study outlined the role of these interfaces in focusing or defocusing the laser
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beam along its path within the eye. This effect is dependent upon the angle of incidence, the number of interfaces, and the change in refractive index across each interface. These studies revealed the main problems to be inadequate power density or nonuniformity of the laser spot, which in turn resulted in non-optimal photocoagulation. Transpupillary photocoagulation through silicone oil demonstrated improved performance compared to perfluoro-noctane, due to the more favorable sequence of refractive indices encountered by the laser beam. These study highlighted the fact that, in many situations, it is difficult, if not impossible, to correctly focus the beam on the target site, due to the vignetting effect of the iris.39-42 In another approach, Azzolini et al.43 reported the effects of vitreous substitutes in endoocular laser photocoagulation, and they also described
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Fig. 11. The refractive correction (spectacle) in diopters in eight aphakic eyes before silicone oil filling, as well as with silicone oil and after removal of intraocular silicone oil. The three measurements in each eye are connected by a line. The two eyes that did not return to the previous degree of hyperopia after removal of silicone oil underwent penetrating keratoplasty at the time of silicone oil removal (dotted lines). (Reproduced from Stefánsson et al.36 by courtesy of the publisher.)
the effect of the air bubbles in the laser beam in endoocular pathology. So far, our insight into substitution in surgery of the anterior and posterior segments of the eye is almost exclusively based on ray-tracing models and in vitro experiments. Laser irradiation of the silicone oil-filled eye may involve some risk. Huy et al.44 have shown that the in vitro exposure of silicone oil to radiation from Nd:YAG lasers results in the formation of transient breakdown gases, which are mainly composed of methane, ethylene, and traces of ethane, as identified by head-space gas chromatography. However, no clinically significant damage has been reported to result from laser irradiation due to the presence of silicone oil or gas in the eye. Moreover, Azzolini et al.43 performed retinal endophotocoagulation through perfluorodecalin in rabbits after vitrectomy. These studies indicated that no extra care is necessary when endocoagulation is performed through perfluorodecalin, provided circular spots are used and the energy is delivered accurately to the target site.
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Fig. 12. Schematic drawing of an aphakic eye. The normal, air-filled, and silicone oil-filled eyes are shown with no contact lens and a flat (planoconcave) corneal contact lens. An object arrow is drawn on the fundus, as well as the image arrow corresponding to it. Each situation where the fundus can be visualized with the slit lamp or operating microscope is marked with a black dot. (Reproduced from Stefánsson and Tiedeman37 by courtesy of the publisher.)
Conclusions Examination and laser treatment of the eye is aided by the use of contact lenses. A variety of contact lenses is available and, in conjunction with the eye, they form a complex optical system. The laser surgeon must understand the fundamentals of the various optical systems that he or she utilizes, in order to maximize the usefulness and minimize the risks involved in the laser treatment. Several aspects of laser irradiation of the eye have been described which must be considered when setting safety limits. The individual laser surgeon must ascertain whether the safety requirements of each and every therapeutic laser application have been satisfied.
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Contact lenses for ophthalmic laser treatment References 1. Fankhauser F, Durr U, Giger H, Rol P, Kwasniewska S: Lasers, optical systems and safety in ophthalmology: a review. Graefe’s Arch Clin Exp Ophthalmol 234:473-487, 1996 2. Fankhauser F, Lotmar W: Photocoagulation through the Goldmann contact glass. Arch Ophthalmol 77:320-330, 1967 3. Fankhauser F, Lotmar W: Methods of photocoagulation through the Goldmann contact glass. Mod Probl Ophthalmol 7:256-272, 1968 4. Riquin D, Fankhauser F, Lörtscher HP: Contact glasses for use with high power lasers. Int Ophthalmol 6:191-200, 1983 5. Rol P, Fankhauser F, Kwasniewska S: A new contact lens for posterior vitreous photodisruption. Invest Ophthalmol Vis Sci 27:946-950, 1986 6. Roussel P, Fankhauser F: Contact glass for use with high power lasers: geometrical and optical aspects. Int Ophthalmol 6:183-190, 1983 7. Rol P, Fankhauser F, Kwasniewska S: Aiming accuracy in ophthalmic laser microsurgery. Ophthalmic Surg 17:278282, 1986 8. Taboada J: Interaction of short laser pulses with ocular tissues. In: Trokel SL (ed) YAG Laser Ophthalmic Surgery, p 17. Norwalk, CN: Appleton Century Croft 1983 9. Born M, Wolf E: Stigmatic imaging of surfaces. In: Principles of Optics, 4th edn, pp 149-150. Oxford: Pergamon Press 1970 10. Rol P, Fankhauser F, Kwasniewska S: Evaluation of contact lenses for laser therapy. I. Lasers Ophthalmol 1:1-20, 1986 11. Lotmar W: Theoretical eye model with aspherics. J Opt Soc Am 61:1522-1529, 1971 12. Nakao S, Fujimoto S, Nagata R, Iwata K: Model of refractive-index distribution in the rabbit crystalline lens. J Opt Soc Am 58:1125-1130, 1968 13. Navarro R, Santamaria J, Bescos J: Accommodation-dependent model of the human eye with aspherics. J Opt Soc Am 2:1273-1281, 1985 14. Pomerantzeff O, Pankratov M, Wang G-JI, Dufault P: Wide angle optical model of the eye. Am J Optom Physiol Optics 61:166-176, 1984 15. Lotmar W, Fankhauser F, Roulier A: Photocoagulation through the Goldmann contact glass. IV. A second model of the attachment to the Zeiss-Oberkochen photocoagulator and some improvements of the Siemens ruby laser coagulation. Arch Ophthalmol 82:314-319, 1969 16. Boldrey EE: A modified contact lens for peripheral retinal evaluation in pseudophakes. Ophthalmology 95:16-17, 1988 17. Fankhauser F, Rol P: Microsurgery with the neodymium: YAG laser: an overview. Int Ophthalmol Clin 25:55-84, 1985 18. Goldmann H: Biomicroscopie du corps vitre et du fonds de l’oeil. Paris: Masson 1957 19. Goldmann H: Fokale Beleuchtung. In: Straub W (ed) Die ophthalmologischen Untersuchungsmethoden, Vol 1, pp 104230. Stuttgart: Enke 1970 20. Rotter H: Zur Theorie der Spaltlampenmicroskopie des Augenhintergrundes. Graefe’s Arch Ophthalmol 152:689718, 1952 21. Mainster MA, Crossman JL, Erickson PJ, Heacock GL: Retinal laser lenses: magnification, spot size, and field of view. Br J Ophthalmol 74:177-179, 1990 22. El Bayadi G: New method of slit lamp microophthalmoscopy. Br J Ophthalmol 37:625-628, 1953 23. Schlegel HJ: Eine einfache Weitwinkeloptik zur Spaltlampenmikroskopischen Untersuchung des Augenhintergrundes. Doc Ophthalmol 26:300-308, 1969 24. Birngruber R, Lorenz B, Weinberg W, Greite JH, Gabel VP: Komplikationen bei der Laserkoagulation durch das Pan-
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25 fundoskop. Fortschr Ophthalmol 79:434-437, 1983 25. Drack AV, Burke JP, Pulido JS, Keech RV: Transient punctate lenticular opacities as a complication of argon-laser photoablation in an infant with retinopathy of prematurity. Am J Ophthalmol 113:583-584, 1992 26. Friberg TR: Principles of photocoagulation using binocular indirect ophthalmoscope laser delivery systems. Int Ophthalmol Clin 30:89-94, 1990 27. Irvine WD, Smiddy WE, Nicholson DH: Corneal and iris burns with the laser indirect ophthalmoscope. Am J Ophthalmol 110:311-313, 1990 28. Morley MG, Frederich AR Jr: Melted haptic as a complication of the indirect ophthalmic laser delivery system. Am J Ophthalmol 113:584-586, 1992 29. Pogrebniak AE, Bolling JP, Stewart MW: Argon laser-induced cataract in an infant with retinopathy of prematurity. Am J Ophthalmol 117:261-262, 1994 30. Robinfeld RS, Pilkerton AR Jr, Zimmerman LE: A corneal complication of indirect ophthalmic laser delivery systems. Am J Ophthalmol 110:206-208, 1990 31. Goldblatt NR: Designing ophthalmic laser systems. In: The Photonics Design & Applications Handbook, Book 3, 36th edn, pp 280-282. Pittsfield, MA: Laurin 1990 32. Dewey D: Corneal and retinal energy density with various laser beam delivery systems and contact lenses. SPIE 1423:105-116, 1991 33. Cibis P, Becker B, Okun E, Canaan S: The use of liquid silicone in retinal detachment. Arch Ophthalmol 68:590599, 1962 34. Landers MB III, Stefansson E, Wolbarsht ML: The optics of vitreous surgery. Am J Ophthalmol 91:611-614, 1981 35. Stefansson E, McCuen BW II, McPherson SD: Biconcave contact lens for examination and laser treatment of the fundus in normal and gas-filled phakic eyes. Am J Ophthalmol 98:806-807, 1984 36. Stefánsson E, Malcolm M, Anderson MM Jr, Landers MB III, Tiedeman JS, McCuen BW II: Refractive changes from use of silicone oil in vitreous surgery. Retina 8:20-23, 1988 37. Stefánsson E, Tiedeman J: Optics of the eye with air or silicone oil. Retina 8:10-19, 1988 38. Launay F, Laroche G, Limon S: Modifications de la réfraction aprés injections intra-oculaires de silicone liquide. J Fr Ophtalmol 5:417-425, 1982 39. Docchio F, Azzolini C, Brancato R: Refractive properties of interfaces due to the use of vitreous substitutes in vitreoretinal surgery: a raytracing approach. 1. Transpupillary laser photocoagulation. Lasers Light Ophthalmol 7:1-13, 1995 40. Docchio F, Azzolini C, Brancato R: Refractive properties of interfaces due to the use of vitreous substitutes in vitreoretinal surgery: a raytracing approach. 2. Endoocular laser photocoagulation. Lasers Light Ophthalmol 7:15-24, 1995 41. Docchio R, Azzolini C, Brancato R: Refractive properties of interfaces due to the use of vitreous substitutes in vitreoretinal surgery: a raytracing approach. 3. Modelling the effect of an air bubble within the irradiation path in endoocular laser photocoagulation. Lasers Light Ophthalmol 7:25-30, 1995 42. Azzolini C, Docchio F, Brancato R: Refractive hazards of intraoperative retinal photocoagulation. Ophthalmic Surg 24:16-23, 1993 43. Azzolini C, Brancato R, Trabucchi G, Camesasca F, Codenotti M, Verdi M: Endophotocoagulation through perfluorodecalin in rabbit eyes. Int Ophthalmol 18:33-36, 1994 44. Huy CP, Larricart P, Warnet JM, Haut J: In vitro laser decomposition of silicone fluid used in detachment of the retina. Ophthalmologica 204:23-26, 1992 45. Fankhauser F, Lotmar W: Skleraindentation and Photokoagulation. Acta Ophthalmol (Kbh) 48:253-260, 1970
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Fundamentals of optical fibers
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Fundamentals of optical fibers Marcel J. Poulain Laboratoire des Matériaux Photoniques, University of Rennes, Rennes, France
Keywords: optical fibers, physics of optical fibers, clinical applications
Introduction Optical fibers are now key components of highcapacity telecommunication networks. They are the result of 30 years of work by major glass and telephone companies. Millions of kilometers are produced every year at a very low cost (2-3 ¤ cents). They are a good example of an industrial success story, combining a challenge, doubts, errors, and luck. The concept of an optical fiber is probably very old, and the transmission of light through glass rods and filaments was already known by the glass makers of ancient Mediterranean civilizations. In 1870, the British scientist John Tyndall1 published the first scientific record of light guiding. In the 1950s, active research was carried out by Kapany with regard to transmitting images through bundles of various types of glass fibers.2 Between 1951 and 1966 most theoretical and experimental questions were addressed, laying the basis for the approaches to follow.3 Over the next 20 years, the key issue remained the manufacture of ultra-high-purity silica fibers. A milestone was reached one day in 1980 when the attenuation of a glass fiber was lowered to 20 dB/km, comparable to the damping an electric current undergoes in a copper wire. This was due to the use of the vapor phase deposition process which had been further improved, resulting in a spectacular decrease in fiber loss.4 The development of optical telecommunications has stimulated the use of optical fibers in other fields, such as manufacturing, sensing, instrumentation, imaging, and medicine.5 In addition to standard telecommunication fibers, a wealth of special fibers has been developed for applications as diverse as metal welding, spectroscopy, imaging, la-
ser surgery, non-invasive optical diagnosis, etc. In the long term, it cannot be excluded that important fiber applications will emerge in fields other than telecommunications. What is an optical fiber? Fiber structure An optical fiber is a waveguide: a beam of light launched at one end of an optical fiber travels down the output with negligible loss. This is illustrated in Figure 1. The optical fiber is made from two transparent materials, usually types of glass: the core and the surrounding cladding. The cross-section of most optical fibers is cylindrical, but other geometrical forms (rectangular, flower-shaped, D-shaped, honeycomb, etc.) have been developed to meet particular needs. A light beam that enters the fiber core is refracted according to Snell’s law: n0 sin θo = n1 sin θi
(1)
θo, θi, n0, and n1 are, respectively, angle of incidence, angle of refraction, refractive index of air, and refractive index of core. As long as the input angle is smaller than the limiting value θc, the refracted beam is totally reflected at the interface between the core and cladding glasses. Provided that the condition for light guiding is fulfilled n2 < n1, (n2 is the refractive index of cladding), and as long as the incident angle is smaller than a critical value θc, the refracted beam is totally reflected at the interface between the core and the cladding. The incident angle θc corresponds to the
Address for correspondence: Marcel J. Poulain, MD, Laboratoire des Matériaux Photoniques, Campus Beaulieu, Bât 10, University of Rennes, F-35042 Rennes, France. e-mail:
[email protected] Lasers in Ophthalmology – Basic, Diagnostic and Surgical Aspects, pp. 27–32 edited by F. Fankhauser and S. Kwasniewska © 2003 Kugler Publications, The Hague, The Netherlands
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Fig. 1. Light propagation in an optical fiber.
maximum angle of total reflection at the core/cladding interface and therefore only depends on n1 and n2 :
This sin θc value is usually referred to as the numerical aperture (NA) of the fiber. The larger NA, the larger the light gathering power of the fiber. Typical values of NA are 0.12 for telecommunication fibers, 0.20 or larger for instrumentation fibers. It should be noted that the retina of the human eye is made up of a large number of rods and cones having, to a large extent, the structure of an optical fiber: a dielectric cylindrical rod, a few micrometers in diameter, surrounded by a dielectric layer of slightly lower refractive index.6 Light guiding has also been demonstrated with metal pipes in which light is reflected onto the walls. However, a fraction of the beam is lost at each reflection on the metallic surface. Even with mirror quality polish, the beam is strongly attenuated after a few hundred reflections. In comparison, the damping undergone at each reflection of a beam zigzagging in a glass fiber is negligible.
fiber. The different modes can easily be seen when an helium-neon laser is coupled to a multimode fiber: the output beam is no longer homogeneous and gives a set of distinct spots when shone on a white screen. The number of different modes depends on core size, NA, and wavelength of the light. Single mode propagation is observed when the core diameter is small enough: this is the case for telecommunication fibers operating at 1.5 µm in which the core size is less than 10 µm. Single mode fibers are not only used for telecommunications, but also for special sensors and interferometric devices. However, light insertion in a small core is difficult, especially when coupled to an external laser source. An additional limitation is caused by the high energy density in the fiber core. Graded index fibers make up the third group of optical fibers. The interface between core and cladding glass can be described as a thick tube in which the value of the refractive index decreases continuously from n1 to n2. This design partly compensates for the path differences between the high and low modes. As a consequence, the beam shape is closer to the gaussian profile than in a simple multimode fiber.
Single mode and multimode optical fibers
Evanescent waves
A beam that propagates in an optical fiber is only allowed to travel a finite number of pathways, the so-called modes. The number of modes allowed depends on the geometry of the fiber, its numerical aperture, and the wavelength of the light beam. Analytical descriptions of the modes are derived from Maxwell’s laws in the weakly guiding approximation, which is the case for fibers of relatively low numerical aperture. In a multimode fiber, low order modes correspond to beams propagating along the fiber axis, while higher modes are characterized by larger θo angles. Path lengths are different for the various modes, and are shorter for the lower modes that require a shorter time to reach the fiber output. For this reason, the temporal width of light pulses is enlarged after transmission through a multimode
It may seem useless to have a cladding glass in an optical fiber since most transmissions take place in air that has a refractive index close to 1, which ensures light guiding. Indeed, there are numerous examples of single index fibers. However, these have serious limitations. The main reason lies in the refraction process at the interface between the two media of the respective refractive indices n1 and n2. Although geometry suggests that a light beam is entirely confined in a fiber core, Maxwell’s relationships show that electric and magnetic fields enter the cladding zone, with two major consequences: they may be subject to physical interactions leading to optical absorption, and they may escape into the outer medium when some conditions are fulfilled. In practice, optical losses are generated at contact points between bare fibers and
sin θc = (n22 – n12)1/2
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(2)
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Fundamentals of optical fibers
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holders; image transmission through bundles of single index fibers suffers from crosstalk between the adjacent fibers. Thus, it is necessary to insulate each fiber with optical cladding, the thickness of which depends on light wavelength and propagation modes. In multimode fibers, low modes are confined to the core, while higher modes penetrate deeper into the outer cladding. Optical transmission The major requirement of an optical fiber is its ability to transmit light with minimal attenuation. Normal optical components have a thickness of less than 1 cm, and the emphasis is on glass homogeneity rather than on high transparency. Specifications for optical fibers are more severe since thickness is exactly fiber length, i.e., meters or kilometers. The expression of transmitted light follows Beer’s law: It = Io exp (-α x)
(3)
where Io and It are the respective intensities of the incident and transmitted beams, x is the path length, and α the absorption coefficient. If a given transmission factor is required, the value of α should be 100,000 smaller for a 1-km fiber than for a 1-cm window. The α value of the fibers is expressed in decibels per meter or kilometer, dB/m or dB/km. For example, the transmission factor of a two-meter long fiber with a 0.5 dB/m attenuation is 80%, as 1 dB loss corresponds to 80% transmission. Intrinsic losses An ideally pure material has a transparency range which is limited by three loss mechanisms: electronic absorption in the UV-visible spectrum, lattice vibrations in the infrared, and Rayleigh scattering. This is exemplified by Figure 2 which shows the typical V shape of the theoretical losses of pure silica. Each material has a specific curve which correlates closely with its chemical composition and structure. The transparency range of silica and other oxide glasses encompasses the visible and near-infrared spectrum. Other materials are required for use at wavelengths greater than 2 µm. Extrinsic losses Real materials contain a set of impurities and defects that limit their performance. Large optical absorption arises from transition metal impurities (Cr, Fe, Co, Ni, Cu, rare earths), even at the subppm level. Anionic impurities such as hydroxyls OH or complex anions (sulphates, phosphates, etc.) are another source of optical losses. Finally, local defects, such as inclusions and bubbles, induce light scattering. The resulting absorption coefficient may be expressed as:
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Fig. 2. Evolution of the theoretical losses of pure silica versus wavelength.
K α = —— λn
(4)
where K is a constant, and exponent n can have different values: n = 0, corresponding to wavelength independent scattering, for defects larger than the wavelength, n = 4 for much smaller defects, and n = 2 (Mie scattering) for intermediate defects. Defects may be randomly distributed along the fiber, more abundant in the cladding or at the core/ cladding interface. Both defects and residual impurities have been drastically reduced in silica fibers, by means of vapor phase processes. Fiber quality is more difficult to achieve in multicomponent and exotic glasses. Insertion losses Coupling light from an external source, e.g., a laser, to an optical fiber can be tricky, especially when a high coupling efficiency is desirable. Coupling is carried out using a focusing lens or a mirror. In addition to Fresnel losses at the input end, part of the beam will escape into the cladding if the spot size or the angle of the focused beam is too large. The problem is less severe in multimode fibers because the core diameter may be large (e.g., > 100 µm). However, the beam profile at the output may be irregular, the more so as some propagation modes change when the fiber is moved. Power transmission Optical fibers are currently in use for laser power transmission, especially with industrial Nd:YAG lasers operating at 1.06 µm. The maximum power density is limited by fundamental phenomena, and there is a damage threshold for a given material at a given wavelength. In practice, this threshold is larger than current requirements, but various de-
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fects or external factors may promote fiber failure upon laser exposition. In the ideal situation, all laser energy is transferred to the fiber. In actual cases, the lost energy is a source of worry, not only because the cost of the laser increases in relation to the energy produced, but also because this lost energy will finally heat the fiber or the surrounding area. The first obvious requirement is that the laser power may not enter the polymeric cladding in the coupling area. While the optical attenuation of the fiber at the laser wavelength is an important factor, it is not the most critical. Point defects with a strong absorption coefficient, e.g., carbon and metal inclusion, are rapidly heated under laser irradiation. Fibers may melt and break locally. These defects also scatter light into the cladding and polymeric coating that protects the fiber. If the scattered intensity is great, coatings may be destroyed and fibers damaged. Fiber ends are critical areas because various impurities may be deposited onto them. These impurities contain water and carbon that have an active part in the failure mechanism. In practice, fibers for laser power transmission are optimized to minimize the absorbing defects,
and they are packaged in order to protect input and output ends from external grime and pollutants. Fiber fabrication Drawing a fiber from a glass rod is a simple exercise, which was already known in ancient Egypt, and is routinely achieved by glass artists or chemistry students. However, the manufacture of an optical fiber requires rigorous control for any contamination factors. Silica fibers are made by drawing high purity preforms at 2000°C using the setup schematically described in Figure 3. Preforms are rods in which the central part consists of a core glass of a higher index of refraction, while the external part is made from cladding glass. These preforms are prepared by a vapor phase process in which silicon chloride reacts with gaseous oxygen. This results in a very high purity material which contains extremely low levels of metal impurities and hydroxyl. Variations of the refractive index are achieved by modification of the vapor composition: germanium, phosphorous, and fluorine can be incorporated in this way. High quality preforms can
Fig. 3. Schematic representation of a draw tower for silica glass preforms.
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Fundamentals of optical fibers
31 and polycrystalline fibers. While the pulling rate is much slower than for glass fibers, this is not a serious problem for short length applications. Hollow core fibers are made from capillaries internally coated with dielectrics under controlled thickness. Choice of fiber Material
Fig. 4. Double crucible for drawing fibers from the glass melt.
be made by using other forms of chemical processing, for example, sol-gel. Preforms of low melting glasses can also be prepared by inserting a rod into a cladding tube or by using other classical methods of glass manufacturing. This is the case for fluoride, borosilicate, germanate, and chalcogenide glasses. Optical fibers may also be drawn directly from the melt using the double crucible method. Both core and cladding glasses are heated in two concentric crucibles at a temperature for which melt viscosity is large enough. Then a step index fiber can be drawn from the bottom of the double crucible, as shown in Figure 4. Special glasses are sometimes difficult to draw into fiber because of their tendency to devitrification. These problems are solved by adjustment of the composition of the glass and optimization of the process. An external polymeric coating is applied to protect the fiber from scratching, to limit the chemical attack of water, and to increase the mechanical strength. Epoxyacrylate resins are the normal coatings, but other polymers such as silicones and polyamide can also be used. Despite their hydrophobic properties, fluorinated polymers (PTFE, FEP) do not create an efficient barrier against hydrolysis, and lead to weaker fibers. Special techniques are used for growing sapphire
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Choosing the optimum optical fiber for a specific use may be an obvious task or a subtle exercise in the balance between various parameters: attenuation, reliability, toxicity, availability, and cost. Fibers from different materials can be used depending on choice criteria. Silica fibers offer a unique set of advantages: lowest attenuations, good mechanical and chemical resistance, transparency range extending from 300 nm to 2 µm. Silica fibers for telecommunications are inexpensive, but the cost increases when a particular size and NA are needed, requiring the manufacture of special preforms. Special silica fibers have been developed for UV transmission, either doped with OH or made from fused quartz. When low attenuation is not important, glass fibers made from borosilicate glass or polymer fibers (PMMA) are a cheaper alternative choice for use in the visible spectrum. Transmission in the mid-infrared spectrum cannot be achieved with silica. This is the case at 3 µm at the emission wavelength of the Er:YAG laser. Possible choices include fluoride fibers made from fluorozirconate glass, germanate glass, and sapphire. This latter material is crystalline, which raises the problem of the optical cladding. Such fibers can be used up to 4-5 µm. Fluoride fibers have low attenuation, as shown in Figure 5. Other materials are required at longer wavelengths.7 Fibers made from sulphide can be used up to 8 or 9 µm, and heavy chalcogenides (Se, Te) glasses allow transmission above 10 µm. However, they still have limitations with regard to optical
Fig. 5. Optical losses from a standard fluorozirconate glass fiber (from Le Verre Fluoré SA).
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losses and toxicity problems. Polycrystalline fibers made from silver halides and hollow core fibers have been developed as an alternative to these chalcogenide fibers at 10.6 µm. Even when fibers are designed for IR transmission, it is useful if they can also transmit visible light: the He-Ne laser beam at 633 nm is currently being used for alignment or for checking the delivery area. Fiber reliability As a general rule, the lifetime of the fiber mainly depends on chemical durability and applied stress. Failure corresponds to breaking, which happens as a crack propagating from a surface flaw. Real fibers have surface flaws resulting from their processing. The initial intrinsic strength of a fiber is related to its chemical composition, and ultimately to the chemical bond energy. Current values are 4 GPa for silica fibers and less than 1 GPa for fluoride and sulphide fibers. The initial strength decreases versus time when the fiber is subject to stress corrosion, or, put more simply, under permanent stress in a humid environment. This phenomenon has been extensively studied in silica fibers.8 As the main characteristic of a fiber lies in its flexibility, it will be moved and bent during operation. The minimum bending radius must be defined. This depends on fiber material and diameter: it is small (a few millimeters) for silica fibers, but it is larger for fluoride, chalcogenide, and polycrystalline fibers. The aging of fibers used for laser transmission also depends on the evolution of internal and end defects under laser irradiation. This aging effect is minimized if the defect density is low enough, and if water concentration and residual stress are controlled. Apart from worsening surface flaws, water has little influence on oxide glasses, which are insoluble and remain transparent. However, things are different for some glasses, such as phosphate and halide. The chemical action of water can weaken the fiber and reduce transmission if its ends and outer surface are corroded. Direct contact with liquid water must be avoided, and various solutions have been successfully employed in harsh environments. Current fiber applications Apart from their use in signal transmission for telecommunications, fibers are also being used in an increasing number of devices. The availability of non-silica fibers has expanded the field of the possible applications.9 Passive and active applications were already foreseen in the early stages of the development of fiber optics.3 The first group of
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applications covers optical fiber sensors, remote chemical analysis, thermal measurements and imaging, reflectometry, optical instrumentation, and also laser power delivery. Most medical applications of fibers, with their specific requirements, belong to this group. These could be expanded very significantly when technological progress leads to the discovery of cheaper disposable fibers. Active fibers are doped fibers from which a laser effect or amplification can be obtained. The large interaction length of the fiber, and the high energy density that can be used for optical pumping, are the main features supporting the development of active fiber devices. Another attractive point is the possibility of generating short wavelength signals using IR laser diodes as pump sources. Powerful laser fibers can be built in this way with very good beam quality, which is more difficult to achieve with semiconducting lasers. Cost aspects and reliability still limit practical applications, but significant changes are likely to occur in the comingyears. Solid-state and compact-fiber lasers could come to replace the large, noisy gas lasers. Conclusions Optical fiber technology has opened the door to a large number of applications, i.e. metal welding, spectroscopy, imaging, laser surgery, non-invasive optical diagnosis, etc. This is possible due to the fact that attenuation of light energy is extremely low, namely 20 dB/km, comparable to the damping an electric current undergoes in a copper wire. Here, the fiber structure and its optical performance are analyzed together with its various applications. References 1. Tyndall J: Proceedings of the Royal Institution of Great Britain 1:446, 1854 2. Kapany NS: Fiber optics. Part 1. Optical properties of certain dielectric cylinders. J Opt Soc Am 47:413, 1957 3. Kapany NS: Fiber Optics: Principles and Applications. New York, NY: Academic Press 1967 4. Miya T, Terunume Y, Hosaka T, Miyashita T: An ultimate low loss single-mode fiber at 1.55 µm. Electronic Lett 15: 106-108, 1979 5. Katzir A: Lasers and Optical Fibers in Medicine. New York, NY: Academic Press 1993 6. Enoch JM: Wavaguide modes in retinal receptors. Science 133:1353, 1961 7. Kokorina VF: Glasses for Infrared Optics. New York, NY: CRC 1996 8. Matthewson MJ, Kurkjan CR: Environmental effect on the static fatigue of silica optical fiber. J Am Ceramic Soc 71:177-183, 1988 9. Poulain M: Fluoride glass fibers: applications and prospects. Proceedings SPIE 3416:2-12, 1998
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On the application of optical fibers in ophthalmology
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On the application of optical fibers in ophthalmology Peter F. Niederer Institute of Biomedical Engineering, University of Zurich and ETH Zurich, Zurich, Switzerland
Keywords: optical fiber endoscopy, cyclophotocoagulation, intraocular application, clinical applications
Introduction Optical fibers are widely used today, both as a carrier of information on the one hand and as a carrier of energy on the other. While the transmission of information is associated with low power levels, high power transport is required if substantial amounts of energy have to be made available in short periods of time. The greatest use of optical fibers is currently in telecommunications where information is relayed over large distances, and minimal distortion and absorption, requiring as little power as possible, are a priority. In turn, applications of optical fibers in medicine are usually restricted to local distances, but are related to both low and high power needs, i.e., information transmission and energy intensive applications can likewise be found. Thereby, optical fiber-based diagnostic and therapeutic procedures have been established in a wide range of clinical disciplines.1 Particulary important in this regard are neurosurgery, dermatology, gynecology, and ophthalmology. A typical information transmission task in clinical medicine consists of remote imaging by way of a fiber bundle, relay lens system, or a GRIN (Gradient Index) rod, i.e., endoscopy. In ophthalmic endoscopy, tips have to be small (diameter of less than about 0.9 mm) such that GRIN-based tips are particularly advantageous (see below). Power requirements of image acquisition, transport, and presentation are per se minimal; in this procedure, aspects relating to image quality, and ease and safety of application are prominent. Another type of physiological or pathophysiological information that can be conveyed by optical fibers consists of various kinds of fluorescent signals. For the generation of such signals, moderate incident power levels are
needed, while the returning signals are in general weak. Other weak optical responses associated with possible fiber applications are, for example, the Doppler shift caused by moving targets, an effect which enables flow measurements in blood vessels or tissues in general, as well as the Raman effect which allows chemical compositions to be determined. Still another use of fibers is made by the transmission of the response of implantable sensor systems, e.g., pressure sensors. In contrast, the transport of higher amounts of energy is necessary for therapeutic applications such as coagulation or optical breakdown production. Moreover, in photodynamic therapy, the delivery of medium to high optical power is necessary. Low and high power transmission techniques, i.e., imaging and treatment needs, are sometimes combined, in particular when therapeutic procedures are carried out under indirect visual control with the aid of an endoscope. Furthermore, an important aspect is related to illumination which is needed for many diagnostic and therapeutic procedures in ophthalmology. Thereby, the illumination which is necessary in endoscopy has to be provided at energy levels at which no tissue damage occurs. In ophthalmic endoscopy, the sensitivity of the retina is the limiting factor. (Dangerous levels are in fact not usually reached; in contrast, insufficient light is the main problem in many cases.) A fiber bundle is often applied for carrying the light necessary for illumination. In contrast to imaging bundles, such fibers are not ordered, but, depending on the application, are mixed in order to obtain the desired distribution of light intensity within the target area. While illumination is typically performed with the aid of a xenon lamp, a popular source of optical
Address for correspondence: Professor Peter F. Niederer, Institute of Biomedical Engineering, University of Zurich and ETH Zurich, Gloriastrasse 35, CH-8092 Zurich, Switzerland. e-mail:
[email protected] Lasers in Ophthalmology – Basic, Diagnostic and Surgical Aspects, pp. 33–42 edited by F. Fankhauser and S. Kwasniewska © 2003 Kugler Publications, The Hague, The Netherlands
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energy in ophthalmology for therapeutic purposes is lasers. Of these, the argon, Nd:YAG and diode lasers are mostly used in connection with fiber-based tools.* The main advantage of laser radiation is the possibility of being able to concentrate the energy in a smaller spot size than can be achieved with ‘normal’ light. Furthermore, optimal wavelengths can be chosen which are adapted to the treatment procedure at hand, and precisely defined pulse lengths and bandwidths are available. This is particularly important in procedures involving fluorescence spectroscopy and photodynamic therapy. However, optical energy can be provided by other sources than lasers, e.g., a xenon lamp. Even the sun has been tested for this purpose.2 In all applications, an important distinction has to be made with respect to the environment in which a fiber, a fiber bundle, a GRIN rod, or a lens is used: In air, the refractive power of surfaces is usually higher than in biological fluids because the difference in the index of refraction of typical materials used for fibers, fiber tips, or lenses, e.g., silica (index of refraction, n = 1.46), is smaller in case of fluids (n = 1.33 for water) than of air (n = 1.0003, all values given in relation to vacuum). Accordingly, the optical pathways are considerably different when such interfaces are present, and the focal lengths or angles of divergence of light beams can also differ decisively. For most applications of fiber systems used in ophthalmology, the particular aspects of wave optics can be disregarded and the analysis restricted to geometrical optics. Ray-tracing is the design method of choice for this purpose.
composition, coherence length/bandwidth, angle of beam divergence, and pulse length. Moreover, from a user’s point of view, installation requirements, size, weight, etc., are of concern. Solid-state lasers have the advantage of minimal maintenance requirements and ease of operation. In addition, diode lasers are small, light weight, simple in terms of instrument complexity, and are easily replaceable in case of malfunction. While the details of the mode composition as such are of minor importance in most ophthalmic applications, pulse length, beam divergence, focal spot size (which, in turn, depends on modal composition and divergence), and energy delivery characteristics are decisive. Depending on the application, lasers are utilized in the continuous wave (cw), free running, and pulsed mode; moreover, Qswitched as well as mode-locking techniques are used for special purposes. The argon (gas) laser is often utilized mainly for two reasons. Firstly, one of its main wavelengths (514 nm) is well adapted to the absorption characteristics of blood, which makes the instrument especially suitable for coagulation purposes; secondly, it has been widely introduced in the medical community. In turn, the Nd:YAG (solid-state) laser, working at 1064 nm, is characterized by its high power capacity and its capability of producing optical breakdowns when run under Q-switched conditions. Furthermore, by frequency-doubling, a wavelength similar to that of the argon laser can be obtained (KTP laser). Finally, infrared diode lasers are well suited for clinical applications because of their small size, such that they can even be installed on the operator’s spectacles. Since the ophthalmic applications of lasers are treated in depth elsewhere in this book, this subject will not be examined in further detail here.
Light sources and optical fibers Lamps used for illumination Lasers Lasers used for therapeutic purposes are essentially characterized by their wavelength, power, mode
The light of the xenon lamp is characterized by a wide spectrum covering the range of ultraviolet to infrared, with a bias towards the blue limit (Fig. 1).
Fig. 1. Typical spectrum of the xenon lamp (UV and infrared filtered). * Excimer lasers are usually employed in corneal surgery; however, this procedure is carried out without the use of optical fibers.
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On the application of optical fibers in ophthalmology On the one hand, its high blue content is advantageous for imaging3 (CCD and CMOS image sensors have a low sensitivity in the blue range), on the other, UV is potentially cancerogenous, such that the UV content of the illumination has to be carefully controlled. In contrast, halogen lamps are non-problematic in this regard, however, they are not suitable for true-color imaging due to their weak blue spectral content. Infrared always has to be filtered because excessive heating has to be prevented and, moreover, sensor response at these wavelengths disturbs the visible (and useful) image information. Optical fibers, optical transmission systems and exit beam shaping The characteristic quantities describing the properties of optical fibers are the numerical aperture, wavelength/diameter ratio, transmission characteristics in terms of bandwidth and damping, as well as water solubility.4 The material used is usually silica, but quartz and chalcogenide fibers can also be chosen for UV and long wavelengths, respectively. Typical core diameters vary between 50 and 600 µm, while numerical apertures cover the range of 0.10.4. A typical fiber transmission system (Fig. 2) for low or high power applications in essence consists of three parts, namely, firstly, an energy control and entrance coupling, where the light is optically coupled into the fiber or fiber bundle, secondly, the fiber or fiber bundle itself, and, thirdly, the exit system. Thereby, the design of the entrance and exit optics has to be adapted to the physical characteristics of
35 the fiber(s), in particular the aperture, as well as the dimensions and divergence of the laser beam (or intensity distribution of a lamp). Furthermore, high power transmission by way of optical fibers is critical with respect to material purity (e.g., the presence of OH- ions) and optical coupling.5 Impurities can cause irregularities in energy transport, local energy concentrations, and ‘hot spots’ causing fiber damage. Moreover, hot spots can occur in case of an unfavorable optical design of the transmission system. A careful design of coupling of optical energy into a fiber is necessary, such that the energy is actually coupled to the fiber without damaging the cladding or fiber jacket. Fiber tips are of particular importance with regard to medical applications. A variety of design strategies can be found, among which, self-focusing tips and beam-shaping microdevices. By the controlled melting of a fiber tip, a rounded surface at the fiber exit exhibiting focusing properties can be achieved (self-focusing tip).6 Figure 3a shows tips in various stages of melting, while in Figure 3b, light propagation is visualized in fluorescein together with ray-tracing modelling. As mentioned in the introduction, the focusing characteristics in air and water differ considerably. Optical breakdown though an optical fiber (Fig. 4) can be achieved in water, although the focusing characteristics are worse than in air.7 Since optical breakdown production is associated with short pulse/high power transmission, care has to be taken with respect to the entrance coupling and fiber quality. Side-firing tips are of use in subconjunctival applications or when the eye is being accessed from
Fig. 2. Laser/fiber/application/observation system. The laser cavity produces coherent light, the intensity of which is controlled by a measurement and adjustment device. The beam expander improves the beam focusing properties. F1 and F2 are focal distances, respectively. Observation is made through a microscope.
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a.
b. Fig. 3. a. Tips in various stages of melting. b. Focal characteristics in air (top) and water (bottom), ray-tracing and visualization. An He-Ne laser was used for demonstration purposes, whereby the beam in air was visualized in dry ice vapor.
behind.8 Light propagation as well as two possible optical designs are shown in Figure 5. Furthermore, beam-shaping can be achieved with the aid of appropriately-designed quartz tips (Fig. 6).5,6 Such devices are particularly advantageous in surgical applications where a precise energy concentration at the tip is desired. Figure 6 shows one particular design, but numerous other configurations are also conceivable.
Fig. 4. Optical breakdown in water (Q-switched Nd:YAG) through an optical fiber.
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Fig. 5. Side firing tip. Visualization in water/fluorescein (top), two possible optical designs (bottom). Sapphire elements are used for beam redirection and focusing, x’ denotes the distance of the beam waist (DC’) from the exit surface.
Fig. 6. Sapphire tips. The beam propagation is shown (top, At denotes the cone angle, div the inclination of the particular beam depicted, i the angle of incidence on the side and θ the angle of incidence on the exit surface). Rays of which the angle of incidence i is smaller than the angle of total reflection will leave the cone, as can be seen in the beam visualization experiment. Optics in air and water are different.
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Endoscopy An endoscope is a remote imaging device in which image transmission is a key element. Such instruments are widely used in medicine today; they can also be applied for special procedures in ophthalmology,9 such as imaging in the posterior chamber and the vitreous cavity, particularly in case of opacities of the media. An endoscope can either be made with the aid of a fiber bundle, which is advantageous whenever flexibility is required, for example, for inspection of the lacrimal duct. Imaging quality increases approximately quadratically with the diameter of such an endoscope, because the number of imaging fibers corresponds to the number of pixels in the image. If flexibility is not required, a GRIN rod is preferable, in particular in a small endoscope, due to its better image quality. Figure 7a shows an ophthalmic GRIN endoscope. When the image quality of fiber bundles and GRIN endoscopes of the same diameter are compared, the superior image quality of the GRIN type becomes apparent (Fig. 8). Illumination is achieved by fibers arranged around the central GRIN rod (Fig. 7b); furthermore, an empty channel is provided in the tip
(not shown in the schematic drawing) in order to insert a fiber for therapeutic applications. Inspection of the lacrimal duct is best carried out with the aid of a fiber bundle endoscope (Fig. 9), because flexibility is advantageous in this application (although the first results were obtained with a rigid endoscope).10 A further, presently still experimental, application of an imaging fiber bundle in ophthalmology consists of a portable slit lamp. With the aid of a fiber bundle, it is possible to project a slit of acceptable quality into a handheld camera which is brought into contact with the cornea (Fig. 10). However, a relatively expensive fiber bundle containing 100,000 individual fibers has to be used for this purpose, in order to achieve good results. Selected clinical procedures Coagulation, controlled destruction of cells, and perforations are typical procedures in which optical fibers can be used. Ab externo and ab interno applications have been proposed in ophthalmology, and contact and non-contact modes have been described.
a.
b. Fig. 7. a. Ophthalmic endoscope (GRIN type). b. Schematic view of a GRIN endoscope. The optical system images the exit surface of the GRIN rod onto the image sensor. Illumination fibers are arranged around the GRIN rod. A selfoc lens is used for image acquisition.
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Fig. 8. Comparison of the resolution of a GRIN rod versus fiber bundle optics of the same diameter (0.5 mm). The working distance is 5 mm in air.
Intraocular procedures
Fig. 9. Endoscopic image of a lacrimal duct, obtained with a fiber bundle endoscope. (Reproduced by courtesy of Dr F.M. Sens, University Hospital Basel.)
Intraocular treatments where fiber-based irradiation systems are used mainly consist of photocoagulation, sclerostomy,11,12 and trabeculoplasty.13,14 While photocoagulation consists of delivering optical energy such that coagulation of the blood is achieved and a blood vessel becomes permanently blocked, sclerostomy and trabeculoplasty are aimed at the removal of tissue by deposition of optical energy. Intraocular photocoagulation with the aid of a fiber delivery system, for example, on the retina, is a routine procedure to date. As mentioned earlier, argon or frequency-doubled Nd:YAG (KTP) laser light is mostly used for this purpose because the relation of the absorption characteristics of the blood versus the surrounding tissue is particularly high in this range of wavelengths. Relatively unsophisticated and simple silica fiber systems and routine optics are utilized. In turn, sclerostomy is applied to improve
Fig. 10. Comparison of the slit of a conventional slit lamp with a slit produced by a fiber bundle. Magnification is 10× (left) and 40x (right), the fiber bundle used has 30,000 individual fibers (top) and 100,000 fibers (bottom).
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Fig. 11. Combined needle/fiber device for endocular applications. The fiber is advanced through the channel of the needle (visualization in water/fluorescein).
Fig. 12. Schematic view of the endoptic application of a combined endoscope/fiber. The fiber tip can be seen through the endoscope. The insert shows an experimental application in a ciliary body (pig eye).
the outflow of chamber fluid in open-angle glaucoma patients by creating an artificial hole in the chamber angle. As the drainage improves, intraocular pressure decreases to normal levels. The aim of using laser energy for this purpose is the creation of a more stable channel compared to a rapidly healing puncture made mechanically.13,14 A similar aim can be seen in trabeculoplasty, in which the trabecular meshwork in the chamber angle is irradiated. Nd:YAG laser pulses, typically delivered through a 200-300 µm fiber, are generally used for sclerostomy. An experimental procedure was proposed by Dürr et al.11 whereby a 200-µm silica fiber was inserted into a 20-22 gauge hypodermic needle, and advanced to the desired location (Fig. 11). Typically, up to 10 J were required to achieve a perforation in an excised pig eye (cw or free-running Nd:YAG laser). Moreover, intracanalicular trabeculostomy, with the aim of improving the outflow of chamber fluid using the Er:YAG laser (2940 nm) and a 300µm quartz fiber, has been evaluated experimentally by Kampmeier et al.8 The pulses were applied ab externo with the aid of a side-firing tip. The advantage of Er:YAG radiation is the high absorption in
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water, such that minimal energy is required to achieve a therapeutic effect.15 For the same reason, the Er:YAG laser is suitable for cataract removal.16 However, special fibers are needed to transport the long wavelength of the Er: YAG laser with minimal loss. With the aid of zirconium-fluoride-based and sapphire fibers, satisfactory lens emulsification has been reported.17 The intraocular application of laser light under endoscopic control18,19 is advantageous when treatment is necessary at a location that cannot well be visualized through the pupil, or in case of opaque media. Figure 12 outlines the experimental application of a combined endoscope/fiber system. Transscleral procedures The transscleral application of optical energy through optical fibers can be carried out in the contact or non-contact mode. In either case, the transmission characteristics of the sclera are decisive (Fig. 13).20 Furthermore, in the contact mode, transmission of the sclera is dependent upon the pressure with which the fiber tip (or exit optics) is applied. 21,22 With
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Fig. 13. Transscleral light propagation. The relative back/forward scatter intensities at various angles of incidence are shown for argon, dye, diode and Nd:YAG lasers, respectively. The Nd:YAG laser radiation has the highest relative transmission.
increasing pressure, transmission increases because of the rearrangement of fibers in the sclera, as well as because of water displacement. Cyclodestruction and cyclophotocoagulation are procedures that are commonly applied in case of refractory glaucoma.23,24 The aim of treatment is to reduce the production of aqueous humor such that the intraocular pressure becomes lower. The Nd:YAG and diode lasers, in combination with fiber delivery systems, are most often used for this purpose. Yet, the successful application of krypton laser irradiation, delivered by way of optical fibers, has been reported as well.25,26 A further treatment where laser light and fiber delivery systems are involved is photodynamic therapy.27 By administration of a photosensitizer, usually photofrin, hematoporphyrin, or phthalocynaine derivatives, cells to be destructed are selectively perfused. Upon irradiation, these cells are deleted as the result of optically-induced oxygen reactions.28-30 Although it is not easy to reach sufficient illumination levels, transscleral illumination has a number of advantages: the corneal region remains free, disturbing reflections from the cornea and intraocular surfaces are absent, and an even distribution of light is achieved. The best results are obtained if a ring of light carrying fibers is applied at the limbal region. Conclusions and outlook A number of noteworthy applications of optical fibers in ophthalmology has been presented in this communication. However, this overview is not exhaustive, since other applications are conceivable or have also been made. Yet, the use of optical fibers is now universal and is part of many standard procedures. As, in particular, micro-optics, fibers and fiber bundles, diode lasers and xenon illumination techniques
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become less expensive, easier to use, and associated with fewer maintenance needs, fiber-based applications in ophthalmology can be expected to become further developed and more widely used. However, after having outlined a number of ophthalmic procedures in which the use of optical fibers is advantageous, it should be pointed out that some applications can be critical. For example, problems can arise whenever small optical (wavelength) shifts have to be measured. Doppler measurements are adversely influenced by a moving fiber because an artifactual shift of wavelength arises.31 Similarly, Raman or other types of spectroscopical analyses can become critical when optical fibers are used and motion artefacts cannot be avoided. Moreover, in case of high power transmission, fiber impurities or imprecise optical alignment of coupling optics are a possible source of malfunction, and particular care is required when such applications are being carried out. In conclusion, research and development in the area of fiber applications in ophthalmology is proceeding and new application techniques can be expected. New types of fibers are being evaluated for their usefulness in medicine: hollow plastic and glass fibers,32 as well as fibers manufactured from special glasses33 (rare earth doped), may open new ways for the delivery of optical energy. And finally, the advent of fiber lasers34 may lead to simpler and less expensive instrumentation. Acknowledgments Many of the results presented in this article were obtained and worked out by Dr Pascal O. Rol, University Eye Clinic, University of Zurich, and Adjunct Professor, University of Miami, who lost his life in an aircraft accident near Zurich in January 2000.
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References 1. Katzir A: Lasers and Optical Fibers in Medicine. New York, NY: Academic Press 1993 2. Feuermann D, Gordon JM: Solar surgery: remote fiber optic irradiation with highly concentrated sunlight in lieu of lasers. Opt Eng 37:2760-2767, 1998 3. Von Orelli A, Lehareinger Y, Rol P, Niederer P, Doswald D, Felber N: High-definition true-colour television for use in minimally invasive medical procedures. Tech Health Care 7:75-84, 1999 4. Poulain MJ: Fundamentals of optical fibers. This volume 5. Rol P, Niederer P, Fankhauser F: High-power laser transmission through optical fibres: applications to ophthalmology. In: Wolbarsht M et al (eds) Laser Applications in Medicine and Biology, Vol 5. pp. 141-198. New York, NY: Plenum Press 1991 6. Rol P, Fankhauser F, Kwasniewska S, Niederer P: A comparison of ophthalmic fiber optic microfocussing systems. Lasers Light Ophthalmol 2:115-124, 1988 7. Rol P, Niederer P, Fankhauser F, Arigoni M, De Haller E: Q-switched pulses and optical breakdown generation through optical fibers. Lasers Light Ophthalmol 3:213-219, 1990 8. Kampmeier J, Stock K, Hibst R, Lang GE, Steiner R, Lang GK: Intracanalicular trabeculostomy: a new approach in glaucoma surgery. Klin Mbl Augenheilk 212:159-162, 1998 9. Rol P, Jenny R, Beck D, Fankhauser F, Niederer P: Optical properties of miniaturized endoscopes for ophthalmic use. Opt Eng 34:2070-2077, 1995 10. Sens FM, Rol PO, Yanar A, Robert YCA: Endoscopy of lacrimal ducts with a rigid GRIN-endoscope. Ophthalmologe 97:418-421, 2000 11. Dürr U, Fankhauser F, England C, Kwasniewska S, Van der Zypen E, Henchoz PD, Rol P, Niederer P: Experimental Nd:YAG, diode and Ho:YAG laser sclerostomy performed ab interno and ab externo. Laser Light Ophthalmol 5:83-93, 1992 12. Latina MA, Dobrogowski M, March WF, Birngruber R: Laser sclerostomy by pulsed dye laser and goniolens. Arch Ophthalmol 108:1745-1750, 1990 13. Atmaca LS, Simsek T: Efficacy of argon laser trabeculoplasty in primary open-angle and pseudoexfoliative glaucoma: longterm follow-up. Ann Ophthalmol 33: 216-220, 1991 14. Sellem E: Trabeculoplasty: are there still indications for this procedure? J Fr Ophtalmol 24:1100-1102, 2001 15. Dietlein S, Kriegstein GK: Erbium:YAG laser trabecular ablation. This volume 16. Dodick JM, Pahlavi IA: Lasers in small-incision cataract surgery. This volume 17. Neubaur CC, Stevens G: Erbium:YAG laser cataract removal: role of fiber-optic delivery system. J Cataract Refract Surg 25:514-520, 1999 18. Jacobi PC, Dietlein TS, Krieglstein GK: Microendoscopic
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trabecular surgery in glaucoma management. Ophthalmology 106:538-544, 1999 Rol P, Sasoh M, Manns F, Edney P, Niederer P, Parel JM: Experimental intraocular laser surgery with a GRIN laser endoscope. SPIE 2673:50-53, 1996 Vogel A, Dlugos C, Nuffer R, Birngruber R: Optical properties of human sclera, and their consequences for transscleral laser applications. Laser Surg Med 11:331-340, 1991 Rol P, Barth W, Schwager M, Zuber N, Fankhauser F, Fankhauser S, Niederer P: Devices for the control of laser transmission across the sclera during transscleral photocoagulation. Ophthalmic Surg 23:459-465, 1992 Rol P, Fankhauser F, Niederer P: Influence of the pressure exerted on the sclera during transscleral photocoagulation. SPIE 1877:249-252, 1993 Hamard P, Gayraud JM, Kopel J, Valtot F, Quesnot S, Hamard H: Contact transscleral cyclophotocoagulation with diode laser in refractory glaucoma: a mid term follow-up study: analysis of 50 patients followed over 19 months. J Fr Ophtalmol 20:125-133, 1997 Schlote T, Kreutzer B, Kriegerowski M, Knorr M, Thiel HJ: Diode laser cyclophotocoagulation in refractory glaucoma. Klin Mbl Augenheilk 211:250-256, 1997 Raivio VE, Immonen IJR, Laatikainen LT, Puska PM: Transscleral contact krypton laser cyclophotocoagulation for treatment of posttraumatic glaucoma. J Glaucoma 10:7784, 2001 Raivio VE, Immonen IJR, Puska PM: Transscleral contact krypton laser cyclophotocoagulation for treatment of glaucoma in children and young adults. Ophthalmology 108: 1801-1807, 2001 Tsilimbaris MK, Naoumidi II, Kozombolis VP, Naoumidi TL, Daskalakis M, Pallikaris IG: Ciliary body PDT in pigmented rabbit eyes: effect of single and repeated treatment. Curr Eye Res 20:469-479, 2000 Van den Bergh H: This volume Yoneya S: This volume Birngruber R: This volume Wardell K, Jakobsson A, Nilsson GE: Laser-Doppler perfusion imaging by dynamic light scattering. IEEE Trans Biomed Eng 40:309-319, 1993 Serafetinides AA, Rickwood KR, Fabrikesi ET, Chourdakis G, Anastassopoulou N: Pulsed 3 µm laser radiation transmission through hollow plastic and hollow glass waveguides. Proc Biomedical Sensors, Fibers, and Optical Delivery Systems. SPIE 3570:12-19, 1998 Lezal D, Pedlikova J, Karel M: Special glasses for passive and active IR fibers for medical and biomedical applications. Proc Biomedical Sensors, Fibers, and Optical Delivery Systems. SPIE 3570:53-61, 1998 Minelly JD: Fiber lasers: launch into medicine, aerospace and material processing. Photonics Spectra 30:128-133, 1996
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Laser speckle Thorsteinn Halldórsson European Aeronautic Defence and Space Company-EADS, Corporate Research Center Germany, Munich, Germany
Keywords: speckle phenomenon, coherent/incoherent light, visual acuity, ocular blood flowmetry
Origin of speckle A phenomenon of light illumination unnoticed by most people is its granular structure, which is normally smoothed by the mixing of different colors in lamp- or natural sources of light, but which can also be observed in certain circumstances, such as during the twinkling of stars or the dancing colors of the sun’s rays transmitted through the leaf cover of trees. Many scientists have investigated this granularity since the time of Newton (1642-1727), who noticed that the scintillation or twinkling could be observed for stars, but not for the planets.1 Today, speckle is casually seen as the intensity and color fluctuation in the distant headlights of a car, or as a mottled light pattern on rear projection screens. Speckles can even be seen in the security holograms on credit cards, which can easily be studied with a magnifying glass. Since the invention of the laser in 1960, there has been a growing interest in the speckle phenomenon, since color smoothing is not present in monochromatic light, and speckle patterns become clearly visible with laser illumination of the surface of many objects. For this phenomenon, it is necessary that the object diffuses the light like a piece of paper, an unpolished metal surface, or that the light propagates through a medium with randomly refractive index fluctuations. On the other hand, no speckle is observed by specular reflection of laser light on polished surfaces such as mirrors, or by laser light transmission through non-turbid liquids or clear glass-like lenses – with the exception of speckles generated at the end of multimode fibers. From this we can conclude that speckle arises from the splitting of the coherent laser beam into a
number of wavelets originating from microscopic elements in the medium, which travel over unequal distances to every point in space surrounding the surface, where they interfere with random but stable phase differences, leading to a time-independent fixed 3-D intensity pattern. If the observation point is moved, the path lengths travelled by the scattered components change, and a new independent value of intensity may result from the interference process. The typical appearance of speckle is depicted in Figure 1 in the case of the scattering of a laser beam from a diffuse surface, which looks similar to a mixture of grains of salt and pepper. Since a stable interference occurs at every point in space, a photographic plate placed at any distance from the object can record the speckle pattern as shown in Figure 1 (objective speckles). If the photographic plate is replaced by an imaging system such as the human eye, shown in Figure 2, the pupil can be considered a new source of wavelets of a random phase, where they are additionally diffracted and imaged by the eye lens on the retina. The interference between these diffraction images is responsible for the random 2-D intensity distribution on the retina, sensed as a granular structure (subjective speckle). A prerequisite for the development of speckle is either sufficient temporal or spatial coherence of the light source. Temporal coherence of a light source is defined as the maximum time delay ∆t between two or more light waves emitted by the source, where the phase shift between the waves in the course of time is stable enough for their interference. This corresponds to a path difference ∆lt of the waves in the direction of the light propagation, given by the relation:
Address for correspondence: Thorsteinn Halldórsson, European Aeronautic Defence and Space Company-EADS, Corporate Research Center Germany, Willy Messerschmittstr. 85521 Ottobrunn, P.O.B. 81663 Munich, Germany. e-mail:
[email protected] Lasers in Ophthalmology – Basic, Diagnostic and Surgical Aspects, pp. 43–49 edited by F. Fankhauser and S. Kwasniewska © 2003 Kugler Publications, The Hague, The Netherlands
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T. Halldórsson produced by: ∆lr = λ z/ds = λ/θ where λ is the mean wavelength, ds the diameter of the source, and z the distance between the source and the observation point, and the ratio: θ = ds/z
Fig. 1. Speckle pattern generated on a rough surface illuminated by a laser.
∆lt = c ∆t where c is the speed of light. Only waves with a shorter path difference than ∆lt can contribute to a speckle pattern. This path difference, called coherence length ∆lt, depends on the spectral bandwidth of the light source ∆λ by the relationship: ∆lt = λ2/∆λ White light at the mean wavelength of λ = 0.5 µm and a bandwidth of ∆λ = 350 nm has a coherence length of less than 1 µm, but a laser with a spectral bandwidth of 0.001 nm has a coherence length of 25 cm. Since the average roughness of diffuse surfaces is much greater than 1 µm, the number of waves passing a point in the space near to a rough scattering surface, which can interfere, is extremely low by illumination with white light, but is very high for laser light, which explains the pronounced speckle pattern produced by laser light, contrary to that of a white light source. However, there is also a second condition necessary for the development of speckle. If the size of the illuminated spot on a scattering surface, or the light source itself, is very small, the phase shifts between waves in a direction transversal to the wave propagation direction from the source become stable enough to interfere – thereby generating a speckle where the waves do overlap, for instance, with the help of an imaging lens, and the source is said to be spatially coherent. The largest transversal distance of waves with a stable phase shift, called spatial coherence length ∆lr of the source, is
is equal to the apparent angular size of the source at the observation point. This relationship tells us that a small incoherent light source with a broad spectral bandwidth, and thus low temporal coherence, can possibly possess sufficient spatial coherence to produce speckle. As mentioned above, this can happen with several natural and technical light sources with a broad spectrum, such as the sun, the stars, or lamps, if they are far enough from the observer. To obtain an estimate of this effect, we have to look at the spatial coherence corresponding to the resolution limit of the human eye of 1 arcmin, which is equal to θ = 3 × 10-4 rad at λ = 0.5 µm, with ∆lr = 1.4 mm. As a consequence of this, a star with a much smaller apparent angular size than this resolution limit of the eye has a spatial coherence length much larger than the diameter of the pupil. It then becomes clear that refractive index fluctuations in the atmosphere induced by winds or turbulence can generate slowly varying interference on the retina, seen as scintillation of the starlight. As Newton pointed out, this is much less noticeable for the planets, because of their larger apparent size and, consequently, their lower spatial coherence. A speckle due to the spatial coherence of laser source illuminating the retina of the eye can also be generated due to fluctuation in the refractive index of the lens or the vitreous humor, if the size of the illuminating source is very small. Another example of speckle due to spatial coherence is that of displays. The spatial coherence length of modern electronic lamp projectors with a small exit pupil can be of the order of several tenths of a micrometer, which is sufficient for generating speckles on a rear projection screen made of a micro-lens array for expanding the projector beam, without introducing an additional path difference
Subjective speckle pattern Fig. 2. Imaging of speckle by the eye.
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in the transmitted light waves. On the other hand, no speckles appear by front projection on a scattering surface with the same projector, because, in this case, the average surface roughness of the diffusing screen is longer than the spatial coherence length of the source. Most lasers have high temporal coherence due to their short spectral width, but also high spatial coherence, due to their small emitting aperture. The spatial coherence is greatly reduced by the scattering of an extended laser beam on a rough surface. In this case, temporal coherence plays the major role in the development of a speckle pattern of the laser illumination. The temporal and spatial coherence of a source are not two different types of coherence, but rather different ways of interpreting the correlation function of light waves, which unify in the mutual coherence function of light waves theory,2,3 but should be kept apart in the practical explanation of speckle phenomena. Here, only time-independent speckle patterns originating from the surface of rigid objects have been considered. If the scattering medium changes with time, the speckle pattern also evolves in time, and is known as ‘dynamic speckle’. But speckle can also be generated by random phase shifts produced by multiple scattering inside the volume of low absorbing optically diffuse materials, such as biological tissue. Since a biological material is not a rigid body and it incorporates time-varying liquid transport, spatial-time modulation of the speckle, so-called ‘bio-speckle’, is observed. The speckle phenomenon is not only encountered by light waves, but related effects also arise in other regions of the electromagnetic spectrum, from Xrays to radio waves. Everyone using a radio receiver in a car driving at the range limit of a radio station is accustomed to the rapidly varying receiving quality, which is due to the field intensity clutters of radio waves, which correspond to speckle. In ultrasound imagery in medicine, the images are overlapped by an acoustic speckle field, which limits their contrast and resolution.
Statistical properties In general, the statistical properties of speckle patterns depend on both the coherence of the incident light and the detailed properties of the random surface or the multiple scattering centers inside the medium. If the scattering introduces path differences greater than one wavelength and the laser is highly coherent, and a large number of scattering centers contribute to the intensity at a point in the observing plane, it can be shown that intensity I at such a point in such a speckle pattern has a negative exponential probability density function and that the probability that the intensity exceeds threshold I is given by:
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Fig. 3. The probability that the intensity in a speckle field of coherent light exceeds a certain level I. The probability for incoherent light is shown for comparison.
P(I) = e-I/ which is shown in Figure 3. The ratio of the standard deviation to the mean intensity is unity for this distribution, so the contrast in the speckle pattern is equal to one. It is interesting to note that the most probable intensity at any point in such a speckle pattern is zero, and that intensities well above the mean value are quite common. The intensity probability function of white light such as sunlight, which is concentrated around the mean value , is shown in Figure 3 for comparison. Another specific property of the speckle is its size distribution, or the coarseness of the granularity in the speckle pattern. The smallest diameter of speckle d is determined by the diameter of the illuminated spot and the distance z of the observer from the spot: dmin = λ z/L or expressed in spatial frequency: fmax= 1/dmin = L/λz shown in Figure 4 for a rectangular spot, with L as the side length, the distance from the spot z, and λ the illumination wavelength. The general conclusion to be drawn from the spectrum in Figure 4 is that, in any speckle pattern, large-scale-size fluctuations are the most populous, and that no scale sizes are present beyond the small size cut-off fmax, which corresponds to a speckle diameter of dmin. The size of the free-space, objective speckles increases linearly with the distance from the illuminated spot. In the case of an imaging system such as the eye, the pupil can be considered a new source of randomly scattered light from a spot with a circular aperture diameter Dp and a distance z, then: d = 1.22 λ z/Dp
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Fig. 4. Probability of occurrence of different spatial frequencies in a speckle pattern.
If the eye is focused on a scattering surface far away, then the image distance z is equal to the focal length f with: d = 1.22 λ f/Dp Since the smallest speckles imaged on the retina are only limited by the diffraction of the eye pupil, the smallest size for f = 17 mm at pupil diameter of Dp = 8 mm is d = 1.3 µm. For a pupil diameter of Dp = 2 mm, the smallest speckle diameter becomes d = 5.2 µm, which is just at the resolution limit of the eye, limited by the distance of the photoreceptors of 5 µm.
General applications Since speckle patterns are bothersome in any kind of imagery – with laser, radar, or sound – several
techniques have been developed in all these areas in order to dispense with them. But, on the other hand, optical speckle patterns contain valuable information about the surface of objects, and have become quite useful in metrology for the measurement of surface roughness and object deformation, which are the subjects of speckle photography and speckle interferometry. Stellar speckle interferometry has become a valuable tool in astronomy. It is based on the fact that telescope images, blurred by atmospheric turbulence, arise from the overlap of a number of different individual speckled images, each of which containing the astronomical information mixed with that of the atmosphere fluctuation. By compensating for the speckle patterns in individual short exposure photographs, the original astronomical information can be extracted. The measurement of dynamic speckle patterns has become an important field in speckle flowmetry in medicine for studying blood flow in biological tissue, and will be discussed further in the application of speckle on the eye in the section on Biospeckle flowmetry versus laser Doppler flowmetry below.
Deterioration of vision and methods to cure this The presence of speckle in an image reduces the ability of a human observer to resolve the details. The spatial information present in the image, particularly in the fine details, is masked by the structure of the speckle pattern. Because of this problem, speckle noise has been an obstacle in the application of lasers in several fields, such as holographic microscopy and holographic endoscopy,
Fig. 5. Argon laser beam: a. direct beam; b. spot of 9 mm in diameter with a speckle pattern on a laser video screen.
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and has also been a major roadblock in the development of laser image projection and its commercial acceptability for approximately 30 years. The severity of this problem is clearly demonstrated by measuring the contrast sensitivity function of the human eye, which is a measure of the spatial resolution capabilities of the vision system. Comparing it for coherent and incoherent illumination – with and without the presence of speckles – the overall contrast sensitivity decreases by a factor of three for lower spatial frequencies, but by a factor of 20 for a frequency of ~3 cycles/degree. It decreases overall when the pupil size decreases, because a reduction in pupil diameter signifies an increase in speckle size.4 A second problem with speckle noise is encountered in laser video color projection where monochromatic red, green, and blue laser lines are used as primary colors, and other color values are generated by a mixture of the primaries. In each image pixel on the video screen, each laser beam of the different primary colors produces an independent speckle pattern with the typical structure shown in Figure 5b, with a reference of the intensity profile of the direct beam being shown in Figure 5a. Since the intensities of three laser beams are added inconsistently – for instance, for producing a white stimulus – the individual speckle intensity peaks for each color will cause imperfect color summation. Thus, a white laser video image consists of an average white background with a dense mask of colored scintillations of subpixel size. Of course, a critical observer will not accept such an imperfect color mix. But in a display where the image is directly scanned on the retina, this problem can be bypassed, as will be discussed below. A number of methods has been developed to reduce speckle, which arises either from the temporal or spatial coherence of the source.5 For reducing the speckle contrast due to temporal coherence of a laser below the 1% level, its spectral bandwidth has to be increased to 8 nm, which is not possible with most lasers available today. To reduce spatial coherence, the apparent source size has to be increased, which frequently has to be done at the cost of the transmission or resolution of the optical system. The most successful method of reducing speckles in imaging systems, including the human eye, is to move the speckle pattern in time and, applying time averaging over a number of speckle patterns, by taking advantage of the limited integration time of cameras and the inertia of the human eye responsiveness.5 If N is the number of individual changing speckled images during the observation time, the averaging will reduce the speckle contrast by a factor of 1/√N. A speckle pattern can be brought into movement by different means in the optical path, by a moving light diffuser, a vibrating multimode fiber, or an acoustic modulator.6 The safest way to solve the speckle problem is to
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avoid speckle. But this is only possible in imaging systems where there is no difference in the light propagation path from the source to the detecting device, the retina of the eye, or an electro-optical detector. Thus, scattering or diffracting must be omitted in the entire optical path, and only polished optical elements such as lenses and mirrors should be used. This has been realized in the retinal laser display and the scanning laser ophthalmoscope (SLO). In both these systems, the basic principle of confocal scanning is used, where a two-axes mirror scanner moves the beam axis along a raster (xy-)pattern and is in focus on the retina. In the retinal laser display, the video information is modulated on the beam without introducing any path differences over the beam cross-section, and the size of the spot is of the order of the resolution limit of the eye itself, preventing the perception of speckle, which could be induced by the traversing of the beam through the eye medium. With the SLO, the backscattered laser light from the focused beam on the retina is imaged onto the detector aperture, which only accepts light from this tiny flying focused spot, and speckle cannot develop from the overlap of backscattered light waves from different areas of the retina because they are separated in time. Speckle due to the spatial coherence of the source can also be avoided by the correct optical design of the system. In surgical laser applications in the eye, speckle may appear when expanded laser beams are used. In photocoagulation of the retina where a multimode fiber coupler is used between the laser and the ophthalmoscope, the number of different modes travelling through the fiber can interfere as a speckle pattern in the image of the fiber on the retina, leading to an inhomogeneous intensity distribution over an expanded beam spot. But, of course, the internal light scattering in the retinal tissue smooths this modulation. In laser keratotomy, speckle of the various spatial modes of the laser leads to an inhomogeneous intensity distribution across the beam spot applied, which also becomes unnoticed because of the intensity of the smoothing by the scanning process.
Laser speckle optometer Since the speckle pattern of a diffuse object observed with the human eye is always in focus on the retina, irrespective of the refractive errors of the observer, it represents an ideal reference for the emmetropic, ‘rested’ eye, with its far point at infinity. Thus, speckle patterns may be used to examine the state of refraction of the eye. If a diffusing surface illuminated by coherent light far away moves perpendicular to the line of sight of an observer, the speckles may appear to move with respect to the surface. Assuming a normal eye,
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movement in a direction opposite to that of the surface indicates under-accommodation, and movement with the surface indicates too strong an accommodation. If the speckles do not move, but just seem to expand or ‘boil’, the observed surface is optically conjugate to the retina. Since its invention by Knoll in 1966,7 the laser speckle optometer has been found to be rather useful in scientific investigations on the dynamic accommodation processes of the eye, but it never became a commercial diagnostic tool because of its limitation of only being able to measure the refractive state of one monochromatic color.
Bio-speckle flowmetry versus laser Doppler flowmetry Fluorescein angiography is a standard clinical method for the measurement of retinal and choroidal hemodynamics. However, it often only provides qualitative information, and is rather demanding in routine clinical use. Thus, some researchers have studied two alternative optical methods extensively for several years: laser Doppler flowmetry and biospeckle flowmetry. With Doppler flowmetry, the velocity-dependent Doppler frequency shift of scattered laser light by the red blood cells is evaluated. Bio-speckle flowmetry is based on measurement of the time-varying properties of speckles, depending on blood flow. There is a fundamental difference to both approaches. The Doppler shift only appears with movement of the blood cells in a direction parallel to the axis of observation, and vanishes in a flow direction perpendicular to the axis. Since the measuring direction has to be perpendicular to the vessels in the retina, this method is limited to sensing of velocity components of the blood cells perpendicular to the direction of the main blood flow. Its accuracy thus depends on how reliable this velocity component is as a measure of the total flow velocities in the vessels, for different flow velocities and vessel sizes. Contrary to this, the fluctuation of biospeckles is independent of the observation axis. With laser Doppler flowmetry, introduced into ophthalmology in 1972 by Riva et al.,8 the backscattered temporal intensity signal from the detector is transformed into the frequency domain by Fourier transform, resulting in the so-called power spectrum, which contains a frequency shift proportional to the blood cell velocity. The laser Doppler flowmeter has mainly been used for localized measurements in blood vessels, but recently Heidelberg Engineering introduced a scanning laser Doppler flowmeter.9 By means of this technique, laser Doppler measurements are performed in a two-dimensional array of points, resulting in two-dimensionally resolved perfusion maps with a matrix of 256 × 64 pixels and a size of 11.2 × 11.2 µm. With laser Doppler flowmetry
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at a single point, high flow velocities of up to about 100 mm/sec can be detected but, due to the much higher data rates in the scanning device, the highest measurable velocity and signal/ratio are considerably lower, or of the order of 10 mm/sec. With bio-speckle flowmetry, compared to laser Doppler flowmetry, the blood-flow direction and scattering geometry have no direct effect on the speckle fluctuations. Fercher et al.10 have proposed a method for visualizing the retinal blood-flow using single-exposure speckle photography, which provides a blood circulation map. However, since the photographic process is troublesome for clinical use, Aizu et al.11 proposed a method for measuring speckle fluctuation with an electronic detector, obtaining the blood rate by analyzing the power spectrum of the speckle intensity fluctuations in the image plane. Later on, Aizu et al.12 improved their technique by measuring the photon correlation in the diffraction (Fourier) plane. This method evaluates the overall activity of various blood flows existing in the illuminated area, rather than the absolute flow velocity at a certain point in a single vessel. For 2-D mapping of flow velocities by monitoring speckle fluctuations similar to those with the scanning laser Doppler flowmeter – but without the need of scanning – measurements of the image contrast of speckles have been used. The intensity fluctuations of the speckles due to the blood flow blur the speckle pattern, and hence reduce its contrast. With a suitable integration time for the exposure, velocity can be mapped as speckle contrast. Konishi and Fujii13 described a bio-speckle flowmeter known as ‘laser flowgraphy’, where the ‘blur’ value is measured in a hard-wired circuit in a matrix of 100 × 100 pixels, with a pixel size of 25 × 25 µm and a rate of 600 frames/sec. The series of 2-D maps taken at 16 frames/sec clearly visualizes the pulsation of the blood flow. Briers and Xiao-Wei14 described a further development in the data processing for this technique, using a CCD camera and a framegrabber to capture an image of the area of interest. The local speckle contrast is computed and used to produce falsecolor map of velocities. The results are similar to those obtained with the scanning laser Doppler technique, but are obtained without the need of scanning. This reduces the time needed for capturing the image from several minutes to a fraction of a second, already a clinical advantage.
Conclusions The principles of the speckle phenomenon depend upon the wave-nature of light and follow laws of interference of light. The measurement of dynamic speckle patterns has become an important field in speckle flowmetry. With bio-speckle flowmetry, compared to laser Doppler flowmetry, the blood-
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flow direction and scattering geometry have no direct effect upon the speckle fluctuations. Konishi et al. have described a bio-speckle flowmeter known as the ‘laser flowgraph’.
9.
References 10. 1. Newton I: Opticks (reprinted by Dover Press, New York, NY, 1952) Book I, Part I, Prop. VIII, Prob. II., 1730 2. Goodman JW: Statistical Optics. New York, NY: John Wiley & Sons 1985 3. Dainty JC (ed): Laser Speckle and Related Phenomena, Topics in Applied Physics, 9. Berlin/Heidelberg: Springer Verlag 1984 4. Artigas JM, Felipe A, Buades MJ: Contrast sensitivity of the visual system in speckle imagery. J Opt Soc Am 11:23452349, 1994 5. Iwai T, Asakura T: Speckle reduction in coherent information processing. Proc IEEE 84:765-781, 1996 6. Wang L, Tschudi T, Boeddinghaus M, Elbert A, Halldórsson T, Pétursson P: Speckle reduction in laser projections with ultrasonic waves. Opt Eng 39:1659-1664, 2000 7. Knoll HA: Measuring ametropia with a gas laser. Am J Optom Arch Am Optom 43:415-418, 1966 8. Riva CE, Ross B, Benedk GB: Laser Doppler measure-
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11.
12.
13.
14.
ment of blood flow in capillary tubes and retinal arteries. Invest Ophthalmol Vis Sci 11:936-944, 1972 Zinser G: Scanning laser Doppler flowmetry: principle and technique. In: Pillunat LE, Harris A, Anderson DR, Greve EL (eds), Current Concepts on Ocular Blood Flow in Glaucoma. pp. 197-204. The Hague: Kugler Publications 1999 Fercher AF, Peukert M, Roth E: Visualization and measurement of retinal blood flow by means of laser speckle photography. Opt Eng 25:731-735, 1986 Aizu Y, Ogino K, Koyama, T, Takai N, Asakura T: Evaluation of retinal blood flow using time-varying laser speckle. In: Adrian RJ (ed) Laser Anemometry in Fluid Mechanics. III. pp. 55-68. Lisbon: Ladoan 1988 Aizu Y, Ogino K, Sugita T,Yamamoto T, Takai N, Asakura T: Evaluation of blood flow at ocular fundus by using laser speckle. Appl Optics 31:3020-3029, 1992 Konishi N, Fujii H: Real-time visualization of retinal microcirculation by laser flowgraphy. Opt Eng 34:753-757, 1995 Briers JD, Xiao-Wei H: Laser speckle contrast analysis (LASCA) for blood flow visualization: improved image processing. In: Priezzhev AV, Asakura T, Briers JD (eds) Proceedings SPIE, Vol 3252, Optical Diagnostics of Biological Fluids III. pp. 26-33. The International Society for Optical Engineering, Washington DC, 1998
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Laser Doppler techniques in ophthalmology Principles and applications
Charles E. Riva1,2 and Benno L. Petrig1 1 Institut de Recherche en Ophtalmologie, Sion; 2Faculté de Médecine, Université de Lausanne, Lausanne; Switzerland
Keywords: ocular blood flow, Doppler velocimetry, clinical applications, apparatus
Introduction The measurement of blood flow in the ocular fundus is of scientific, as well as of clinical, interest. Its scientific value lies in the possibility of gaining insight into the physiology of deep vascular beds that are under local and central nervous control. Its clinical potential lies in the early assessment of alterations in blood flow, whether associated with specific ocular diseases or resulting from systemic ailments. Furthermore, evaluation of the effect of treatment on the disturbed blood flow represents an important area of application of such a measurement. Ideally, in order to be clinically applicable, blood flow measuring techniques should be reproducible, accurate, and sensitive enough to be able to reveal early pathological alterations. Their spatial resolution should permit measurements at discrete sites of the retinal, optic nerve, and choroidal vascular systems. Their temporal response should be fast enough to allow the investigation of blood flow regulatory responses evoked by various physiological stimuli. The advent of the laser, a device that emits optical waves of almost single frequency, has made it possible to detect, with extremely high resolution, the Doppler shift that light undergoes when scattered by a moving particle. This has allowed the measurement of a broad range of velocities (from µm/s to many km/s). In 1972, this capability led to the first report on the application of the Doppler effect to the measurement of blood velocity, which was obtained from a rabbit retinal arteriole using a helium-neon laser.1 In 1974, Tanaka et al. published the first laser Doppler velocimetry (LDV) measurement of blood in retinal vessels of human
volunteers.2 After the pioneering article by Stern,3 who proposed examining the hemodynamics in the tissue of the skin using LDV, Riva et al.4 described a method to measure blood velocity in the human optic nerve microcirculation. This method was then extended to measuring blood flow by laser Doppler flowmetry (LDF) in the vascular bed of the cat optic nerve head (ONH),5 human ONH6,7 and subfoveal choroid.8 In 1995, Michelson and et al.9 first reported LDF mapping of blood flow in the microcirculation of the human retina by means of a scanning laser ophthalmoscope. Due to the limitations of the space and scope of this review, the present chapter is by no means an exhaustive description of LDV and LDF. Regrettably, numerous important LDV and LDF studies could not be cited. Part of the material in this chapter has been reproduced with permission from previous publications.7,10,11 First, we will describe the LDV technique for measuring blood flow in the main retinal vessels and will discuss some recent developments. Then, the principle of LDF for the measurement of red blood cell (RBC) flux in the tissues of the ONH, subfoveal choroid, and iris will be introduced and applications discussed. The Doppler effect The basis of LDV and LDF is the Doppler effect first described in 1842 by the Austrian physicist Christian Doppler in an article entitled On the Colored Light of Double Stars and Some Other Heavenly Bodies, which describes the frequency shift that a sound or light wave undergoes when emitted from an object that is moving away or towards an observer. It manifests itself, for example,
Address for correspondence: Charles Riva, D.Sc., Institut de Recherche en Ophtalmologie, Grand Champsec 64, CP 4168, CH1950 Sion 4, Switzerland. e-mail:
[email protected] Lasers in Ophthalmology – Basic, Diagnostic and Surgical Aspects, pp. 51–59 edited by F. Fankhauser and S. Kwasniewska © 2003 Kugler Publications, The Hague, The Netherlands
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C.E. Riva and B.L. Petrig which is the so-called Doppler shift. Its magnitude depends upon V, αs and αi, the index of refraction n of the medium containing the particle and the wavelength λ of the laser light in vacuo. For a particle moving at a speed of 1 cm/s, with αs = 80°, αi = 90°, n = 1 (air), and λ = 632.8 nm (heliumneon laser), ∆f is equal to 2.7 × 103 Hz. Although extremely small compared to fi (~5 × 1014 Hz), this shift can be detected using optical mixing spectroscopy.12 Laser Doppler velocimetry measurements of blood velocity in individual retinal vessels
Fig. 1. The Doppler effect. The frequency of the light scattered by the particle (velocity V) in the direction defined by αs is shifted in frequency by an amount ∆f compared to that of the incident light ( fi) from the direction defined by αi (see text for explanation). (Adapted from Riva and Petrig11 by courtesy of the Publisher.)
in the increase in the pitch of the siren of an approaching ambulance. Consider a single particle such as an RBC moving at velocity V in the direction shown in Figure 1. A laser beam of single frequency fi is incident to this particle at an angle αi to the direction of V. The incident light is scattered by the particle in various directions. In the direction of the detector, defined by the angle αs, the frequency of this light will differ from fi by an amount ∆f equal to ∆f = V (cos αs – cos αi) n / λ
Doppler shift power spectrum for RBCs moving in individual retinal vessels Particles suspended in a fluid moving at constant mean velocity through a tube (particles much smaller than the tube diameter) have velocities that depend upon their radial position in the tube. With Poiseuille flow (Fig. 2A), for example, the distribution of velocities as a function of radial position has a parabolic shape, with the peak velocity at the center and zero velocity at the wall. To this distribution of velocities, a spectrum of Doppler shifts corresponds, the so-called Doppler shift power spectrum (DSPS). In our model (Fig. 2B), the DSPS extends from ∆f = 0, which corresponds to the velocity of the particles at the wall of the tube, to ∆f = ∆fmax, where ∆fmax is the frequency shift arising from the particles moving with the centerline velocity, Vmax. ∆fmax is the ‘cutoff frequency’
Fig. 2. Top: Laminar flow in a tube with a parabolic velocity profile (Poiseuille flow). Bottom left: The Doppler shift power spectrum (DSPS) corresponding to this flow. Scattering particles suspended in the fluid are assumed to be uniformly distributed within the cross-section of the tube. The cutoff frequency corresponds to the center line velocity Vmax. The shaded rectangles A, B, and C in the DSPS are the frequencies originating from the particles contained in the corresponding flow annuli (bottom right). (Adapted from Riva and Petrig11 by courtesy of the Publisher.)
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Laser Doppler techniques in ophthalmology of the DSPS. Furthermore, if the particles are uniformly distributed across the tube, it can be shown that the DSPS is constant from ∆f = 0 to ∆f = ∆fmax.13 ∆fmax is measured using the autodyne mode of optical mixing spectroscopy. A description of this technique is beyond the scope of this paper and the interested reader is referred to a previous publication.14 Using bidirectional LDV, an absolute measurement of Vmax can be obtained by detecting the scattered light along two scattering directions, and determining the DSPS for each direction.15 Vmax (cm/s) is derived from the cutoff frequencies ∆fmax1 and ∆fmax2 through the relation Vmax = k (|∆fmax1 – ∆fmax2|). The constant k depends upon the intraocular angle between the two scattering directions, the index of refraction of the ocular media, and the wavelength of the incident laser beam, all of which are known. Retinal blood flow through a retinal vessel is defined as Q = S × Vmean, where S = π D2/4 is the cross sectional area, D the diameter of the vessel and Vmean the mean blood velocity. D is usually measured from fundus photographs taken in monochromatic light at around 570 nm in order to obtain maximum contrast of the blood column relative to the background. The accuracy of Q depends upon that of Vmax and D2, and therefore, it is twice as sensitive to the accuracy to which D can be measured. The relationship between Vmean and Vmax depends on the velocity profile. For a parabolic profile, Vmean = Vmax / 2. Instrumentation A bidirectional retinal LDV device consists of optical systems (i) to deliver a laser beam to a given
53 site on a retinal vessel; (ii) to collect the light scattered by the RBCs along two directions; and (iii) to allow observation of the fundus by an operator. In addition, there is a fixation target for the subject to reduce eye movements. Bi-directional laser delivery and detection systems have been incorporated into a slit-lamp microscope15,16 and into various fundus cameras.17,18 Electronics and signal processing The light detector transforms the incident laser light into an electrical current (photocurrent) signal, which is further processed by electronic filtering and amplification. This resulting so-called ‘Doppler signal’ contains the sum of the Doppler shift components from all moving RBCs. These signal components are separated from each other using fast Fourier analysis, followed by power spectrum estimation. The resulting DSPS is essentially the histogram (distribution) of signal power as a function of Doppler shift frequency (Figs. 2B and 5). Depending on the type of application (LDV or LDF), the DSPS is evaluated by different means of curve fitting models19 and statistical analysis20 in order to derive the various flow parameters described in this review. Applications of laser Doppler velocimetry Physiology Some of the findings obtained in normal volunteers have provided new insight into the physiology of the retinal circulation. For example, they demonstrated high temporal resolution recordings of the pulsatile time course of RBC velocity in retinal vessels (Fig. 3),15,19,21,22 and established the rela-
Fig. 3. Relative RBC center-line velocity (proportional to ∆fmax) in a retinal artery of a human volunteer. A. At normal intraocular pressure (IOP); and B. at IOP elevated to diastolic retinal artery blood pressure. The diastolic velocity is reduced to about zero when IOP is elevated, whereas the systolic velocity is increased above its baseline value, suggesting that the retinal vasculature is counteracting the change in perfusion pressure. Note that the effect of a blink lasts only for a short time. (Reproduced from Petrig and Riva19 by courtesy of the Publisher.)
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tionship of Vmean versus D and of Q versus D.23 Q was found to increase with D2.76 for arteries and D2.84 for veins. This is in good agreement with the values predicted by Murray’s law,24 which stipulates that Q varies as D3 for a vascular system that minimizes its resistance for a given blood volume. The LDV technique is particularly suitable for investigating retinal blood flow regulation in response to various physiological stimuli. These include acute increases in mean ocular perfusion pressure (PPm) achieved using isometric exercises,25 decreases in PPm induced by increasing IOP,26 increases in arterial oxygen and carbon dioxide tension,27 and light/dark transitions.28-31 Clinical applications A number of LDV studies have led to a better understanding of the effect of diabetes on the retina. These examined (i) the retinal circulatory changes during the natural history of diabetes;32,33 (ii) the effect of poor glycemic control on retinal hemodynamics;34 (iii) the response of retinal blood flow to hyperoxia in patients with various degrees of retinopathy;35 and (iv) the effect of various treatment modalities, such as panretinal laser photocoagulation.36,37 Some of the most important LDV findings obtained in diabetic patients have been reviewed by Grunwald and Riva.38 The effect of various antihypertensive medications on the retinal circulation and its autoregulation in normal volunteers and patients with ocular hypertension and primary open-angle glaucoma have been investigated by Grunwald and colleagues.39-41 Recent developments As mentioned above, the relationship between Vmean and Vmax depends upon the velocity profile of the RBCs. In various situations, such as at sharp bends, arteriolar branchings, venous junctions,42 arteriovenous crossings, and impending or partial vessel occlusions, this profile deviates from the parabolic shape.43 Recently, methods have been developed to determine this profile, because a measurement of the RBC velocity profile could be of help in the early diagnosis and treatment of various retinal circulatory impairments. In particular, determination of the velocity gradient at the vessel wall could provide valuable information on the wall shear rate, a quantity that plays an important role in the control of blood flow.44,45 The technique of ‘color Doppler optical coherence tomography’ (CDOCT) allows cross-sectional imaging of blood flow.46 In CDOCT, laser light of low coherence scattered by the RBCs is made to interfere with a strong reference beam from a retroreflector external to the eye. Figure 4A shows a flow profile obtained from a 176-µm retinal vessel using CDOCT.46
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Fig. 4. A. Depth-resolved velocity profile of RBCs in a 176µm diameter vessel in an undilated human eye. The profile was detected and digitized in 10.3 msec and thus does not exhibit any effects of the pulsatile nature of the flow. (Reproduced from Yazdanfar et al.46 by courtesy of the Publisher.) B. Velocity profile obtained from a 150-µm retinal vein of a human eye by scanning a 670-nm coherent laser beam across a straight portion of the vessel.48
Using lasers with different coherence lengths to obtain Doppler shifts from RBCs moving in various volumes of increasing depths from the vessel wall provides another approach to monitor in vivo the velocity gradient at the vessel wall.47 With this ‘variable coherence optical Doppler velocimetry’ (VCODV) only the light from RBCs moving at a depth of less than half the coherence length is efficiently detected. Yet another technique for measuring the RBC velocity profile consists of scanning a 12-µm diameter coherent laser beam perpendicularly across retinal vessels.48 Such scans have demonstrated a parabolic shape of the velocity profile in straight portions of retinal vessels (Fig. 4B).
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Laser Doppler techniques in ophthalmology Laser Doppler flowmetry measurements of blood flow in the optic nerve head, subfoveal choroid and iris The Doppler shift power spectrum in the case of red blood cells moving in the microvascular bed of a tissue When a laser beam illuminates RBCs moving through a network of capillaries at various velocities and in different directions, the light scattered by the RBCs consists of a summation of waves with various Doppler shifts ∆fi. It can be generally noted that: (i) Most of the light emerging from the tissue has been scattered solely by static structural components of the tissue. This non-shifted light acts as a reference signal that is mixed with the shifted scattered light at the surface of the photodetector. The photocurrent contains only the components oscillating at the Doppler shift frequencies (∆fi). (ii) The light scattered by the RBCs contains Doppler shift frequencies that can be positive or negative, depending on the direction of RBC motion relative to the incident light and detector. Since the detector does not discriminate between positive and negative frequency shifts, the DSPS only spans the positive frequencies (Fig. 5).
Fig. 5. DSPS of the optic nerve tissue. Arrow indicates the mean frequency, which is proportional to the mean velocity of the RBCs. (Reproduced from Riva and Petrig11 by courtesy of the Publisher.)
Hemodynamic parameters derived from the Doppler shift power spectrum Assuming the validity of Bonner and Nossal’s theory,49 the following flow parameters are obtained from the DSPS. Vel is the mean speed of the RBCs moving in the sampling volume (proportional to the mean Doppler frequency shift). Vol is the number of moving RBCs in the sampling volume (proportional to the area under the DSPS curve). The total RBC flux in the sampling volume is F = Vel × Vol, where Vel is expressed in Hz, and Vol and F in arbitrary units.
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55 This means that the LDF technique provides only relative blood flow measurements for the following reasons. Laser radiation upon a tissue will undergo scattering and absorption, both influencing the penetration pattern of the laser light. Penetration may differ from one region of a tissue to another, depending upon the optical properties of the tissue. Thus, spatial or temporal variations in tissue structure and vascularization, as is the case, for example, in the ONH in glaucoma, will affect the LDF measurements. Furthermore, direct comparison between the flow values from different tissues may not be valid due to variations in optical properties resulting from differences in tissue structure and composition. The measured quantity F is usually referred to as ‘blood flow’. However, what is actually measured is the flux of the RBCs. Blood flow is only directly proportional to RBC flux if the hematocrit remains constant during an experiment. Laser Doppler flowmetry measurement modes Two LDF measurement modes have been implemented. The first is the continuous mode for online, continuous recording of the flow parameters at a discrete site of the vascular beds of the ONH, subfoveal region of the choroid, or iris. The second mode is based on a scanning laser technique and provides a two-dimensional image of the RBC flux in the capillaries of the ONH and retina. Continuous recording of the laser Doppler flowmetry signal in the optic nerve head and subfoveal choroid In ONH LDF, as originally described by Riva et al.,5 an optical system adapted to a standard fundus camera delivers a laser beam to a discrete site on the optic disc. The scattered light is collected by an optical fiber placed in the retinal plane of the camera and guided to a photodetector. An area of the fundus (30° in diameter) is illuminated in red-free light, allowing the observation and positioning of the laser beam at the disc, away from visible blood vessels. The effective diameter at the disc from which the scattered light is collected is approximately 150 µm. A NeXT computer system with dedicated software is used for the LDF analysis.20 This software allows averaging of the Doppler signal in phase with the heart cycle so that precise measurements of RBC flux variations during the systolic and diastolic phases can be obtained. Recordings of Vel, Vol and F from the ONH of a normal volunteer are shown in Figure 6. Scanning laser Doppler flowmetry Scanning laser Doppler flowmetry (SLDF) combines the techniques of LDF and scanning laser ophthalmoscopy.50,51 It is based on the principle of
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Fig. 6. Recording of the LDF flow parameters from the optic nerve of a human volunteer. The diameter of the beam at the optic disc was approximately 150 µm. On the right, the time course of these parameters during two heart cycles was obtained by averaging the recordings over 20 seconds (shaded area) in phase with the heart cycle.
confocal microscopy and has a nominal depth resolution of approximately 300 µm. The Heidelberg Retinal Flowmeter (HRF) performs quick measurements at 16,384 different locations in a two-dimensional grid, and provides an image of the retinal perfusion. The perfusion map can be used for a qualitative visualization of the network of perfused vessels and capillaries. It is also possible to define a measurement region of variable size (a 100 x 100 µm window is commonly used), to place it interactively anywhere in the perfusion map, and to measure the average perfusion values inside this region. This technique is limited to the detection of Doppler shifts below 2000 Hz. Laser Doppler flowmetry sample volume A central question in the application of LDF to the ONH is the depth of the sampling volume. This depth determines the relative contribution to the Doppler signal of the superficial layers, those supplied by the central retinal artery, and the deeper layers supplied by the posterior ciliary arteries. These two vascular beds may have different blood flow regulation properties. Furthermore, the deep layers of the ONH appear to be particularly susceptible to ischemic disorders, including glaucoma. Investigations on a model system suggests that, when the light-collecting aperture coincides with the tissue volume illuminated by the probing laser, layers of the ONH tissue of as deep as 300 µm contribute to the LDF signal.52 In humans, however, although the depth of tissue sampling in the ONH remains to be experimentally assessed, it appears that the LDF technique detects predominantly the motion of RBCs within the intraocular region of the ONH.53 A study on monkey eyes54 concluded that LDF is predominantly sensitive to blood flow changes in the superficial layers of the ONH, and less to those layers of the prelaminar and deeper regions of the ONH, and that their relative proportions are still unknown. The weaker signal from the deep layers cannot be separated from the dominant signal from the superficial layers to exclusively study the circulation in the deep layers.
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Linearity of laser Doppler flowmetry Linearity between F and actual blood flow has been documented for various tissues, such as the skin, skeletal muscle, cerebral cortex, nerves, etc.55 Experiments have shown that the assumption of LDF linearity is valid for the ONH and choroidal circulations.8,11 Applications of laser Doppler flowmetry The high spatial and temporal resolution of the LDF technique, particularly in the continuous mode, makes this technique most suitable for investigation of the regulatory processes of blood flow in response to various physiological stimuli. Optic nerve head LDF investigations of ONH blood flow in humans include the response of this flow to: (i) decreases in PPm induced by increases in IOP;53,56 (ii) increases in PPm produced by increases in systemic blood pressure with isometric exercises;57 (iii) hyperoxia, breathing carbogen and mixtures of O2 and CO2;58-60 and (iv) increased neuronal activity (Fig. 7).15,61 LDF and SLDF have been used to measure retinal and ONH microvascular perfusion in normal, ocular hypertensive, and glaucomatous eyes,62-66 under various clinical conditions and using different provocation tests. Other clinical investigations include the study of retinal vein occlusions67 and retinal blood flow in age-related macular degeneration.51,68 Choroid and iris LDF measurements of choroidal blood flow in the foveal region of the human fundus are recent. Studies have been performed in humans on the effect of: (i) increases and decreases of PPm;69,70 (ii) valsalva maneuvers;8 (iii) breathing of various gas mixtures (pure O2, various mixtures of O2 and
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Fig. 7. Optic nerve blood flow response to a 15-Hz diffuse luminance flicker (horizontal bar) in a normal volunteer.
CO2);71 and (iv) the effect of light.72 Investigations on the effect of aging, age-related macular degeneration, and choroidal neovascularization have been reported.73,74 LDF has been applied to investigate iris blood flow in humans, as well as the effect of acute physical exercise75 and increased IOP76 on this flow.
7.
Conclusions
9.
The LDV and LDF techniques have been applied over a number of years. They are powerful methods for noninvasively investigating changes in blood flow in the retina, ONH, subfoveal choroid, and iris. Both techniques have high sensitivity and a temporal response fast enough to reveal the changes in blood flow during the cardiac cycle, and in response to acute changes in various physiological stimuli. These capabilities open new avenues in the understanding of the regulation of blood flow in the various tissues of the eye.
8.
10.
11.
12.
Acknowledgments 13.
The authors wish to thank Martial Geiser for his help in preparing the manuscript and Eric Logan for providing figure 4B. This study was supported in part by the Swiss National Science Foundation (Grant No. 32-43157), the Priority Program ‘Optics II’ (Grant No. 491), the EMDO Foundation, Ciba Vision Ophthalmics, the Mobilière Suisse, and the Loterie Suisse Romande.
14.
15.
16.
References 1. Riva CE, Ross B, Benedek GB: Laser Doppler measurements of blood flow in capillary tubes and retinal arteries. Invest Ophthalmol 11:936-944, 1972 2. Tanaka T, Riva CE, Ben-Sira I: Blood velocity measurements in human retinal vessels. Science 186:830-831, 1974 3. Stern MD: In vivo evaluation of microcirculation by coherent light scattering. Nature 254:56-58, 1975 4. Riva CE, Grunwald JE, Sinclair SH: Laser Doppler measurement of relative blood velocity in the human optic nerve head. Invest Ophthalmol Vis Sci 22:241-248, 1982 5. Riva CE, Harino S, Petrig BL, Shonat RD: Laser-Doppler flowmetry in the optic nerve. Exp Eye Res 55:499-506, 1992 6. Riva CE, Falsini B, Geiser MH, Petrig BL: Optic nerve head blood flow response to flicker. In: Pillunat LE, Har-
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metry, pp 349-383. Boston: Kluwer Academic Publ 1989 22. Petrig BL, Riva CE: Near-infrared retinal laser Doppler velocimetry and flowmetry: new delivery and detection techniques. Appl Opt 30:2073-2078, 1991 23. Riva CE, Grunwald JE, Sinclair SH, Petrig BL: Blood velocity and volumetric flow rate in human retinal vessels. Invest Ophthalmol Vis Sci 26:1124-1132, 1985 24. Murray CD: The physiological principle of minimum work. I. The vascular system and the cost of blood volume. Proc Nat Acad Sci US 12:207-214, 1926 25. Robinson F, Riva CE, Grunwald JE, Petrig BL, Sinclair SH: Retinal blood flow autoregulation in response to an acute increase in blood pressure. Invest Ophthalmol Vis Sci 27:722-726, 1986 26. Riva CE, Sinclair SH, Grunwald JE: Autoregulation of retinal circulation in response to decrease of perfusion pressure. Invest Ophthalmol Vis Sci 21:34-38, 1981 27. Pakola SJ, Grunwald JE: Effects of oxygen and carbon dioxide on human retinal circulation. Invest Ophthalmol Vis Sci 34:2866-2870, 1993 28. Feke GT, Zuckerman R, Green GJ, Weiter JJ: Responses of human retinal blood flow to light and dark. Invest Ophthalmol Vis Sci 24:136-141, 1983 29. Riva CE, Grunwald JE, Petrig BL: Reactivity of the human retinal circulation to darkness: a laser Doppler velocimetry study. Invest Ophthalmol Vis Sci 24:737-740, 1983 30. Riva CE, Logean E, Petrig BL, Falsini B: Effet de l’adaptation à l’obscurité sur le flux rétinien. Klin Mbl Augenheilk 216:309-310, 2000 31. Riva CE, Petrig BL, Grunwald JE: Near infrared retinal laser Doppler velocimetry. Lasers Ophthalmol 1:211-215, 1987 32. Feke GT, Tagawa H, Yoshida A, Goger DG, Weiter JJ, Buzney SM, McMeel JW: Retinal circulatory changes related to retinopathy progression in insulin-dependent diabetes mellitus. Ophthalmology 92:1517-1522, 1985 33. Grunwald JE, Riva CE, Sinclair SH, Brucker AJ, Petrig BL: Laser Doppler velocimetry study of retinal circulation in diabetes mellitus. Arch Ophthalmol 104:991-996, 1986 34. Grunwald JE, Riva CE, Baine J, Brucker AJ: Total retinal volumetric blood flow rate in diabetic patients with poor glycemic control. Invest Ophthalmol Vis Sci 33:356-363, 1992 35. Grunwald JE, Riva CE, Brucker AJ, Sinclair SH, Petrig BL: Altered retinal vascular response to 100% oxygen breathing in diabetes mellitus. Ophthalmology 91:14471452, 1984 36. Grunwald JE, Riva CE, Brucker AJ, Sinclair SH, Petrig BL: Effect of panretinal photocoagulation on retinal blood flow in proliferative diabetic retinopathy. Ophthalmology 93:590-595, 1986 37. Feke GT, Green GL, Goger DG, McMeel JW: Laser Doppler measurements of the effect of panretinal photocoagulation on retinal blood flow. Ophthalmology 89:757-762, 1982 38. Grunwald JE, Riva CE: Retinal blood flow in diabetes. In: Belcaro GV, Hoffmann U, Bollinger A, Nicolaides AN (eds) Laser Doppler, pp 223-247. London: Med-Orion Publ Co 1994 39. Grunwald JE: Effect of two weeks of timolol maleate treatment on the normal retinal circulation. Invest Ophthalmol Vis Sci 32:39-45, 1991 40. Grunwald JE, Delehanty J: Effect of topical carteolol on the normal human retinal circulation. Invest Ophthalmol Vis Sci 33:1853-1856, 1992 41. Grunwald JE, Mathur S, Dupont J: Effects of dorzolamide hydrochloride 2% on the retinal circulation. Acta Ophthalmol Scand 75:236-238, 1997
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42. Pries AR, Secomb TW, Gaehtgens P: Biophysical aspects of blood flow in the microvasculature. Cardiovasc Res 32:654-667, 1996 43. Fankhauser F, Bebie H, Kwasniewska S: The influence of mechanical forces and flow mechanisms on vessel occlusion. Lasers Surg Med 6:530-532, 1987 44. Koller A, Kaley G: Endothelial regulation of wall shear stress and blood flow in skeletal muscle microcirculation. Am J Physiol 260:H862-H868, 1991 45. Nerem RM, Alexander RW, Chappell DC, Medford RM, Varner SE, Taylor WR: The study of the influence of flow on vascular endothelial biology. Am J Med Sci 316:169175, 1998 46. Yazdanfar S, Rollins AM, Izatt JA: Imaging and velocimetry of the human retinal circulation with color Doppler optical coherence tomography. Opt Lett 25:1448-1450, 2000 47. Logean E, Schmetterer LF, Riva CE: Optical Doppler velocimetry at various retinal vessel depths by variation of the source coherence length. Appl Opt 39:2858-2862, 2000 48. Logean E, Schmetterer LF, Dorsaz JR, Riva CE: Red blood cell velocity profile in retinal vessels by confocal laser Doppler velocity. Invest Ophthalmol Vis Sci 42:S238, 2001 49. Bonner R, Nossal R: Model for laser Doppler measurements of blood flow in tissue. Appl Opt 20:2097-2107, 1981 50. Michelson G, Schmauss B: Two dimensional mapping of the perfusion of the retina and optic nerve head. Br J Ophthalmol 79:1126-1132, 1995 51. Zinser G: Scanning laser Doppler flowmetry. In: Pillunat LE, Harris A, Anderson DR, Greve EL (eds) Current Concepts on Ocular Blood Flow in Glaucoma, pp 197-204. The Hague: Kugler Publ 1999 52. Koelle JS, Riva CE, Petrig BL, Cranstoun SD: Depth of tissue sampling in the optic nerve head using laser Doppler flowmetry. Lasers Med Sci 8:49-54, 1993 53. Riva CE, Hero M, Titzé P, Petrig BL: Autoregulation of human optic nerve head blood flow in response to acute changes in ocular perfusion pressure. Graefe’s Arch Clin Exp Ophthalmol 235:618-626, 1997 54. Petrig BL, Riva CE, Hayreh SS: Laser Doppler flowmetry and optic nerve head blood flow. Am J Ophthalmol 127:413425, 1999 55. Shepherd AP, Öberg PA: Laser-Doppler blood fowmetry. Boston: Kluwer Academic Publ 1990 56. Pillunat LE, Stodtmeister R, Wilmanns I, Christ T: Autoregulation of ocular blood flow during changes in intraocular pressure. Graefe’s Arch Clin Exp Ophthalmol 223:219223, 1985 57. Movaffaghy A, Chamot SR, Petrig BL, Riva CE: Blood flow in the human optic nerve head during isometric exercise. Exp Eye Res 67:561-568, 1998 58. Geiser MH, Riva CE, Dorner GT, Diermann U, Luksch A, Schmetterer L: Response of choroidal blood flow in the foveal region to hyperoxia and hyperoxia-hypercapnia. Curr Eye Res 21(2):669-676, 2000 59. Harris A, Douglas RA, Pillunat L, Joos K, Knighton RW, Kagemann L, Martin BJ: Laser Doppler flowmetry measurement of changes in human optic nerve head blood flow in response to blood gas perturbations. J Glaucoma 5:258265, 1996 60. Lietz A, Hendrickson P, Flammer J, Orgül S, Haefliger IO: Effect of carbogen, oxygen and intraocular pressure on Heidelberg retina flowmeter parameter ‘Flow’ measured at the papilla. Ophthalmologica 212:149-152, 1998 61. Riva CE, Falsini B, Logean E: Flicker-evoked response of human optic nerve head blood flow: luminance versus chromatic modulation. Invest Ophthalmol Vis Sci 42:756-762, 2001
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Laser Doppler techniques in ophthalmology 62. Grunwald JE, Piltz J, Hariprasad SM, Dupont J: Optic nerve and choroidal circulation in glaucoma. Invest Ophthalmol Vis Sci 39:2329-2336, 1998 63. Nicolela MT, Hnik P, Drance SM: Scanning laser Doppler flowmeter study of retinal and optic disk blood flow in glaucomatous patients. Am J Ophthalmol 122:775-783, 1996 64. Michelson G, Langhans MJ, Groh MJM: Perfusion of the juxtapapillary retina and the neuroretinal rim area in primary open angle glaucoma. J Glaucoma 5:91-98, 1996 65. Wolf S, Arend O, Sponsel WE, Schulte K, Cantor LB, Reim M: Retinal hemodynamics using scanning laser ophthalmoscopy and haemorheology in chronic open-angle glaucoma. Ophthalmology 100:1561-1566, 1993 66. Strenn K, Matulla B, Wolzt M, Findl O, Bekes MC, Lamsfuss U, Graselli U, Rainer G, Menapace R, Eichler HG, Schmetterer L: Reversal of endothelin-1-induced ocular hemodynamic effects by low-dose nifedipine in humans. Clin Pharmacol Ther 63:54-63, 1998 67. Avila CP, Bartsch DU, Bitner DG, Cheng L, Mueller AJ, Karavellas MP, Freeman WR: Retinal blood flow measurements in branch retinal vein occlusion using scanning laser Doppler flowmetry. Am J Ophthalmol 126:683-690, 1998 68. Kruger A, Matulla B, Wolzt M, Pieh S, Strenn K, Findl O, Eichler HG, Schmetterer L: Short-term oral pentoxifylline use increases choroidal blood flow in patients with agerelated macular degeneration. Arch Ophthalmol 116:2730, 1998
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59 69. Riva CE, Titzé P, Hero M, Movaffaghy A: Choroidal blood flow during isometric exercises. Invest Ophthalmol Vis Sci 38:2338-2343, 1997 70. Riva CE, Titzé P, Petrig BL: Effect of acute decrease of perfusion pressure on choroidal blood flow in humans. Invest Ophthalmol Vis Sci 38:1752-1760, 2002 71. Schmetterer L, Wolzt M, Lexer F, Alschinger C, Gouya G, Zanaschka G, Fassolt A, Eichler HG, Fercher Friedrich A: The effect of hyperoxia and hypercapnia on fundus pulsations in the macular and optic disc region in healthy young men. Exp Eye Res 61:685-690, 1995 72. Longo A, Geiser M, Riva CE: Subfoveal choroidal blood flow in response to light-dark exposure. Invest Ophthalmol Vis Sci 41(9):2678-2683, 2000 73. Grunwald JE, Hariprasad SM, Dupont J, Maguire MG, Fine SL, Brucker AJ, Maguire AM, Ho AC: Foveolar choroidal blood flow in age-related macular degeneration. Invest Ophthalmol Vis Sci 39:385-390, 1998 74. Grunwald JE, Hariprasad SM, Dupont J: Effect of aging on foveolar choroidal circulation. Arch Ophthalmol 116:150154, 1998 75. Michelson G, Groh M, Gründler A: Regulation of ocular blood flow during increases of arterial blood pressure. Br J Ophthalmol 78:461-465, 1994 76. Chamot SR, Movaffaghy A, Petrig BL, Riva CE: Effet d’une diminution de la pression de perfusion oculaire sur le flux sanguin dans l’iris mesuré par fluxmétrie laser Doppler. Klin Mbl Augenheilk 214:302-304, 1999
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Principles of optical coherence tomography Christoph K. Hitzenberger Institute of Medical Physics, University of Vienna, Vienna, Austria
Keywords: dual beam OCT, en-face OCT, Doppler OCT, clinical results, interferometry
Introduction In modern ophthalmology, there is an increasing demand for high-resolution imaging of the ocular tissues. In particular, the diagnosis of retinal disorders has been dramatically improved by the introduction of modern imaging techniques such as scanning laser ophthalmoscopy, fluorescein angiography, and nerve fiber polarimetry. While these techniques are able to image retinal structures at a high transverse resolution, depth resolution is limited by the numerical aperture of the eye and by aberrations of its optical components to about 300 µm. This prevents the use of these technologies for measuring and imaging the depth of individual fundus layers (an exception is nerve fiber polarimetry, which can determine the thickness of the retinal nerve fiber layer by measuring its birefringence). Optical coherence tomography (OCT) has overcome this limitation. Based on the special coherence properties of superluminescent diodes (SLDs) and femtosecond pulse lasers, OCT enables the recording of cross-sectional images of transparent and semi-transparent tissues at a high resolution.1,2 The main application field of OCT is retinal imaging,3-6 while another ophthalmic application field is imaging of the anterior eye segment.7 OCT imaging is similar to ultrasound B-mode imaging, except that near infrared light is used instead of sound to probe the tissue. OCT images are made up of several optical A scans (in analogy to ultrasound A scans). The basic ranging technique of OCT that provides the optical A scans is partial (or low) coherence interferometry (PCI or LCI), a technique developed after the mid-1980s for measurement of intraocular distances.8-12 PCI enables high-depth resolution that is completely indepen-
dent of the numerical aperture and the aberrations of the eye. Instead, the axial resolution solely depends on the coherence properties of the probing light beam. In standard applications, an SLD with a bandwidth of ~25 nm yields an axial resolution of ~10 µm. This provides OCT images whose resolution is at least one order of magnitude better than that of ultrasound B-mode imaging. By using stateof-the-art femtosecond pulse lasers as the light source, a depth resolution of ~3 µm can be achieved.13 Another advantage of OCT, compared to ultrasound imaging, is that it operates as a non-contact technique and is therefore much more comfortable for the patient and avoids any risk of corneal infection. This chapter describes the basic principles of OCT, the generation of optical A scans, the synthesis of A scans to an OCT image, and discusses axial and transversal resolution, image contrast, and the advantages and limitations of the technique. Alternative methods and extensions, such as dual beam OCT, transversal (or en-face) OCT, and Doppler OCT, are discussed briefly, and finally, an outlook regarding future developments is presented.
Basics of optical coherence tomography Partial coherence interferometry Optical A scans are generated by PCI, the basic ranging technology of OCT. The goal of PCI is, very precisely, to locate the depth position of the back-reflecting or backscattering layers within a sample. Contrary to classical interferometry, which uses laser light of high spatial and temporal coher-
Address for correspondence: Prof. Christoph K. Hitzenberger, Institute of Medical Physics, University of Vienna, Währinger Strasse 13, A-1090 Vienna, Austria. email:
[email protected] Lasers in Ophthalmology – Basic, Diagnostic and Surgical Aspects, pp. 61–72 edited by F. Fankhauser and S. Kwasniewska © 2003 Kugler Publications, The Hague, The Netherlands
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Fig. 1. Basic principle of partial coherence interferometry.
ence, PCI needs a light source that is only partly coherent. It has to have high spatial coherence, but very low temporal coherence (or short coherence length lc).14 Roughly speaking, high space coherence means that all light components across the beam cross-section are in phase with each other, i.e., any parts of the beam that have travelled the same distance from the source can interfere with each other. Short coherence length means that only those beam components that have travelled approximately the same distance from the light source can interfere with each other (components whose travel distances from the light source differ by more than lc cannot interfere). Figure 1 illustrates the basic principle of PCI. A low coherence light source (usually an SLD) emits a short coherence light beam towards a Michelson interferometer. At the beam splitter, the beam is divided into two components: the reference beam directed towards the movable reference mirror, where it is retroreflected, and the sample beam directed towards the sample. This beam is backreflected or backscattered, usually at several interfaces within the sample whose positions are to be measured. Both beam components are recombined at the beam splitter and superimposed on the photodetector. In order to determine the positions of the sample’s interfaces, the reference mirror is moved, while the light intensity is measured as a function of the position of the mirror. The signal is recorded on a PC and displayed on the computer monitor after the measurement. Any time the path difference from the beam splitter to the reference mirror equals one of the path differences from the beam splitter to one of the sample interfaces, interferometric light modulation occurs: an oscillating signal intensity is recorded. The frequency of the oscillations equals the Doppler shift frequency of the reference light caused by the moving reference mirror, and acts as the carrier frequency of the coherence signal. At the position of maximum os-
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cillation strength, the reference mirror position corresponds to the position of the sample interface (see Fig. 1). It should be mentioned that distances measured by this method are optical ones, i.e., geometric distances multiplied by the (group) refractive index of the sample.9 In order to obtain the true geometric distances, the optical distances have to be divided by that index. Usually, only the envelope of the interferometric signal is recorded. This saves computer memory and improves the readability of the signal. The position of a sample interface is now indicated by a signal peak in the A scan. The width of the peak is equal to the ‘round trip’ coherence length (lc/2, because the beam travels back and forth in the interferometer), and determines the depth resolution of the technique. Figure 2 illustrates the application of PCI for measuring intraocular distances. The bulk Michelson interferometer is now replaced by a fiber optic Michelson interferometer, as is used in the majority of present OCT set ups.1,4-7 The advantages of the fiber optic set up are its compact, light-weight design, its robustness, and ease of alignment. The beam splitter is replaced by a fiber optic 50:50 X-coupler, and the reference mirror by a retroreflector. An SLD and a photodetector are directly coupled to the interferometer via single mode fibers; therefore, these components need no further alignment. The eye is placed in the sample path and, as indicated in Figure 2, a signal peak is recorded every time the retroreflector position coincides with an intraocular interface. In this way, intraocular distances can be determined with micrometer precision.9-12,15
Synthesis of optical coherence tomograms In order to obtain cross-sectional OCT images, several optical A scans are recorded at adjacent sample
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Fig. 2. Partial coherence interferometry for intraocular ranging. Principle of optical A-scan recording.
positions. For this purpose, a pair of scanning mirrors deflects the probing beam to different sample positions and an A scan is recorded at each position (see Fig. 3). The signal intensities are converted to gray levels or color values, and the individual scans are mounted to form a two-dimensional gray level or false color image showing a cross-section of the sample, the OCT tomogram. In conventional OCT of the retina, realized by the first commercially available system, the ZeissHumphrey OCT scanner, 100 A scans are typically recorded within ~1 second. The transversal orientation of the tomographic section can be arbitrarily chosen by tilting the scanning mirrors between the individual A scans in a suitable, predetermined way. Thereby, horizontal, vertical, arbitrarily oriented, and even circular sections (e.g., around the optic nerve head) can be recorded. Image contrast Conventional OCT is based on the intensity backreflected or backscattered at object structures. Backreflection occurs at interfaces within the sample where the refractive index changes abruptly. However, such an interface is only observable in OCT images if the light is retroreflected, i.e., if the interface is oriented (nearly) perpendicular to the probing light beam. In this case, the interface shows up as a bright line. Backscattering is caused by particles having a refractive index different from that of the surrounding matrix. Backscattering occurs in all directions, however, the intensity will vary with backscattering angle, depending on particle size, shape, and refractive index mismatch to
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Fig. 3. Synthesis of an OCT tomogram of the retina by recording and mounting several A scans.
the matrix.16 Therefore, backscattering structures are visible in OCT images independently of the orientation of the backscattering structure. Backscattering surfaces show up as bright boundaries, volume backscatterers (particles dispersed in a surrounding matrix) are imaged as ‘speckled’ structures; due to attenuation by scattering, the backscattered intensity decreases exponentially with depth. Absorption is another effect that attenuates light remitted by the sample. Strongly-absorbing material can shadow tissue structures beneath, and render them invisible. It should be mentioned that OCT is a ‘high pass’ technique.2 This means that only areas where the refractive index changes over a very short distance (of the order of a wavelength or less) give rise to detectable backscattered (or back–reflected) intensity. Areas with only slowly spatially varying refractive indexes remain invisible in standard OCT (enhanced OCT contrasting techniques, e.g., polarization-sensitive OCT or phase-contrast OCT, which are currently under development, can improve this situation in certain cases). As already mentioned, scattering (and also absorption) attenuates the probing light exponentially with depth. Furthermore, the back-reflection or backscattering coefficient can vary by a large amount, depending on the optical properties of the tissue structures. This causes a large variation in backscattered intensity, usually too large to be
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Fig. 4. OCT image recorded in the retina of a human eye in vivo with a Zeiss-Humphrey OCT scanner. A horizontal cross-section through the optic nerve head is shown. The direction towards the fovea is indicated. RPE: retinal pigment epithelium.
displayed on a linear color or gray level scale. Therefore, a logarithmic intensity scale is usually adopted in OCT. Figure 4 shows an OCT tomogram recorded horizontally through the optic nerve head of a human eye in vivo. It illustrates the image contrast typically obtained from retinal structures: a strong backscattered intensity is obtained from the retinal nerve fiber layer and from the posterior surface of the retina, probably caused by the retinal pigment epithelium (RPE) and the choriocapillaris. Intermediate layers show weaker backscattering with a typical ‘speckled’ appearance. Due to light attenuation, only very weak light intensity can be observed from structures below the intense RPE/ choriocapillaris band. Structures with a rather constant refractive index, such as the vitreous or edema (not shown here), do not backscatter light and are displayed in black. Blood absorbs the wavelength used for retinal OCT imaging (~800–850 nm). Therefore, blood vessels can shadow structures beneath them. System performance Light source The most important component of an OCT system with respect to its performance is the light source. Via absorption and scattering coefficients, the central wavelength determines how deep the light penetrates into the tissue. The bandwidth defines the depth resolution, and the power determines the sensitivity of the system. OCT light sources have to fulfill very special criteria: good spatial coherence is required, while the temporal coherence has to be very low (short coherence length). Three types of light sources that fulfill these criteria are presently in use for OCT
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applications: femtosecond pulse lasers, broadband fiber optic sources, and superluminescent diodes. Femtosecond pulse lasers achieve the shortest coherence lengths and therefore the best axial resolution.13,17 However, they are very expensive and difficult to operate. Therefore, their use for OCT applications is presently restricted to a very few research laboratories. Broadband fiber optic sources are derived from the fiber optic amplifiers used in the telecommunications industry. Because of their origin, they are presently restricted to the wavelengths most commonly used in telecommunications: wavelength bands centered around 1300 and 1550 nm. The advantage of these sources is that they provide rather a high output power (up to ~30 mW) from a single mode fiber, which ensures excellent spatial coherence. This type of source is preferably used for OCT of scattering tissue, such as skin, teeth, mucosa, etc., because the longer wavelengths of 1.3–1.55 µm are less scattered and allow greater penetration depth (~1–2 mm) into these types of tissue. However, these wavelengths cannot be used for retinal OCT because they are absorbed by the water contained in the ocular media. An ophthalmic application for 1300 nm light is OCT of the anterior eye segment,7 where the better transmission of this wavelength is used to image the sclera and iris. SLDs are most commonly used in OCT. They are comparatively cheap, lightweight, easy to operate, and available in a comparatively large selection of wavelengths. Available center wavelengths presently range from 670-1550 nm. The wavelength range used for retinal OCT is ~800–850 nm. In this range, output powers from ~0.5-20 mW are available from single mode fibers. The ocular media are usually perfectly transparent in this wavelength range, and the retina tolerates more light intensity than at the shorter wavelength of 670 nm.
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With a probing beam power of 750 µW at λ0 = 0.83 µm, the maximum permissible exposure time of a point at the retina is eight seconds18. Within this time, several OCT images can be recorded. Furthermore, since the beam scans the retina (i.e., the light energy is distributed over several points of the retina), OCT imaging is well below laser safety limits. Resolution in OCT One of the main advantages of OCT is the complete decoupling of axial (depth) resolution from transverse resolution. Contrary to scanning laser ophthalmoscopy, axial resolution is influenced by neither the numerical aperture of the eye nor aberrations of its optical components. The axial resolution of OCT is determined by the coherence length lc of the probing light. For a light source with a Gaussian emission spectrum with emission bandwidth ∆λ (full width at half maximum, FWHM) and center wavelength λ0, the axial resolution in air, for a dispersion balanced interferometer, is given by the following equation:19
δx = 2ln2
2ln2 λ20 δz = . π ∆λ
fλ πd
(3)
(1)
A single reflecting interface in the sample arm gives rise to a coherence signal with an FWHM width δz. δz equals the so-called round trip coherence length which is equal to lc/2. If we assume a center wavelength of λ0 h 830 nm for retinal OCT (compare Light source above), the main parameter determining the axial resolution is ∆λ. The larger the bandwidth, the better the resolution. For a typical SLD bandwidth of 25 nm, we obtain δz h 12 µm in air. For OCT imaging within a sample with a refractive index of n > 1, the resolution is improved: δz δzMedium = n .
(2)
(Strictly speaking, we would have to use the group refractive index ng instead of the usual phase refractive index n9. However, for the ocular media we are interested in, the difference between n and ng is only in the order of ~1%.) If we assume an index of n h 1.3520 for ocular media, we obtain δzMedium h 9 µm. It should be mentioned that the resolution figures provided by equations (1) and (2) are only valid if the interferometer is dispersion-compensated. In other words, the amount of (group) dispersion introduced by dispersive media (lenses, prisms, glass fibers, etc.) in the sample arm has to be compensated by an equal amount of dispersion by equivalent media in the reference arm. Otherwise the resolution is degraded. This is not usually a problem for the technical components in the interfero-
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65
where f is the focal length of the lens (in the case of retinal imaging, the focal length of the optical components of the eye), and d the FWHM beam diameter of the collimated beam. With a beam diameter of about 1–2 mm, a transverse resolution of the order of ~10 µm can be expected. This is equal to the transverse resolution obtained with scanning laser ophthalmoscopes. Beam diameters larger than ~3–4 mm would not further improve the transverse resolution, because of the aberrations caused by the ocular optical components. Furthermore, a larger beam diameter would reduce the depth of focus considerably, requiring dynamic focusing,24 which would make the system more complex and expensive. It should be mentioned that the transverse resolution defined by equation (3) requires that the transverse sampling rate is high enough, i.e., optical A scans are recorded at transverse sampling intervals ≤ δx. However, this is usually not the case in retinal OCT. A typical OCT image of the retina consists of 100 A scans equally spaced over a transverse scanning width of ~3 mm. In this case, the transverse resolution is equal to the transverse sampling distance, i.e., 30 µm in our example. System sensitivity An important issue in OCT imaging is detection sensitivity. Sensitivity can be defined as the smallest signal just discernible from noise. OCT signals are not generated by the tissue, but are caused by backscattered light. The amount of backscattered
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light depends on the incident light intensity and the sample reflectivity Rs. Since we are not interested in the absolute amount of backscattered light but in the tissue parameter Rs, a suitable definition for the sensitivity S in the context of OCT is the ratio between the reflectivity of a perfectly reflecting mirror (R = 1) and the smallest sample reflection coefficient, Rs,min, yielding a signal strength equal to noise: S =
1 Rs,min
.
(4)
The sensitivity is usually specified in dB: r 1 i u. S(dB) = 10logu u u qRs,mint
(5)
As a rule of thumb, a sensitivity of at least 50 dB is necessary to observe a signal from the retina. However, with S = 50 dB, the strongest backscatterer at the retina, the RPE/choriocapillaris complex, is just discernible from noise, and no further details can be observed. For good quality OCT images of the retina, a sensitivity of about 90–100 dB is required. The sensitivity of an OCT system is determined by noise. Several noise components contribute to the overall noise. The main noise components are shot noise (quantum noise), electronic amplifier noise, and excess noise caused by the light source.25,26 If SLDs with low to medium output power are used, as in retinal OCT, and if the electronic components are properly chosen, shot noise is the factor limiting the sensitivity (shot noise limitation is the optimum case; this type of noise is caused by the inherent quantum nature of light, and cannot be eliminated). The main factors determining shot noise are light power and detection bandwidth. In the shot-noise-dominated regime, sensitivity increases linearly in proportion to the power incident on the sample. Therefore, light power should be as high as the safety limit permits. Sensitivity is inversely proportional to detection bandwidth. For high sensitivity, a low bandwidth is preferable, however, this implies a slow scanning speed and long recording times. Therefore, there is a trade-off between recording time and sensitivity. The recording time of 100 A scans within one second, as presently employed in the ZeissHumphrey OCT scanner, seems to be a good compromise. Motion artifacts Ocular motions such as microsaccades or axial eye motions during the recording can cause image artifacts and distortions. Microsaccades are small, rapid rotational eye movements with an amplitude
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of 1-20 minutes of arc. If a large microsaccade occurs during the recording period, the tomographic scan should be repeated. Axial eye motions are more difficult to avoid. Even very small axial movements of about a few tens of microns can severely distort the image. Such movements can easily occur, especially in elderly patients. In order to correct for distortions caused by axial eye motions, digital image processing techniques can be used. Figure 5 presents an example.27 An OCT image was recorded across the fovea of a human retina. Figure 5a shows the raw image. Axial motions are clearly visible: there is an overall backward drift of the eye, overlaid with a small tremor. Figure 5b shows the image after digital post-processing. As can be seen, motion artifacts are considerably reduced by this technique. However, it should be mentioned that the result of the image processing technique depends on the assumption of what the image should look like. One of the common image processing techniques assumes that the RPE should resemble a smooth, straight line. Each deviation is assumed to be caused by axial eye motions. Consequently, the image is corrected to display a flat RPE. While the assumption is meaningful in most cases, there are retinal disorders in which the RPE is not flat. The image processing algorithm would also flatten these images and the disorder might be missed. Therefore, care has to be taken if postprocessed images are used. Another technique that avoids this type of motion artifact problem is the so-called dual-beam OCT technique. This will briefly be discussed in the following section. Alternative and extended OCT techniques Dual-beam OCT The dual-beam OCT technique completely eliminates artifacts caused by axial eye motions. It is based on dual-beam PCI, which was initially introduced to measure axial eye length and other intraocular distances with micrometer precision.8,9,15 A wide range of applications of dual-beam PCI in physiological studies and in therapeutic applications, especially in the context of cataract surgery, have been reported.12,28,29 The main idea of dual-beam PCI and OCT is to use the cornea as a reference surface. Both the reference and the sample beam are directed towards the eye (they form a coaxial ‘dual beam’, see Fig. 6a). Interferometric modulation is recorded at the photodetector, if the path difference cornea–retina is matched by a corresponding path difference on an external Michelson interferometer. Since no absolute positions are measured, but just path differences, any influence of axial eye motions is completely eliminated. In other words: the sample
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Fig. 5. Motion artifacts in OCT imaging. A horizontal cross-section through the fovea of a human eye was recorded in vivo. Data acquisition time: 2.4 seconds. a. Raw image. Artifacts caused by axial eye motions during the recording time are clearly visible. b. Post-processed image. Motion artifacts are considerably reduced. RNFL: retinal nerve fiber layer. (Reproduced from Swanson et al.27 by courtesy of the publisher.)
(retina) moves synchronously with the reference (cornea). Figure 6b is an example of a horizontal section recorded across the fovea of a human eye in vivo.30 Because of the slow scanning speed used with the experimental set up, the transverse resolution is limited by microsaccades to ~100 µm. No digital image post-processing was necessary for producing the tomogram seen in this Figure. The disadvantage of the technique is the increased system complexity and the rather complicated alignment procedure needed to receive the beams reflected at the cornea and retina simultaneously. En-face OCT Conventional OCT images are made up of A scans, i.e., the information on the reflectivity distribution
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within the sample is obtained by subsequent axial scans recorded at different transversal sample positions. In order to obtain information on a whole 3D sample volume, A scans have to be recorded along a 2D grid distributed over the sample. Other scanning patterns can also be used to obtain 3D information. The information can be obtained by transversally ‘slicing’ the sample volume:31 a complete transversal image corresponding to an object depth defined by the reference arm length is recorded by raster scanning the sample beam across the object. Thereby, the distribution of backscattered light within a slice of thickness equal to the coherence length is recorded. After recording one transversal slice, the reference arm length is altered and the next transversal image slice is obtained. This process is repeated until a whole 3D data set has been recorded. While the A-scan tech-
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a.
b. Fig. 6. Dual-beam OCT. a. Schematic. BS: beam splitter; OL: optical length of eye ball. Interferometric signals are recorded, if d h OL. b. Horizontal cross-section through the fovea of a human eye recorded in vivo. Axial motion artifacts are eliminated. A synthesized light source with an effective bandwidth of 50 nm was used, yielding an improved axial resolution of ~6 µm. (Reproduced from Baumgartner et al.30 by courtesy of the publisher.)
nique provides a carrier frequency based on the Doppler shift of the reference light induced by the scanning reference mirror, transversal, or en-face, OCT usually needs additional optical components, such as phase modulators or frequency shifters, for generation of the carrier frequency. Podoleanu et al. have adapted the transversal OCT technique for retinal imaging.32 They used the path length modulation induced by transversally raster scanning the retina with the sample
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beam by a galvo mirror pair for generating the carrier frequency. The disadvantage of this method is that, due to the nonlinear path length dependence on the scanning angle, a frequency spread of the photoelectric signal occurs, leading to a reduced signal-to-noise ratio. The advantage is that no additional phase modulators are needed. The images thus recorded look rather fragmented because, due to the short lc, the coherence condition within one transverse image is only met for small
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Fig. 7. 3D data set recorded by en-face OCT. A 3D data set was recorded from a human optic nerve head in vivo. Sections along the x-z and y-z planes are shown. RPE: retinal pigment epithelium. (Reproduced from Podoleanu et al.34 by courtesy of the publisher.)
areas of the retina. Therefore, several subsequent transverse OCT images have to be recorded and mounted to a 3D data set with the help of a simultaneously recorded SLO image33 (an SLO detection system is integrated with the instrument). From the resulting 3D data set, sections of arbitrary orientation through the retina can be derived by software.34 This can be a considerable advantage because section geometries do not have to be defined in advance; instead, they can be chosen after data acquisition, based on the recorded information. Reported recording times are about 0.5 seconds for one transversal image; a stack of 112 en-face images can be recorded within 56 seconds. Figure 7 shows an image derived from a 3D data set of the optic nerve head of a human eye recorded in vivo.34 Two mutually perpendicular sections across the nerve head, oriented perpendicular to the retinal surface (along the x-z and y-z planes; z: axial direction), are shown. Imaged structures comprise the RPE, retinal nerve fiber layer, and lamina cribrosa.
effect is used in laser Doppler velocimetry to measure flow velocity in a wide range of applications. The short coherence light sources used in OCT enable this effect to be used for depth resolved flow velocity measurements and flow imaging. Different schemes for Doppler OCT have been reported.35-37 The method first reported by Izatt et al.37 is based on the conventional A-scan OCT technique described above under Basics of optical coherence tomography, and is therefore best suited for a brief explanation regarding the basic principles of Doppler OCT in the context of this book. As mentioned above (Partial coherence interferometry), an oscillating coherence signal is generated every time the reference mirror position coincides with the position of a backscatterer in the sample arm. The frequency f0 of this signal equals the Doppler frequency shift of the reference beam caused by the reference mirror moving with speed vr: f0 =
Doppler OCT If light is backscattered by a moving particle, its frequency is shifted by the Doppler frequency. This
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2vr λ0
.
(6)
If the scatterer in the sample arm moves with a speed vs, the frequency of the sample beam is also
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Fig. 8. Color Doppler OCT. A cross-section superior to the optic nerve head was recorded in a human eye in vivo. Structural information is encoded in gray scale; direction and magnitude of blood flow are encoded by color (red or blue) and intensity, respectively. ILM: inner limiting membrane; RPE/CC: retinal pigment epithelium/choriocapillaris complex. (Reproduced from Yazdanfar et al.38 by courtesy of the publisher.)
Doppler shifted, by a frequency: fs =
2vs cosα λ0/n
,
(7)
where α is the angle between the incident sample beam and the direction of the scatterer movement, and n is the refractive index of the medium in which the scatterer moves. The resulting frequency detected by the photodetector is: fDet = f0 ± fs ;
(8)
the sign in equation (8) depends on the movement direction of the scattering particle. To measure and image flow velocity by OCT requires the recording of the full interferometric signal. Just to record the signal envelope is not sufficient. After recording the A scans making up an OCT image, each A scan is analyzed by a shorttime fast Fourier transform, which provides the local distribution of frequencies within the A scan. Deviations from the center frequency f0 are caused by moving scatterers. These deviations are extracted and converted into color values, indicating the local particle speed. The data sets are then mounted in the usual way in order to provide a map of velocity distribution within the sample. The signal intensities contained in the A scans can be separately displayed as conventional OCT images of the sample structure. Figure 8 shows a Doppler OCT image recorded in this way.38 A cross-section recorded in vivo in a human retina, at a location superior to the optic nerve head, is shown. A structural OCT image
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displaying reflected intensities (gray scale image) is superimposed with flow information reconstructed from Doppler shifts. The sample contains two vessels, branches of the central retinal artery and vein superior, whose cross-sections can be seen in the tomogram. Blood cells of opposite flow directions cause positive and negative frequency shifts in the light backscattered from within the two vessels, indicated in red and blue. The saturation of the color is a measure of the blood cell speed. It should be mentioned that there is a trade-off between spatial resolution and velocity resolution. A good velocity resolution requires a larger Fourier window, i.e., a larger segment of an A scan is needed to obtain the necessary amount of data. However, a longer A-scan segment implies reduced spatial resolution in an axial direction. More sophisticated Doppler OCT schemes that overcome this problem have recently been reported.39,40 Conclusions and Outlook Several improvements and extensions of OCT are presently under development. The aim of these developments is essentially improved resolution, speed, and image contrast. Improved axial resolution requires light sources of a higher bandwidth. Recently, the first retinal OCT images recorded with a state-of-the-art Ti: Al2O3 femtosecond pulse laser were presented.13 Corresponding images can be seen in Figures 4 and 5 of the chapter by J.S. Schuman et al. (see p. 147). Using a bandwidth of 155 nm, a depth re-solution of ~3 µm was obtained. Comparison with a standard resolution image clearly demonstrates the dramatic
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Principles of optical coherence tomography improvement in resolution. Implementation of this technique into clinically-useful OCT systems depends on the availability of a cheap and robust light source. With presently available standard OCT systems, one OCT image consisting of 100 A scans is recorded in one second. Acquiring 3D information is presently rather time consuming. Different approaches for solving this problem are under investigation. Rapid scanning optical delay lines have been reported that achieve more than 1000 A scans per second.41,42 Parallel OCT schemes recording up to 58 × 58 A scans simultaneously with speciallydesigned detector arrays have been suggested43. Finally, a variant of en-face OCT that can record a 3D data set consisting of 32 transverse OCT images within one second has been announced.44 All these techniques depend on light sources providing sufficient output power to achieve high sensitivity. Image contrast in conventional OCT is based on the intensity of backscattered light. Several structures yield poor or no contrast on an intensity basis. Other properties of light, as wavelength and polarization, can be exploited to improve contrast or to generate other types of contrast. OCT at wavelengths with different absorption coefficients in oxygenated and de-oxygenated blood has been suggested for measuring and imaging oxygen saturation in vessels.45 Polarization-sensitive OCT has been reported as a means of measuring the thickness of the birefringent nerve fiber layer for glaucoma diagnostics.46 These developments seem promising technologies for future, enhanced OCT applications. However, much additional research and developmental work has to be done before they can be converted into robust, clinically-applicable tools. Acknowledgments The author acknowledges the permission of several authors and editors to reproduce figures from their works. Furthermore, the author wishes to thank Professor A.F. Fercher, head of the Institute of Medical Physics, University of Vienna, and several other co-workers, for fruitful discussions. Part of the work reported in this chapter was financed by the Austrian Science Foundation.
References 1. Huang D, Swanson EA, Lin CP, Schuman JS, Stinson WG, Chang W, Hee MR, Flotte T, Gregory K, Puliafito CA, Fujimoto JG: Optical coherence tomography. Science 254:1178-1181, 1991 2. Fercher AF, Hitzenberger CK: Optical coherence tomography in medicine. In: Asakura T (ed) International Trends in Optics and Photonics, ICO IV, pp 359-389. Berlin: Springer 1999 3. Fercher AF, Hitzenberger CK, Drexler W, Kamp G, Sattmann H: In vivo optical coherence tomography. Am J Ophthalmol 116:113-114, 1993
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71 4. Hee MR, Izatt JA, Swanson EA, Huang D, Schuman JS, Lin CP, Puliafito CA, Fujimoto JG: Optical coherence tomography of the human retina. Arch Ophthalmol 113:325-332, 1995 5. Puliafito CA, Hee MR, Lin CP, Reichel E, Schuman JS, Duker JS, Izatt JA, Swanson EA, Fujimoto JG: Imaging of macular diseases with optical coherence tomography. Ophthalmology 102:217-229, 1995 6. Schuman JS, Hee MR, Puliafito CA, Wong C, PedutKloizman T, Lin CP, Hertzmark E, Izatt JA, Swanson EA, Fujimoto JG: Quantification of nerve fiber layer thickness in normal and glaucomatous eyes using optical coherence tomography. Arch Ophthalmol 113:586-596, 1995 7. Radhakrishnan S, Rollins AM, Roth JE, Yazdanfar S, Westphal V, Bardenstein DS, Izatt JA: Real-time optical coherence tomography of the anterior segment at 1310 nm. Arch Ophthalmol 119:1179-1185, 2001 8. Fercher AF, Mengedoht K, Werner W: Eye-length measurement by interferometry with partially coherent light. Opt Lett 13:186-188, 1988 9. Hitzenberger CK: Optical measurement of the axial eye length by laser Doppler interferometry. Invest Ophthalmol Vis Sci 32:616-624, 1991 10. Huang D, Wang J, Lin CP, Puliafito CA, Fujimoto JG: Micron-resolution ranging of cornea anterior chamber by optical reflectometry. Lasers Surg Med 11:419-425, 1991 11. Hitzenberger CK: Measurement of corneal thickness by low coherence interferometry. Appl Opt 31:6637-6642, 1992 12. Hitzenberger CK, Drexler W, Dolezal C, Skorpik F, Juchem M, Fercher AF, Gnad HD: Measurement of the axial length of cataract eyes by laser Doppler interferometry. Invest Ophthalmol Vis Sci 34:1886-1893, 1993 13. Drexler W, Morgner U, Ghanta RK, Schuman JS, Kärtner FX, Hee MR, Ippen EP, Fujimoto JG: New technology for ultrahigh resolution optical coherence tomography of the retina. In: Lemij HG, Schuman JS (eds) The Shape of Glaucoma: Quantitative Neural Imaging Techniques, pp 75-104. The Hague: Kugler Publications 2000 14. Born M, Wolf E: Principles of Optics. Oxford: Pergamon 1987 15. Drexler W, Baumgartner A, Findl O, Hitzenberger CK, Sattmann H, Fercher AF: Submicrometer precision biometry of the anterior segment of the human eye. Invest Ophthalmol Vis Sci 38:1304-1313, 1997 16. Schmitt JM, Kumar G: Optical scattering properties of soft tissue: a discrete particle model. Appl Opt 37:27882797, 1998 17. Bouma B, Tearney GJ, Boppart SA, Hee MR, Brezinski ME, Fujimoto JG: High-resolution optical coherence tomographic imaging using a mode-locked Ti:Al2O3 laser source. Opt Lett 20:1486-1488, 1995 18. American National Standards Institute: Safe use of lasers, ANSI Z 136.1. New York, NY: American National Standards Institute 1993 19. Swanson EA, Huang D, Hee MR, Fujimoto JG, Lin CP, Puliafito CA: High-speed optical coherence domain reflectometry. Opt Lett 17:151-153, 1992 20. Drexler W, Hitzenberger CK, Baumgartner A, Findl O, Sattmann H, Fercher AF: Investigation of dispersion effects in ocular media by multiple wavelength partial coherence interferometry. Exp Eye Res 66:25-33, 1998 21. Hitzenberger CK, Baumgartner A, Drexler W, Fercher AF: Dispersion effects in partial coherence interferometry: implications for intraocular ranging. J Biomed Opt 4:144151, 1999 22. Wilson T: Confocal microscopy. In: Wilson T (ed) Confocal Microscopy, pp 1-64. London: Academic Press 1990
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23. Gerrard A, Burch JM: Introduction to Matrix Methods in Optics. London: John Wiley & Sons 1975 24. Lexer F, Hitzenberger CK, Drexler W, Molebny S, Sattmann H, Sticker M, Fercher AF: Dynamic coherent focus OCT with depth-independent transversal resolution. J Mod Optics 46:541-553, 1999 25. Podoleanu AG, Jackson DA: Noise analysis of a combined optical coherence tomograph and a confocal scanning ophthalmoscope. Appl Opt 38:2116-2127, 1999 26. Rollins AM, Izatt JA: Optimal interferometer designs for optical coherence tomography. Opt Lett 24:1484-1486, 1999 27. Swanson EA, Izatt JA, Hee MR, Huang D, Lin CP, Schuman JS, Puliafito CA, Fujimoto JG: In vivo retinal imaging by optical coherence tomography. Opt Lett 18:1864-1866, 1993 28. Drexler W, Findl O, Menapace R, Rainer G, Vass C, Hitzenberger CK, Fercher AF: Partial coherence interferometry: a novel approach to biometry in cataract surgery. Am J Ophthalmol 126:524-534, 1998 29. Findl O, Drexler W, Menapace R, Heinzl H, Hitzenberger CK, Fercher AF: Improved prediction of intraocular lens power using partial coherence interferometry. J Cat Ref Surg 27:861-867, 2001 30. Baumgartner A, Hitzenberger CK, Ergun E, Stur M, Sattmann H, Drexler W, Fercher AF: Resolution improved dual beam- and standard optical coherence tomography: a comparison. Graefe’s Arch Clin Exp Ophthalmol 238:385-392, 2000 31. Izatt JA, Hee MR, Owen GM, Swanson EA, Fujimoto JG: Optical coherence microscopy in scattering media. Opt Lett 19:590-592, 1994 32. Podoleanu AG, Dobre GM, Seeger M, Webb DJ, Jackson DA, Fitzke FW, Halfyard AS: Low coherence interferometry for en-face imaging of the retina. Lasers Light Ophthalmol 8:187-192, 1998 33. Podoleanu AG, Jackson DA: Combined optical coherence tomograph and scanning laser ophthalmoscope. Electron Lett 34:1088-1090, 1998 34. Podoleanu AG, Rogers JA, Jackson DA, Dunne S: Three dimensional OCT images from retina and skin. Opt Express 7:292-298, 2000 35. Wang XJ, Milner TE, Nelson JS: Characterization of fluid
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flow velocity by optical Doppler tomography. Opt Lett 20:1337-1339, 1995 Chen Z, Milner TE, Srinivas S, Wang X, Malekafzali A, Van Gemert MJC, Nelson JS: Noninvasive imaging of in vivo blood flow velocity using optical Doppler tomography. Opt Lett 22:1119-1121, 1997 Izatt JA, Kulkarni MD, Yazdanfar S, Barton JK, Welch AJ: In vivo bidirectional color Doppler flow imaging of picoliter blood volumes using optical coherence tomography. Opt Lett 22:1439-1441, 1997 Yazdanfar S, Rollins AM, Izatt JA: Imaging and velocimetry of the human retinal circulation with color Doppler optical coherence tomography. Opt Lett 25:1448-1450, 2000 Zhao Y, Chen Z, Saxer C, Xiang S, De Boer JF, Nelson JS: Phase-resolved optical coherence tomography and optical Doppler tomography for imaging blood flow in human skin with fast scanning speed and high velocity sensitivity. Opt Lett 25:114-116, 2000 Van Leeuwen TG, Kulkarni MD, Yazdanfar S, Rollins AM, Izatt JA: High-flow-velocity and shear-rate imaging by use of color Doppler optical coherence tomography. Opt Lett 24:1584-1586, 1999 Tearney GJ, Bouma BE, Fujimoto JG: High speed phaseand group- delay scanning with a grating based phase control delay line. Opt Lett 22:1811-1813, 1997 Rollins AM, Kulkarni MD, Yazdanfar S, Ung-arunyawee R, Izatt JA: In vivo video rate optical coherence tomography. Opt Express 3:219-229, 1998 Bourquin S, Seitz P, Salathe RP: Optical coherence topography based on a two-dimensional smart detector array. Opt Lett 26:512-514, 2001 Hitzenberger CK, Zhou Q, Trost P, Schmode S, Baumgartner A: High-speed three-dimensional optical coherence tomography for retinal imaging. ARVO Abstract no 2856, 2002 Faber DJ, Mik EG, Van der Meer FJ, Aalders MCG, Van Leeuwen TG: Oxygenation measurements in blood with optical coherence tomography. Proc SPIE 4619 (in press) Ducros MG, Marsack JD, Rylander HG III, Thomsen SL, Milner TE: Primate retina imaging with polarization-sensitive optical coherence tomography. J Opt Soc Am A 18:2945-2956, 2001
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From physical energy to biological effect
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From physical energy to biological effect How retinal laser treatment affects diabetic retinopathy
Einar Stefánsson Department of Ophthalmology, Landspítali, University of Iceland, Reykjavík, Iceland
Keywords: diabetes, diabetic retinopathy, photocoagulation, theory of photocoagulation, retinal oxygenation
•
Abstract Retinal photocoagulation has been used for diabetic and other ischemic retinopathies for decades. While most ophthalmologists are convinced of its clinical effect, many do not understand how if works, i.e., the mechanism of effect is unclear. The purpose of this chapter is to present a theory on how the physical energy of visible laser light, such as that of the argon laser, is transferred into a biological effect and a therapeutic effect on retinopathy.
Introduction The physiological mechanism of photocoagulation can be seen in following steps: • The physical light energy is absorbed in the melanin of the retinal pigment epithelium. The adjacent photoreceptors are destroyed and are replaced by a glial scar, and the oxygen consumption of the outer retina is reduced. • Oxygen that normally diffuses from the choriocapillaris into the retina can now diffuse through the laser scars in the photoreceptor layer without being consumed in the mitochondria of the photoreceptors. • This oxygen flux reaches the inner retina in order to relieve inner retinal hypoxia and raise oxygen tension. As a result, the retinal arterioles constrict and the blood flow decreases. • Vasoconstriction increases arteriolar resistance, decreases hydrostatic pressure in the capillaries and venules, and reduces edema formation according to Starling’s law. • Hypoxia relief reduces the production of growth factors such as vascular endothelial growth factor (VEGF), and neovascularization is reduced or halted.
Retinal laser photocoagulation improves inner retinal oxygenation, which affects retinopathy through the relief of hypoxia and consequent changes in growth factor production and hemodynamics.
Light coagulation and laser treatment of the retina were introduced to ophthalmology around the middle of the last century. They are widely used for the treatment of diabetic retinopathy and other ischemic retinopathies. In this chapter we will focus on argon laser treatment and diabetic retinopathy, even though the theory applies to all visible light photocoagulation for ischemic retinopathies. The usefulness of retinal photocoagulation was established in clinical trials, such as the Diabetic Retinopathy Study1,2 and the Early Treatment Diabetic Retinopathy Study.3-8 The mechanism of treatment was not investigated in these trials and eludes many ophthalmologists to this day. Ophthalmic physiologists have strived to understand the physiologic mechanism using histological, physiological, and clinical research techniques. This paper proposes a general theory on the mechanism of action of retinal laser treatment, based on these studies.
Histology Visible light that is sent into the eye is predominantly absorbed in the melanin of the retinal pigment epithelium, and the heat results in thermal destruction of retinal pigment epithelium cells, the adjacent photoreceptors and, to some degree, the choriocapillaris.9-14 Figure 1 shows a pigmented rabbit retina with a two-week-old argon laser burn. The outer retina and retinal pigment epithelium are photoco-
Address for correspondence: Einar Stefánsson, MD, PhD, Department of Ophthalmology, Landspítali, University of Iceland, 101 Reykjavík, Iceland. e-mail:
[email protected] Lasers in Ophthalmology – Basic, Diagnostic and Surgical Aspects, pp. 73–78 edited by F. Fankhauser and S. Kwasniewska © 2003 Kugler Publications, The Hague, The Netherlands
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Fig. 1. Light micrograph of a two-week-old argon laser burn in a pigmented rabbit retina. The outer retina and retinal pigment epithelium are photocoagulated and the photoreceptors are absent, whereas the inner retina is relatively intact. RPE: pigmented retinal pigment epithelium; CC: choriocapillaris; CV: large choroidal vessels. (Reproduced from Novak et al.21 by courtesy of the publisher.)
are replaced by glia, which have few mitochondria and a low oxygen consumption. This means that the laser scars function as windows where oxygen consumption is low and it can diffuse from the choroid through the photoreceptor layer into the inner retina (Fig. 2). This should lead to increased oxygen tension in the inner retina. Oxygen physiology
Fig. 2. Schematic drawing indicating an oxygen flux coming from the choroid and passing through a laser scar into the ischemic inner retina. In the laser scar, photoreceptors are replaced by glia, and the oxygen consumption is decreased. The oxygen flux from the choroid would normally be consumed by the photoreceptors, but since the oxygen consumption of the glia is less, the oxygen flux reaches the inner retina, where ischemic hypoxia may be relieved. (Redrawn with permission from Stefánsson E, Graefe’s Arch Clin Exp Ophthalmol 228:120-123, 1990.)
agulated and the photoreceptors are absent, whereas the inner retina is relatively intact. The histology predicts the physiological effect. In the normal situation, oxygen and nutrients diffuse from the choriocapillaris into the retina and these are consumed by the photoreceptors, which have a very high density of mitochondria and a high oxygen consumption. In laser scars, the photoreceptors
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Oxygen tension in the inner retina following laser treatment has been studied in a number of experimental animals and diabetic patients. The improved oxygenation following retinal laser treatment was first shown in rhesus monkeys, when Stefánsson et al.15 showed that retinal oxygen tension was much higher in laser-treated areas than in untreated areas of the same retina. These findings have been confirmed by a number of researchers,16-18 including Pournaras et al., who showed that laser treatment reverses retinal hypoxia induced by branch retinal vein occlusion.19 Diddie and Ernest found an initial increase in retinal oxygen tension following retinal laser treatment, but this effect was short lived.20 Figure 3 shows data from rabbits, part of whose retina was photocoagulated and the other part not. The oxygen tension of the retina was much higher over the photocoagulated area compared to the untreated area. This effect was seen even as early as one day following treatment.21 Oxygen tension measurements from the preretinal vitreous of human diabetics undergoing vitrectomy have shown the same result, where the oxygen tension over lasertreated areas of the retina was significantly higher than the oxygen tension over untreated areas of the same retina.22
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From physical energy to biological effect
Fig. 3. Graph showing preretinal oxygen tension (torr = mmHg) in rabbits where part of the rabbit retina was photocoagulated and another part was not. The oxygen tension in the retina was significantly higher in the photocoagulated area compared to the untreated area. This effect was seen even as early as one day following treatment. (Reproduced from Novak et al.21 by courtesy of the publisher.)
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Fig. 4. Graph showing the correlation between decreasing arteriolar diameter following panretinal photocoagulation in the Diabetic Retinopathy Study and the regression of disc neovascularization. There was a statistically significant correlation between the vasoconstriction and the disappearance of disc neovascularization, according to the DRS graders. (Reproduced with permission from Wilson et al., Am J Ophthalmol 106:657-664, 1988.)
Vascular physiology The retinal circulation autoregulates its own diameter and blood flow. It dilates in hypoxia and constricts if the oxygenation is increased. The improved oxygenation of the inner retina should lead to constriction of the retinal arterioles and venules, and to decreased blood flow following laser treatment. This was originally proposed by Wolbarsht et al.,23,24 and has been tested and demonstrated in numerous other studies.25-32 Hessemer and Schmidt suggested that total ocular blood flow is also decreased after scatter laser treatment of the retina.33 Wilson et al. used photographs and data from the Diabetic Retinopathy Study (DRS) 2 to evaluate the diameters of retinal vessels in patients treated with panretinal photocoagulation for proliferative diabetic retinopathy.34 On DRS photographs, arteriolar and venular constriction was seen following both laser and xenon arc photocoagulation. Furthermore, there was a statistically significant correlation between the vasoconstriction and the disappearance of disc neovascularization, according to the DRS graders (Fig. 4). Remky et al. also observed constriction of the retinal venules following laser treatment.35 Retinal vasoconstriction is also seen following macular grid photocoagulation.36 The temporal arterioles and venules and their macular branches all constrict significantly following photocoagulation for diabetic macular edema, and the same is also true of laser treatment in branch retinal vein occlusion.37 In diabetic macular edema, Kristinsson et al. documented progressive vasodilatation prior to the onset of edema, and constriction following laser treatment as the edema disappears (Fig. 5).38
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Fig. 5. Graph showing the diameter of a temporal retinal arteriole in a diabetic patient. In 1989, the female patient was diagnosed with diabetic macular oedema (DMO) and received macular grid laser treatment. The graph demonstrates progressive vasodilatation prior to the onset of edema and constriction following laser treatment as the edema disappears. (Reproduced from Kristinsson et al.36,38 by courtesy of the publisher.)
Proliferative retinopathy Improved oxygenation of the inner retina and the relief of hypoxia lower the production of VEGF in the retina, and hinder neovascularization. Aiello et al.39 and Augustin et al.40 showed that VEGF levels in the vitreous of diabetic patients are elevated in patients with proliferative retinopathy, and are significantly lower following retinal photocoagulation. Lip et al.41 showed that VEGF levels were elevated in the blood of diabetics with proliferative
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Fig. 6. Schematic flow diagram explaining the mechanism of effect for retinal photocoagulation on retinal neovascularization and macular edema in diabetic retinopathy and related retinopathies. See text for detailed explanation.
retinopathy, and fell following panretinal photocoagulation. Previously, Boulton et al.42 and Smith et al.43 had shown that VEGF and VEGF receptors in the diabetic retina are present in proliferative retinopathy, and are reduced in eyes that have had laser treatment. Pournaras et al.44 demonstrated that hyperoxia also reduces VEGF production in ischemic retina, suggesting that laser treatment and hyperoxia have the same effect on VEGF production. Miller et al.45 showed, in an experimental animal, that VEGF levels are elevated when retinal ischemia is induced by means of a branch vein occlusion, demonstrating the correlation between ischemia (hypoxia) and VEGF levels. In addition to VEGF reduction following retinal photocoagulation, reduced vasodilatation and endothelial stretching may also reduce the effect of growth factors on vascular endothelium. This was first suggested in 1983,46 based on clinical observations, and recent laboratory data have pointed to the role of vascular tissue stretching in vasoproliferation.47-50 Suzuma et al. found that capillary stretching increases thymidine uptake and VEGF production in capillaries.51 Retinal hypoxia stimulates neovascularization in two ways, firstly by hypoxia-induced VEGF production, and secondly through the autoregulatory
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dilatation of capillaries, which stimulates growth directly. Reduction in retinal hypoxia through laser treatment will reduce VEGF formation and capillary stretching, and reduce neovascularization through both mechanisms (Fig. 6). Macular edema By definition, edema is an abnormal accumulation of water in a tissue, and the development and regression of edema is based on the movement of water between vascular and tissue compartments. Starling’s law describes the steady-state water exchange between the vascular compartment and the extracellular tissue compartment. Hydrostatic pressure in the vessel drives water into the tissue, and this is opposed by oncotic (osmotic) pressure differences between blood and tissue. In the normal state, these forces are in balance and there is no net movement of water between tissue and vascular compartments. However, if the hydrostatic pressure in the capillaries and venules is increased, this drives water into the tissue and creates edema, whereas decreased hydrostatic blood pressure would decrease edema, assuming that the oncotic pressures are constant.
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From physical energy to biological effect The retinal arterioles serve as resistance vessels and control the hydrostatic pressure downstream. Dilated arterioles have less resistance, and consequently the blood flow and hydrostatic pressure are increased downstream in the capillaries and venules, where high hydrostatic pressure dilates these thinwalled vessels, according to LaPlace’s law. The diameter of the arterioles and venules in the retina is an indicator of the hydrostatic blood pressure in the retinal microcirculation. We have observed that, prior to the development of diabetic macular edema, the retinal arterioles and venules gradually dilate.38 Bek also observed the role of vasodilatation in diabetic macular edema.52 Following macular laser treatment, the arterioles and venules constrict as the retinal edema regresses (Fig. 5).36 The same pattern is seen in branch retinal vein occlusion.37 The regression of retinal edema following macular laser treatment is easily understood in light of Starling’s law. The laser treatment reduces oxygen consumption in the outer retina, and extends the oxygen flux from the choroid into the inner retina. This causes the arterioles to constrict (autoregulation), their resistance is increased to the fourth power of the radius, and the hydrostatic pressure downstream is decreased. Starling’s law predicts that this will reduce the water flux from the vessel into the tissue, and the oncotic pressure will now manage to drive the water back into the vessels and reduce the edema. The decreased hydrostatic pressure will also reduce the dilatation of the venules, as has been observed (Fig. 6). Since vascular hydrostatic pressure plays a central role in Starling’s law, this explains why arterial hypertension is so important in diabetic edema, and why reducing the blood pressure may be helpful against diabetic macular edema.53 Conclusions This chapter describes a general theory on the mechanism of action of retinal photocoagulation in diabetic retinopathy and related retinopathies (Fig. 6). The theory focuses on the effect of the treatment on retinal oxygenation, and the consequence of the oxygenation effect. The same mechanism can be used to understand the effect of laser treatment on different aspects of diabetic retinopathy, as well as the effect of vitrectomy on diabetic retinopathy.54 References 1. Diabetic Retinopathy Study Research Group: Photocoagulation treatment of proliferative diabetic retinopathy. Trans Am Acad Ophthalmol Otol 85:82-106, 1978 2. Diabetic Retinopathy Study Research Group: Photocoagulation treatment of proliferative diabetic retinopathy: clinical application of Diabetic Retinopathy Study (DRS) findings. DRS Report No. 8. Ophthalmology 88:583-600, 1981
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77 3. Early Treatment Diabetic Retinopathy Study Research Group: Photocoagulation for diabetic macular edema: early treatment diabetic retinopathy study report No. 1. Arch Ophthalmol 103:1796-1806, 1985 4. Early Treatment Diabetic Retinopathy Study Research Group: The Early Treatment Diabetic Retinopathy Study Report No. 7. Ophthalmology 98:741-756, 1987 5. Vatne HO, Syrdalen P: Long-term visual results after laser treatment of proliferative diabetes retinopathy in childhood onset diabetes. Acta Ophthalmol (Scand) 66(2):161-164, 1988 6. Stefánsson E, Bek T, Porta M, Larsen N, Kristinsson JK, Agardh E: Screening and prevention of diabetic blindness. Acta Ophthalmol (Scand) 78(4):374-385, 2000 7. Kristinsson JK: Diabetic retinopathy: screening and prevention of blindness. Acta Ophthalmol (Scand) 75(Suppl 223):176, 1997 8. Stenkula S: Photocoagulation in diabetic retinopathy: a multicentre study in Sweden. Acta Ophthalmol Suppl 162:1100, 1984 9. Campbell CJ, Rittler MC, Swope CH et al: The ocular effects produced by experimental lasers: IV. The argon laser. Am J Ophthalmol 67:671-681, 1969 10. Powell JO, Bresnick GH, Yanoff M et al: Ocular effects of argon laser radiation: II. Histopathology of chorioretinal lesions. Am J Ophthalmol 71:1267-1276, 1971 11. Apple DJ, Goldberg MF, Wyhinny G: Histopathology and ultrastructure of the argon laser lesion in human retinal and choroidal vasculature. Am J Ophthalmol 75:595-609, 1973 12. Tso MOM, Wallow JHL, Elgin S: Experimental photocoagulation of the human retina: I. Correlation of physical, clinical and pathological data. Arch Ophthalmol 95:10351040, 1977 13. Wallow JHL, Tso MOM, Elgin S: Experimental photocoagulation of the human retina: II. Electron microscopic study. Arch Ophthalmol 95:1041-1050, 1977 14. Suomalainen VP: Comparison of retinal lesions produced by transscleral krypton laser photocoagulation, transpupillar krypton laser photocoagulation and cryocoagulation. Acta Ophthalmol (Scand) 71(2):224-229, 1993 15. Stefánsson E, Landers MB III, Wolbarsht ML: Increased retinal oxygen supply following panretinal photocoagulation and vitrectomy and lensectomy. Trans Am Ophthalmol Soc 79:307-334, 1981 16. Molnar I, Poitry S, Tsacopoulos M, Gilodi N, Leuenberger PM: Effect of laser photocoagulation on oxygenation of the retina in miniature pigs. Invest Ophthalmol Vis Sci 26:14101414, 1985 17. Funatsu H, Hori S, Yamashita H, Kitano S: Effective mechanisms of laser photocoagulation for neovascularization in diabetic retinopathy. Nippon Ganka Gakkai Zasshi 100:339349, 1996 18. Funatsu H, Wilson CA, Berkowitz BA, Sonkin PL: A comparative study of the effects of argon and diode laser photocoagulation on retinal oxygenation. Graefe’s Arch Clin Exp Ophthalmol 235:168-175, 1997 19. Pournaras CJ, Tsacopoulos M, Strommer K, Gilodi N, Leuenberger PM: Scatter photocoagulation restores tissue hypoxia in experimental vasoproliferative microangiopathy in miniature pigs. Ophthalmology 97:1329-1333, 1990 20. Diddie KR, Ernest JT: The effect of photocoagulation on the choroidal vasculature and retinal oxygen tension. Am J Ophthalmol 84:62-66, 1977 21. Novak RL, Stefánsson E, Hatchell DL: The effect of photocoagulation on the oxygenation and ultrastructure of avascular retina. Exp Eye Res 50:289-296, 1990 22. Stefánsson E, Machemer R, McCuen BW, deJuan E, Peterson J: Retinal oxygenation and laser treatment in patients with
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diabetic retinopathy. Am J Ophthalmol 113:36-38, 1992 23. Wolbarsht ML, Landers MB: The rationale of photocoagulation therapy for proliferative diabetic retinopathy: a review and a model. Ophthalmic Surg 11:235-245, 1980 24. Wolbarsht ML, Landers MB III, Stefánsson E: Vasodilation and the etiology of diabetic retinopathy: a new model. Ophthalmic Surg 12(2):2104-2107, 1981 25. Feke GT, Green GJ, Goge DG, McMeel JW: Laser Doppler measurements of the effect of panretinal photocoagulation on retinal blood flow. Ophthalmology 89:757-762, 1982 26. Richard G, Kreissig I: Effect of laser therapy in diabetic retinopathy on the hemodynamics of retinal vessels. Klin Mbl Augenheilk 186(2):107-109, 1985 27. Grunwald JE, Riva CE, Brucker AJ, Sinclair SH, Petrig BL: Effect of panretinal photocoagulation on retinal blood flow in proliferative diabetic retinopathy. Ophthalmology 93:590595, 1986 28. Grunwald JE, Brucker AJ, Petrig BL, Riva CE: Retinal blood flow regulation and the clinical response to panretinal photocoagulation in proliferative diabetic retinopathy. Ophthalmology 96:1518-1522, 1989 29. Fujio N, Feke GT, Goger DG, McMeel JW: Regional retinal blood flow reduction following half fundus photocoagulation treatment. Br J Ophthalmol 78:335-338, 1994 30. Mendivil A, Cuartero V: Ocular blood flow velocities in patients with proliferative diabetic retinopathy after scatter photocoagulation: two years of follow-up. Retina 16(3): 222-227, 1996 31. Mendivil A: Ocular blood flow velocities in patients with proliferative diabetic retinopathy after panretinal photocoagulation. Surv Ophthalmol Suppl 41:S89-95, 1997 32. Hiroshiba N, Ogura Y, Nishiwaki H, Miyamoto K, Honda Y: Alterations of retinal microcirculation in response to scatter photocoagulation. Invest Ophthalmol Vis Sci 39(5): 769-776, 1998 33. Hessemer V, Schmidt KG: Influence of panretinal photocoagulation on the ocular pulse curve. Am J Ophthalmol 123(6):748-752, 1997 34. Wilson CA, Stefánsson E, Klombers L, Hubbard LD, Kaufman S, Ferris FL: Optic disc neovascularization and retinal vessel diameter in diabetic retinopathy. Am J Ophthalmol 106:131-134, 1988 35. Remky A, Arend O, Beausencourt E, Elsner AE, Bertram B: Retinal vessels before and after photocoagulation in diabetic retinopathy. Determining the diameter using digitized color fundus slides. Klin Mbl Augenheilk 209(2/3):7983, 1996 36. Gottfredsdóttir MS, Stefánsson E, Jónasson F, Gíslason I: Retinal vasoconstriction after laser treatment for diabetic macular edema. Am J Ophthalmol 115(1):64-67, 1993 37. Arnarsson A, Stefánsson E: Laser treatment and the mechanism of edema reduction in branch retinal vein occlusion. Invest Ophthalmol Vis Sci 41(3):877-879, 2000 38. Kristinsson JK, Gottfredsdóttir MS, Stefánsson E: Retinal vessel dilatation and elongation precedes diabetic macular oedema. Br J Ophthalmol 81(4):274-278, 1997 39. Aiello LP, Avery RL, Arrigg PG, Keyt BA, Jampel HD, Shah ST, Pasquae LR, Thieme H, Iwamoto MA, Park JE: Vascular endothelial growth factor in ocular fluid of patients with diabetic retinopathy and other retinal disorders. N Engl J Med 331(22):1480-1487, 1994 40. Augustin AJ, Keller A, Koch F, Jurklies B, Dick B: Effect
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of retinal coagulation status on oxidative metabolite and VEGF in 208 patients with proliferative diabetic retinopathy. Klin Mbl Augenheilk 218:89-94, 2001 Lip PL, Belgore F, Blann AD, Hope-Ross MW, Gibson JM, Lip GY: Plasma VEGF and soluble VEGF receptor FLT1 in proliferative retinopathy: relationship to endothelial dysfunction and laser treatment. Invest Ophthalmol Vis Sci 41(8):2115-2119, 2000 Boulton M, Foreman D, Williams G, McLeod D: VEGF localisation in diabetic retinopathy. Br J Ophthalmol 82(5): 561-568, 1998 Smith G, McLeod D, Foreman D, Boulton, M: Immunolocalisation of the VEGF receptors FLT-1, KDR, and FLT4 in diabetic retinopathy. Br J Ophthalmol 83(4):486-494, 1999 Pournaras CJ, Miller JW, Gragoudas ES, Husain D, Munoz JL, Tolentino MJ, Kuroki M, Adamis AP: Systemic hyperoxia decreases vascular endothelial growth factor gene expression in ischemic primate retina. Arch Ophthalmol 115:1553-1558, 1997 Miller JW, Adamis AP, Shima DT, D’Amore PA, Moulton RS, O’Reilly MS, Folkman J, Dvorak HF, Brown LF, Berse B: Vascular endothelial growth factor/vascular permeability factor is temporally and spatially correlated with ocular angiogenesis in a primate model. Am J Pathol 145(3): 574-584, 1994 Stefánsson E, Landers MB III, Wolbarsht ML: Oxygenation and vasodilatation in relation to diabetic and other proliferative retinopathies. Ophthalmic Surg 14:209-226, 1983 Seko Y, Fujikura H, Pang J, Tokoro T, Shimokawa H: Induction of vascular endothelial growth factor after application of mechanical stress to retinal pigment epithelium of the rat in vitro. Invest Ophthalmol Vis Sci 40(13):32873291, 1999 Li Q, Muragaki Y, Ueno H, Ooshima A: Stretch-induced proliferation of cultured vascular smooth muscle cells and a possible involvement of local renin-angiotensin system and platelet-derived growth factor (PDGF). Hypertens Res 20(3):217-223, 1997 Zeidan A, Nordstrom I, Dreja K, Malmqvist U, Hellstrand P: Stretch-dependent modulation of contractility and growth in smooth muscle of rat portal vein. Circ Res 87(3):228234, 2000 Hudlicka O: Is physiological angiogenesis in skeletal muscle regulated by changes in microcirculation? Microcirculation 5(1):5-23, 1998 Suzuma I, Hata Y, Clermont A, Pokras F, Rook SL, Suzuma K, Feener EP, Aiello LP: Cyclic stretch and hypertension induce retinal expression of vascular endothelial growth factor and vascular endothelial growth factor receptor-2: potential mechanisms for exacerbation of diabetic retinopathy by hypertension. Diabetes 50(2):444-454, 2001 Bek T: Diabetic maculopathy caused by disturbances in retinal vasomotion: a new hypothesis. Acta Ophthalmol (Scand) 77(4):376-380, 1999 United Kingdom Prospective Diabetes Study Group: Intensive blood-glucose control with sulphonylureas or insulin compared with conventional treatment and the risk of complications in patients with type 2 diabetes (UKPDS 33). Lancet 352:837-853, 1998 Stefánsson E: The therapeutic effects of retinal laser treatment and vitrectomy: a theory based on oxygen and vascular physiology. Acta Ophthalmol (Scand) 79:435-440, 2001
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High-resolution multiphoton imaging and nanosurgery of the cornea
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High-resolution multiphoton imaging and nanosurgery of the cornea using femtosecond laser pulses Karsten König Center for Lasermicroscopy, Faculty of Medicine, Friedrich Schiller University Jena and JenLab GmbH, Jena, Germany
Keywords: laser surgery, femtosecond laser, optical tomography, cornea
Introduction Conventional laser techniques for corneal surgery are based on the application of high energy ultraviolet (UV) nanosecond (ns) laser pulses for the precise photoablation of stromal tissue.1 In general, an argon fluorite (ArF) excimer laser at 193 nm, at a low pulse repetition rate of some hundred Hz, is used as laser source. The energy of a single photon of 6.4 eV is sufficient to break molecular bonds (photodecomposition). For example, a dissociation energy of 3.6 eV is required for the C-C bond as well as the C-O bond, and 4.8 eV for the O-H bond. An energy of 6.4 eV is necessary to break C=C bonds. As a result of 193-nm laser exposure, fragments of low molecular weight are ejected. This non-thermal process is called ablative photodecomposition. Due to the high absorption coefficients, the light penetration depth for 193-nm radiation is limited to < 5 µm.2 The radiation is mainly absorbed by cytoplasm of the outermost cell layer. Therefore, ablation can only occur at the surface. Non-invasive intratissue ablation is impossible with excimer lasers. Ablation of in-depth tissue, such as the stroma underneath the epithelial layer, first requires the removal of the superficial layer. This can be done either (i) optically with the same laser source, (ii) mechanically with a microtome, or (iii) by mechanical manipulation after chemical disconnection of the epithelial layer from the stroma. Photorefractive keratectomy (PRK)3 removes the epithelial layer optically by subsequent layer-by-layer UV photoablation, starting from the outermost layer. After the epithelium has been removed, photoablation of the stroma can be performed. PRK has the disadvantage of a long healing period and relatively severe postoperative pain.
The laser-assisted in situ keratomileusis (LASIK) method implies UV laser ablation of corneal stromal structures after partial removal of the epithelium by means of a microtome.4,5 A corneal flap is produced. After the laser procedure, the flap is relocated, and it covers the stroma again. Although the epithelium and Bowman’s layer can be preserved, LASIK faces problems due to microtome-related complications and the weak binding forces between the flap and the stroma. The healing process is much faster than with PRK. When LASEK (laser epithelial keratomileusis) is being used,6 excimer laser treatment follows the removal of the epithelium, using a combined knife technique and 20% alcohol. The postoperative pain level is between those of PRK and LASIK. Although millions of excimer laser treatments on human cornea have been performed within the last couple of years, the application of high-energy UV photons and their possible long-term harmful effects are still under discussion. A haze of unknown origin is frequently observed within a few years of corneal surgery with the ArF excimer laser, after PRK and LASEK for the treatment of higher corrections.2 A new direction in laser corneal surgery implies the use of ultrashort wavelength lasers in the visible (VIS) and near infrared (NIR) spectral range. Due to the high light penetration depth in this wavelength region, non-invasive intratissue surgery becomes possible. Laser beams which are focussed into the tissue are required in order to confine the surgical effect to the intratissue region of interest (Fig. 1). Clinical studies have been performed with nanosecond, picosecond (ps), and femtosecond (fs) laser pulses. When using the pulsed Nd:YLF laser at 1053 nm with a 30-60 ps pulse width, better results
Address for correspondence: Karsten König, PhD, Center for Lasermicroscopy, Faculty of Medicine, Friedrich Schiller University Jena, Teichgraben 7, D-07743 Jena, Germany. www.mti.uni-jena.de/clm Lasers in Ophthalmology – Basic, Diagnostic and Surgical Aspects, pp. 79–89 edited by F. Fankhauser and S. Kwasniewska © 2003 Kugler Publications, The Hague, The Netherlands
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Fig. 1. Principle of intrastromal surgery using VIS/NIR laser beams focused to an intratissue region of interest.
were obtained than when ns laser pulses were used. However, no fully contiguous intrastromal effect could be obtained.7,8 Femtosecond lasers work much more precisely than ps and ns lasers. Using femtosecond laser pulses, precise intrastromal cuts in animal cadaver eyes and human eyes can be obtained.9-16 The American company InterLase Inc. provides a femtosecond laser system (INTRALASE™ FS, Intralase, Irvine) operating with a 3-µm spot at 1053 nm for the optical production of flaps, based on photodisruptive effects.17 However, destructive collateral effects have been reported, based on the formation of large gas bubbles and their strong disruptive mechanical effects. Bubble
diameters of more than 12 µm have been hypothesized.15 Also, strong destructive mechanical shock waves can occur. A third side-effect is uncontrolled beam propagation towards the inner eye by a nonlinear effect known as self-focussing. All these destructive collateral effects have been observed with amplified femtosecond laser systems providing high laser peak energies in the range of microjoules (µJ) or even millijoules (mJ). The intensity of these destructive effects changes in parallel with the pulse energy. If it were possible to use less strong pulses at lower pulse energy, by maintaining the ability to ablate corneal tissue, this collateral damage should be reduced or even avoided. In this chapter, we report on the use of nanojoule (nJ) and sub-nJ NIR femtosecond laser pulses of compact non-amplified laser sources for precise nanoprocessing of biological structures without significant collateral damage. Using NIR femtosecond laser pulses at low subnJ pulse energy, we were able to achieve cut sizes of sub-200 nm in biological structures (Fig. 2).18 A minimum full-width half-maximum (FWHM) cut size of 85 nm was determined after femtosecond laser treatment of human chromosome 1. This amazing low cut size is even below the diffraction limited laser spot size. Multiphoton effects which occur in the central part of the spot only make such cuts and holes possible below the size of the illumination spot. Therefore, NIR femtosecond lasers enable similar or better nanoprocessing of cells and tissues than UV lasers, although the diffraction-limited spot size at 800 nm is a factor of four larger than the value for a high quality 193-nm laser beam (spot size changes with wavelength). Using nJ and sub-nJ NIR femtosecond pulses, it was possible to optically knock out single organelles and the dissection of chromosomes within living animal cells, without collateral damage to the surround-
Fig. 2. Highly-focussed NIR femtosecond laser beams of sub-nJ pulse energy and 80 MHz repetition frequency allow precise nanoprocessing of biological structures. Laser-induced, sub-200-nm cuts and holes produced in human chromosomes are depicted.
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High-resolution multiphoton imaging and nanosurgery of the cornea ing organelles or cellular membranes.19,20 We were also able to drill transient holes in cell membranes, and to realize optical gene transfer and efficient targeted transfection with foreign DNA.21 After modification of our femtosecond laser system, this nanotechnology was used for the first time to realize precise femtosecond laser surgery of corneal structures with an intrastromal cut size in the sub-micron range.22 But not just ultraprecise surgical corneal procedures can be performed with NIR femtosecond laser pulses of low pulse energy, we also used the same system to realize three-dimensional (3D) optical tomography of the cornea, with subcellular resolution based on multiphoton-excited autofluorescence and other non-linear effects. In contrast to excimer lasers where ablation occurs layer-by-layer and the ablative effect is confined to the outermost structure, intratissue femtosecond laser procedures require sophisticated monitoring systems in order to make sure of the right location of the focal volume at the target structure of interest. Such a system can be based on optical coherence tomography (OCT), however, the typical spatial resolution is in the range of 10-100 µm. Better subcellular resolution can be obtained with confocal scanning microscopes based on reflected backscattered light (for a review, see, Koester,23 this volume). In order to obtain sufficient depth resolution, outof-focus photons have to be suppressed. This can be done by the introduction of a spatial filter (pinhole). By changing the focal plane, optical sectioning (optical tomography) can be performed. A major advantage of this confocal imaging technique is the possibility of using low-cost NIR laser diodes as light sources. Disadvantages include the use of a precise adjusted pinhole, low photon collection efficiency, artifacts due to the detection of some scattered outof-focus photons, and limitation to signals due to changes in the refractive index. Today’s laser surgical systems have no integrated high-resolution 3D monitoring devices. In this chapter, we also report on a pinhole-free method of optical sectioning based on multiphoton excitation of endogenous biomolecules. The use of NIR femtosecond laser pulses allows the detection of fluorescence photons emitted from endogenous fluorophores, such as the coenzyme NAD(P)H which emits in the blue/green spectral range. Collagen can also be selectively imaged. This multiphoton imaging method with a spatial resolution of 1 µm or better, provides different information than confocal reflection microscopy. In particular, multiphoton imaging allows highly sensitive functional imaging. NAD(P)H, which is mainly located in mitochondria, acts as a sensitive bioindicator because only the reduced form of this coenzyme is fluorescent (neither NAD nor NADP are fluorescent).24 In general, fluorescence detection is a very sensitive method compared to the detection of transmitted or reflected photons. In principle, single fluorescence photons can be detected.
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We used our 3D diagnostic tool for multiphoton imaging with subcellular resolution to determine the target of interest for nano-/microsurgery as well as to control the surgical procedure. In particular, NAD(P)H and collagen have been imaged. After localization of the intratissue target by multiphoton imaging, laser ablation was performed with the same system. Immediately after laser treatment, optical sectioning was carried out in order to ascertain the surgical effect. Femtosecond laser systems Ultrafast laser systems emitting femtosecond pulses are considered to be the highly-precise surgical and diagnostic laser tools of the future. Potential fields of application for femtosecond laser surgery include treatment of the cornea and other ocular structures, the teeth, inner ear, and brain. The first clinical studies with femtosecond laser pulses have been conducted for the optical production of corneal flaps in LASIK.16 In a variety of experiments in material sciences, it was shown that femtosecond lasers may generate a more precise cut or hole in a variety of materials than nanosecond or picosecond laser pulses. So far, the shortest laser pulse produced is of the order of attoseconds (1 as = 10-18 s).25 Attosecond laser pulses have been realized in the X-ray spectral range. The shortest pulses obtained in the NIR spectral range are about 3 fs (3 × 10-15 s). In this chapter, we report on applications of ultrashort pulses in the range of about 100 fs. During 100 fs, light travels a distance of 30 µm in air, within the cornea (n = 1.4) 21 µm. Ultrashort laser pulses are not monochromatic. For example, the spectral bandwidth of a 50-fs pulse is about 19 nm. This broad spectrum has to be considered in the choice of coated optics. Due to optical dispersion effects, the pulse width of the femtosecond laser pulse will be increased during transmission through glass. For example, the transmission of a 100-fs laser pulse through a typical microscope results in a pulse width at the target of about 170 fs, whereas an incident 10-fs pulse will have a pulse width at the target of more than 1 ps. The additional use of pulse compression units can realize a pulse width at the target which is nearly as good as at the laser output. Femtosecond lasers as sources of pulsed infrared radiation belong to high-risk class IV. The most popular femtosecond laser is a modelocked solid state laser with a titanium (Ti)-doped sapphire as the laser medium. The Ti3+ ion is responsible for the action of the laser in the NIR spectral range. These lasers are capable of producing pulses of 100 fs or less at a high repetition rate of about 80 MHz (80 million pulses per second). Recent developments include turn-key, compact, sealed tuneable Ti:sapphire laser systems which include the pump laser system (diode pumped frequency doubled neodymium:yttrium vanadate Nd:
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Fig. 3. Pulse train of a femtosecond laser at 80 MHz with 100fs pulse width and 1-W mean power. L: length of laser cavity, c: velocity of light.
YVO4 cw laser at 532 nm) and the Ti:sapphire laser oscillator in one box. Ti:sapphire lasers can be tuned over a wavelength range of between about 690 and 1060 nm. For example, the compact system
MaiTai (Spectra Physics Inc., USA) has a typical tuning range of 780-920 nm, the Chameleon system (Coherent Inc., USA) of 720-930 nm. In a non-amplified, 1-W, mode-locked Ti:sapphire laser system, the output pulse width is in the range of 60-100 fs and the repetition rate is about 80 MHz (Fig. 3). The pulse energy is 1 W/80 MHz = 12.5 nJ. Further developments in femtosecond laser technology include the use of other laser materials such as Cr:LiSaF, neodymium-doped yttrium lithium fluoride (Nd:YLF, emission at 1047 nm), and the construction of frequency-doubled fiber lasers at 780 nm. So far, material processing including cornea surgery has been conducted with high-pulse-energy amplified femtosecond pulses in the µJ to mJ range, and µm-sized illumination spots. In order to obtain such high energy pulses, the output beam of the Ti:sapphire laser oscillator has to be amplified. A typical amplified femtosecond laser system consists of the pump-laser unit, laser oscillator, a pulse stretcher to transform the pulses into picosecond pulses that do not harm the amplifier material, the amplifier with additional pump source unit, and a pulse compressor. The repetition frequency of amplified
Fig. 4. Principle scheme of the commercial ultrafast amplifier Hurricane system from Spectra Physics. The 80-MHz MaiTai, which is pumped by a laser-diode bar-driven green cw laser, serves as the seed laser. The output beam of sub-100-fs laser pulses is introduced into a compact pulse stretcher in order to obtain picosecond pulses. These pulses are amplified within the regenerative amplifier, which is pumped by a diode-pumped Q-switched frequency-doubled Nd:YLF laser. Finally, the amplified picosecond pulses are compressed to obtain amplified sub-130-fs laser pulses of 1 mJ pulse energy at 1 kHz repetition rate. The standard wavelength is 800 nm. The system is not tuneable.
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systems is in the order of kHz (Fig. 4). Amplified laser systems are complex, expensive, and their use requires a laser expert. In this chapter, we focus on non-amplified 80 MHz fs laser systems emitting picojoule (pJ) and nanojoule (nJ) laser pulses. Multiphoton effects Nano- and microsurgery, as well as fluorescence diagnostics of cells and tissues with NIR femtosecond laser systems, are based on multiphoton effects. Multiphoton effects, in particular two-photon absorption, were predicted in 1931 by a young woman, Maria Göppert-Mayer, in her doctoral dissertation (Fig. 5).26 Her supervisor was Max Born. GöppertMayer hypothesized that one molecule can simultaneously absorb two photons (two-photon absorption) within a short temporal window (100 as). This means that two low-energy NIR photons can be absorbed by the molecule, and can lead to an excited state that would normally require high-energy UV or blue photons. From this excited state, the molecule can emit fluorescence in the visible range. Therefore, NIR photons can induce visible fluorescence in the blue, green, yellow, and red spectral ranges (Fig. 6). Because the process of fluorescence depends on absorbed photon pairs, there is a nonlinear, squared dependence between fluorescence intensity and incident light intensity or laser power, respectively. Göppert-Mayer’s, the later Nobel prize laureate, two-
Fig. 5. Photograph of Maria Göppert-Mayer, who predicted multiphoton processes in her PhD thesis. (Source: University Goettingen)
Fig. 6. Scheme of two-photon excited fluorescence, multiphoton-induced optical breakdown, and formation of second harmonic generation (SHG).
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photon hypothesis was in direct contrast to the dogma that the fluorophore has to emit at a longer wavelength than the fluorescence excitation light. It took 30 years to confirm her ideas, due to the low probability of such a quantum event. Photon fluxes of about 1024 photons per second and per square centimeter are required, which could not be provided until the laser was invented in 1960. In 1961, Kaiser and Garret reported first two-photon excited fluorescence using CaF2:Eu2+ crystals.27 Another two-photon effect, the second-harmonic generation (SHG) was found by Franken et al.28 SHG occurs in certain non-centrosymmetric molecules and produces light at exactly half the incident laser light (Fig. 6). In contrast to fluorescence light, SHG occurs immediately (within femtoseconds compared to nanoseconds), it has a low spectral bandwidth and the same direction as the incident light. In 1970, Rentzepis et al. produced images of threephoton excited fluorescence in organic dyes.29 Based on these studies, Sheppard and Kompfner suggested the utilization of multiphoton-induced fluorescence and harmonic generation for a new type of microscopy, nonlinear microscopy.30 In 1978, they wrote: “In the scanning optical microscope, nonlinear interactions are expected to occur between the object and highly focused beam of light, which we hope will open new ways of studying matter in microscopic detail hitherto not available.” Their idea was to use highly focused laser beams which provide the required high-photon flux density in the focal volume. By means of a scanning unit (flying-spot-technology), nonlinear excitation within the focal volume can be used to probe the target in three dimensions. In 1990, Denk et al. from Cornell University produced the first multiphoton microscope for applications in life sciences, based on femtosecond laser pulses from a dye laser at high repetition frequency.31 Their studies on two-photon excited fluorescence revolutionized fluorescence microscopic imaging. Two- and three-photon excitation of fluorescence in NIR require MW/cm2 and GW/cm2 laser intensities which are produced within a sub-femtoliter focal volume of a high numerical aperture objective. Using 80 MHz sub-200 fs laser pulses, mean powers at the target of 10 mW or less of the tightly focussed laser beam are sufficient. The probability of twophoton excitation decreases with the forth power of the distance from the focus. There is no out-of-focus nonlinear effect, such as generation of VIS fluorescence or photodamage. This is in contrast to conventional one-photon imaging in which reflected photons and fluorescence arise within the entire illumination cone and in which pinholes are required for optical sectioning in order to suppress out-offocus photons. A major advantage of multiphoton imaging is the possibility of pinhole-free optical sectioning by (i) moving the tiny multiphoton excitation volume across the target of interest, and (ii) the efficient collection of fluorescence and SHG photons with respect to the intratissue position of the excitation volume.
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The efficiency of two-photon excitation and the fluorescence yield emulates the following relationship:31 n ≈ P2α / (τf2) . π2NA4 / (hcλ)2 with n = the number of absorbed photon pairs, P = mean power, α = molecular two-photon absorption coefficient, τ = pulse width, f = repetition frequency, NA = numerical aperture, h: Plack quantum of action, c: velocity of light, and λ = wavelength. Because the fluorescence yield depends on a P2/τ relationship in two-photon microscopy, the efficiency increases for high peak power and low pulse width. Despite the high light intensity, safe imaging can be performed below certain intensity thresholds, as demonstrated on laser-exposed single hamster ovarian cells and their next generations.32 Squirrel et al. imaged whole hamster embryos for 24 hours with a femtosecond laser scanning microscope without any damage to the embryos, in contrast to shortwavelength visible light.33 However, when the light intensity is increased to TW/cm2 light intensities, immediate destructive effects occur. The intensity is high enough to induce optical breakdown and plasma formation. At 800nm laser wavelength, four-photon absorption is sufficient to induce ionization and the formation of quasi-free electrons, leading to optical breakdown and plasma formation in water, and organic molecules (Fig. 6). Within the plasma, material is removed due to high temperatures, and plasma-filled bubbles are formed. However, there is another optomechanical effect causing photodisruption. Photodisruption is based on rapid expansion of the laser-induced plasma with Gigapascal pressures and the development and further collapse of cavitation bubbles, accompanied by the formation of destructive shock waves. So far, plasma-mediated ablation of the cornea by highenergy ultrashort laser pulses is mainly due to photodisruptive effects. Photodisruption also leads to destructive effects outside the illumination spot and causes undesired side-effects. In order to obtain a desired highly-localized destructive effect without significant collateral damage, small bubble diameters and low optomechanical effects in the surroundings are required. The threshold for optical breakdown in water is decreased by a factor of more than 100 when comparing 100-fs with 3-ns pulses accompanied by less transformation into destructive mechanical energy (factor 6).13 Kurtz et al. reported that femtosecond pulses require about one-tenth the energy of picosecond and nanosecond pulses to produce corneal disruption.14 But also the photodisruptive effects of µJ femtosecond laser pulses count as destructive effects outside the illumination volume. In addition to optomechanical effects such as the formation of large cavitation bubbles, gas bubbles, and shock waves, self-focusing effects occur. Self-focusing effects based on laser-induced modulation of the refractive index of the illuminated
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tissue occur at high pulse energy and weak focusing, and lead to uncontrolled cutting effects in the direction of the incident laser light. Self-focusing and photodisruptive effects depend linearly on pulse energy. It would helpful to induce multiphoton effects, including optical breakdown and plasma-mediated material ablation, with femtosecond laser pulses at pulse energies as low as possible, to realize highlylocalized destructive effects within the illumination spot only. Experimental set-up and biological samples Preliminary studies on nanoprocessing and multiphoton imaging of the cornea with pJ and nJ non-amplified femtosecond laser pulses have been conducted with an inverted femtosecond laser scanning microscope (FLSM, JenLab GmbH, Jena, Germany; Fig. 7). This microscope is based on the Zeiss LSM 410 system. The 80-MHz, 1-W beam of a tunable MaiTai laser was introduced into the FLSM via the LasCon interface, which consists of a beam expander, fast shutter, beam attenuator, power control, synchronization unit, and mirrors for beam adjustment. The FLSM contains baseport detectors, such as photomultipliers (PMT) and CCD cameras. Autofluorescence and SHG images were obtained by processing the PMT signals depending on the position of the galvomirrors (Fig. 5). In order to avoid the detection of backscattered laser light, two short-pass filters SP 730 (Chroma Techn., USA) were placed in front of the PMT. In part, the Ti:sapphire Vitesse laser (Coherent Inc., USA) at 800 nm was also used. Transillumination of light from a halogen lamp through the whole ex-vivo porcine eye was monitored on-line with a baseport CCD camera. The laser beam was focussed on its sub-micron diffractionlimited spot size by a 40× objective of 1.3 numerical aperture (NA). The pulse width at the sample was determined to be 170 fs pulses by autocorrelation techniques. For two-photon autofluorescence
Fig. 8. Fresh porcine eyes were chemically marked using AgNO3 (upper part). The eyes were placed in special eye chambers (JenLab GmbH).
imaging, a mean power of 5-10 mW at the target surface was used. For corneal processing, a beam with a mean power of 80 mW was used. This power corresponds to 1 nJ pulse energy. We studied the effect of the ablation of corneal structures on porcine eyes that were placed in special tissue chambers with a 170-µm glass window (MiniCeM-biopsy, JenLab GmbH, Jena). In order to localize the imaged and laser-processed structures after laser treatment, parts of the cornea were marked with silver nitrate (Fig. 8). Multiphoton imaging
Fig. 7. The femtosecond laser scanning microscope from the JenLab GmbH, consisting of the MaiTai compact laser system, an interface between the laser and microscope, and the modified LSM 410 microscope.
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Optical sectioning of corneal structures was performed by beam scanning using fast galvomirrors and a piezo-driven z-stage. A typical scan in one focal plane consisted of 512 × 512 pixels and took one or eight seconds. The illuminated field covered 320 × 320 µm = 0.1 mm2. The intratissue focal plane was varied in z-steps of 1 or 5 µm. Figure 9 demonstrates a typical intracorneal image showing two-photon excited autofluorescence. It is obvious that multiphoton imaging provides subcellular spatial resolution. The fluorescence arises mainly from mitochondria. Using a variety of broadband filters in the detection path, maximum emission was found to be in the blue/green spectral range. Very likely, the origin of this autofluorescence is mainly determined by the reduced coenzyme NAD (P)H. The nuclei are non-fluorescent. Different cells can be clearly differentiated between, as can nuclear
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Fig. 9. High-resolution image of two-photon excited autofluorescence of the cornea at 20-µm tissue depth. The autofluorescence is based on NAD(P)H located in the mitochondria of epithelial cells.
area and cytoplasm. Images from different tissue depths are depicted in Figure 10. In addition, the dependence on excitation wavelength is shown in the same figure. The 12 images in the figure represent only some out of a large stack of images. As shown by these representative optical sections, 3D multiphoton-induced autofluorescence imaging enables the various corneal tissue layers (epithelium, Bowman’s layer, stroma) and individual cells and collagen structures to be distinguished. By tuning the laser wavelength to 750 nm, autofluorescent mitochondria can clearly be seen. The mitochondrial fluorescence became less intense when the laser was tuned to longer excitation wavelengths. This behavior correlates with the expected two-photon fluorescence excitation spectrum of NAD(P)H. Because the onephoton absorption band is around 340 nm, the optimum two-photon excitation wavelength should be 750 nm or shorter. Significant changes of the images at 750, 790, and 830 nm occur in the depth range of the junction between the epithelial layer and stroma, and in deeper tissue. The image at 750 nm represents autofluorescence signals, the 830-nm image mainly SHG radiation, while the image taken at 790 nm exhibits a mixture of fluorescence and SHG signals. The SHG
Fig. 10. Autofluorescence and SHG images at different tissue depths and excitation wavelengths from a stack of images. Images above: 750 nm laser excitation; middle: 790 nm; below: 830 nm.
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signal arises from the non-centromeric molecule collagen. Significant SHG signals cannot be obtained from structures within the epithelial layer. SHG imaging can be used to determine clearly the onset of stromal tissue with an axial resolution of about 1 µm. An SHG signal was found if excitation wavelengths of 790 nm or higher were used. In principle, SHG radiation also occurs at laser wavelengths shorter than 790 nm. However, no SHG signal in the UV can be imaged with this experimental set-up due to the reduced UV transmission of short-pass filters in front of the detector. For that reason, the 750-nm signal represents autofluorescence without contributions from SHG radiation, while the 830-nm signal is mainly influenced by backscattered SHG light from collagen structures at 415 nm. Nano- and microprocessing Autofluorescence and SHG imaging were used to find a particular intratissue target region of interest, and to ‘park’ the laser beam there. In order to perform a precise single intrastromal cut, the mean laser power was increased to 80 mW (1 nJ pulse energy). The intense 80-mW beam was scanned along one 512 pixel line at a typical beam dwell time on a pixel of 4 µs. This corresponds to about 320 pulses and a real exposure time of 320 × 170 fs = 54 ps per pixel. The photomultiplier became saturated (gray level: 255) during the laser exposure, probably due to plasma luminescence. Therefore, recording the PMT signal during laser treatment provides on-line information about multiphoton-induced optical breakdown and plasma formation. Interestingly, seconds and minutes after the laser treatment, the effect of intense laser exposure can also be imaged with high spatial resolution at a lower laser power. An intratissue, highly-fluorescent structure of sub-micron lateral size was formed along the cut. Figure 11 shows multiphoton images of stroma after the laser procedure along five lines. The region of intense laser exposure can clearly be seen by the luminescent lines with a typical dimension of 0.8 ± 0.4 µm. In order to study the femtosecond laser effects more carefully, laser-treated eyes underwent histological examination. Upon analysis of hematoxilin/ eosin (HE)-stained cryosections, ultrathin sections revealed precise intraocular sub-micron cuts without collateral damage (Fig. 12). In part, the cuts even went through single nuclei of individual cells without any further visible damage. Laser scanning microscopy of these sections revealed a cut size of about half a micron. This corresponds to the resolution limit in VIS/NIR light microscopes. In order to obtain more precise information on the cut size and quality of the NIR femtosecond laser procedure, scanning force microscopy and electron microscopy (EM) were conducted. Figures 13 and 14 demonstrate EM images obtained after a laser procedure in which a cube of material was ablated. To obtain such a cube, the bottom was first prepared by scanning a squared region deep into the stroma.
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Fig. 11. Nanoprocessing along line scans. The effect of the surgical procedure can be monitored by multiphoton imaging. After laser treatment, an intratissue highly luminescent structure of sub-micron lateral size is formed along the cut.
Fig. 12. Histological examination of HE-stained cryosections after laser exposure by 488-nm laser scanning microscopy reveals precise sub-micron line cuts. Even a single nucleus at 90-µm tissue depth can be clearly cut through without visible collateral damage (image in the left corner).
Thereafter, the four walls were prepared up to the surface by x,z scans and y,z scans. As can be seen in the figures, material can be removed by relatively precise laser cuts. At higher magnification (Fig. 14), the image demonstrates in detail that collagen fibers can be cut through without any significant, large damage zones. There is a small visible layer along the cut of about 100 nm in size, which could be formed as the result of photothermal or photome-
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Fig. 13. Electron micrographs of laser-treated stromal and epithelial tissue.
Fig. 14. REM images demonstrate clear cuts through collagen fibers.
chanical effects. The origin of these tiny changes in tissue structure still have to be investigated. Nevertheless, the image demonstrates that no significant signs of collateral destructive effects were present. In addition to the removal of material, the successful removal of small intratissue structures, such
Fig. 15. During line scans with a power near the threshold for optical breakdown, some small bubbles with diameters of less than 5 µm can be monitored.
as single intraocular cells, was realized by singlepoint-illumination. During and immediately after laser treatment, the formation, localization and lifespan of intratissue bubbles were monitored by video. Typically, between three and seven bubbles occurred as fluorescent sites along a 320-µm line. Using microsecond beam dwell times per pixel, a maximum mean diameter of 5 µm of intrastromal bubbles with a mean lifespan of 1.8 ± 0.3 s were recorded. The relatively long lifespan corresponds to observations by others who noted long-lived gas-filled bubbles (oxygen, hydrogen, methan) compared to short-lived cavitation bubbles. The possibility to remove intratissue material without destruction of the epithelial surface is demonstrated in Fig. 16. Conclusions These preliminary studies clearly show that nJ NIR femtosecond pulses at TW/cm2 intensities of nonamplified compact MHz lasers have the potential for
Fig. 16. Histological image of an HE-stained cryosection of corneal tissue after laser ablation of intratissue material.
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High-resolution multiphoton imaging and nanosurgery of the cornea highly-precise, intratissue processing. No significant collateral damage by optical or optomechanical effects, such as intratissue self-focusing and photodisruption, were seen, in contrast to amplified NIR femtosecond laser surgery with µJ pulse energies. Using fs 80 MHz laser pulses at pJ pulse energies, high-resolution optical tomography of the cornea can be performed, based on two-photon excited NAD(P)H autofluorescence and SHG formation in the collagen structures. It should be possible to build a compact femtosecond laser system for both diagnostic and therapeutical eye procedures. Further studies, including in-vivo animal studies, still have to be conducted in order to discover the optimum pulse energy, pulse length, and spot size for refractive intracorneal surgery. Acknowledgments We would like to thank Oliver Krauss for his experimental contributions to this chapter. Helmut Hörig performed the EM studies, Mrs Möller and Mrs Hitschke the histology. This work was supported by the German Science Foundation (DFG, KO1361/10-3), the Ministry of Science, Research and Art of the State of Thuringia, and the German Ministry of Science, Research and Technology (BMBFT 01ZZ0105).
References 1. Srinivasan R: Ablation of polymers and biological tissue by ultraviolet lasers. Science 234:559-565, 1986 2. Niemz MH (ed): Laser-Tissue Interaction. Springer 1996 3. Trokel SL, Srinivasan R, Braren B: Excimer laser surgery of the cornea. Am J Ophthalmol 96:710-715, 1983 4. Pallikaris IG, Siganos DS: Excimer laser in situ keratomileusis and photorefractice keratectomy for correction of high myopia. J Refract Surg 10:498-510, 1994 5. Mrochen M, Kaemmerer M, Seiler T: Wavefront-guided laser in situ keratomileusis: early results in three eyes. J Refract Surg 16:116-121, 2000 6. Kornilovsky IM: Clinical results after subepithelial photorefractive keratectomy (LASEK). J Refract Surg 17:S222-223, 2001 7. Ito M, Quantock AJ, Malhan S, Schanzlin DJ, Krueger RR: Picosecond laser in situ keratomileusis with a 1053-nm Nd:YLF laser. J Refract Surg 12:721-728, 1996 8. Gimbel H, Coupland S, Ferensowisc M: Review of intrastromal photorefractive keratectomy with the Nd:YLF laser. Int Ophthalmol Clin 37:95-102, 1997 9. Stern D, Schoenlein RW, Puliafito CA, Dobi ET, Birngruber R, Fujimoto JG: Corneal ablation by nanosecond, picosecond, and femtosecond lasers. Arch Ophthalmol 107:587592, 1989 10. Kurtz RM, Horvath C, Liu HH, Krueger RR, Juhasz T: Lamellar refractive surgery with scanned intrastromal picosecond and femtosecond laser pulses. J Refract Surg 14:541-548, 1998 11. Lubatschowski H, Maatz G, Heisterkamp A, Hetzel U, Drommer W, Welling H, Ertmer W: Applications of ultrashort laser pulses for intrastromal refractive surgery.
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Graefe’s Arch Clin Exp Ophthalmol 238:33-39, 2000 12. Noack J, Vogel A: Laser-induced plasma formation in water at nanosecond to femtosecond time scales: calculation of thresholds, absorption coefficients, and energy density. IEEE J Quantum Electron 35:1156-1167, 1999 13. Vogel A, Noack J, Nahen K, Theisen D, Busch S, Parlitz U, Hammer DX, Noojin D, Rockwell BA, Birngruber B: Energy balance of optical breakdown in water at nanosecond to femtosecond time scales. Appl Phys B 68, 271-280, 1999 14. Kurtz RM, Liu X, Elner VM, Squier JA, Du D, Mourou GA: Plasma-mediated ablation in human cornea as a function of pulse width. J Refract Surg 13:653-658, 1997 15. Juhasz T, Kastis GA, Suarez C, Bor Z, Bron WE: Timeresolved observations of shock waves and cavitation bubbles generated by femtosecond laser pulses in corneal tissue and water. Lasers Surg Med 19:23-31, 1996 16. Juhasz T, Loesel FH, Kurtz RM, Horvath C, Bille JF, Mourou G: Corneal refractive surgery with femtosecond lasers. IEEE J Quantum Electron 5:902-910, 1999 17. www.intralase.com 18. König K, Riemann I, Fritzsche W: Nanodissection of human chromosomes with near-infrared femtosecond laser pulses. Opt Lett 26:819-821, 2001 19. König K, Riemann I, Fischer P, Halbhuber KJ: Intracellular nanosurgery with near infrared femtosecond laser pulses. Cell Mol Biol 45:195-201, 1999 20. König K: Multiphoton microscopy in life sciences. J Microsc 200:83-104, 2000 21. Tirlapur UK, König K: Targeted transfection by femtosecond laser. Nature 418:290-291, 2002 22. König K, Krauss O, Riemann I: Intratissue surgery with 80 MHz nanojoule femtosecond laser pulses in the near infrared. Optics Express 10:171-176, 2002 23. Koester CJ: Confocal microscopy of the eye. In: Fankhauser F, Kwasniewska S (eds) Laser in Ophthalmology: Surgical and Diagnostic Aspects. This volume 24. Schneckenburger H, König K: Fluorescence decay kinetics and imaging of NAD(P)H and flavins as metabolic indicators. Opt Eng 31:1447-1451, 1992 25. Drescher M, Hentschel M, Kienberger R, Uiberacker M, Yakovlev V, Scrinzi A, Westerwalbesloh T, Kleineberg U, Heinzmann U, Krausz F: Time-resolved atomic inner-shell spectroscopy. 26. Göppert-Mayer M: Über Elementarakte mit zwei Quantensprüngen. Ann Phys 9:273-295, 1931 27. Kaiser W, Garret CGB: Two-photon excitation in CaF2: Eu2+. Phys Rev Lett 7:229-231, 1961 28. Franken PA, Hill AE, Peters CW, Weinreich G: Generation of optical harmonics. Phys Rev Lett 7:118-119, 1961 29. Rentzepis PM, Mitschele CJ, Saxmann AC: Measurement of ultrashort laser pulses by three-photon fluorescence. Appl Phys Lett 17:122-124, 1970 30. Sheppard CJR, Kompfner R: Resonant scanning optical microscope. Appl Opt 17:2879-2882, 1978 31. Denk W, Strickler JH, Webb WW: Two-photon laser scanning fluorescence microscopy. Science 248:73-76, 1990 32. König K, Becker TW, Fischer P, Riemann I, Halbhuber KJ: Pulse length dependence of cellular response to intense nearinfrared laser pulses in multiphoton microscopes. Opt Lett 24:113-115, 1999 33. Squirrel JM, Wokosin DL, White JG, Barister BD: Longterm two-photon fluorescence imaging of mammalian embryos without compromising viability. Nature Biotechnol 17:763-767, 1999
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Selective absorption by melanin granules and selective cell targeting Charles P. Lin Wellman Laboratories of Photomedicine, Massachusetts General Hospital, Harvard Medical School, Boston, MA, USA
Keywords: laser trabeculoplasty, laser photocoagulation, cavitation, selective targeting, retinal pigment epithelium, RPE, melanin, melanosome
Introduction The use of short-pulsed lasers in ophthalmology dates back to the early 1960s. Soon after the development of the first ruby laser by Maiman,1 attempts were made to use this new light source for photocoagulation of the retina.2-4 However, the ruby laser did not always create reproducible lesions and was associated with a substantial risk of retinal hemorrhage.4-7 When the argon ion laser became available about a decade later, it replaced the ruby laser as the instrument of choice for retinal photocoagulation because its wavelength was more favorable and because its continuous-wave (cw) output allowed better control of energy delivery. The pulsed ruby laser is no longer in use in ophthalmology today. Several other ophthalmic lasers introduced in the 1970s and 1980s require short pulses to achieve their tissue effects, but they employ wavelengths that fall outside the visible spectrum, and do not require pigment absorption. Intraocular photodisruption using the Q-switched Nd:YAG laser,8,9 for example, is based on the principle of optical breakdown in transparent media,10-12 and relies on a nonlinear rather than a linear absorption mechanism. The fundamental output of Nd:YAG laser at 1064 nm is chosen to minimize absorption by the retina at this wavelength. Photoablation with the excimer laser at 193 nm, on the other hand, uses a wavelength that is strongly absorbed by the cornea and does not penetrate deeper into the eye. In the last few years, there has been renewed interest in using short-pulsed lasers to target pigmented structures in the eye.13-17 Both selective laser trabeculoplasty (SLT, see the chapter by Latina in this volume) and selective RPE laser treatment (SRT, see the chapter by Roider, Brinkmann and Birn-
gruber) employ short pulses of green light, with melanin being the intended absorber. Moreover, both these new procedures aim to achieve cell-selective therapy, with the trabecular meshwork (TM) cells and the retinal pigment epithelial (RPE) cells as their respective targets, while minimizing collateral damage to adjacent tissue. Thus, the SLT targets the TM cells while preserving the underlying collagen beams of the meshwork; similarly, the SRT targets RPE cells while preserving the adjacent photoreceptors.13-17 How can such selectivity be achieved? Given the problems encountered with the pulsed ruby laser in the earlier studies, how is it possible to make reproducible lesions in the TM and the RPE, and confine the damage to the target cells? What prevents gross disruption and creation of hemorrhage? This article focuses on the mechanism of short-pulse laser interaction with pigmented cells and with individual subcellular pigment granules. Better understanding of the cell damage mechanism is important, not only in the development of new laser targeting strategies, but also in setting safety standards for short-pulse laser exposures.18 Absorption of pulse laser radiation by melanin granules The TM cells and RPE cells both contain numerous melanin granules, which are the primary absorbers for visible radiation in their respective parts of the eye.19-21 Energy deposition into the pigment absorbers results in a rise in tissue temperature; the spatial extent of the temperature increase depends on the duration of the laser pulse. Kapany et al. provided the following description on the impact of
Address for correspondence: Charles P. Lin, PhD, Wellman Laboratories of Photomedicine, Massachusetts General Hospital, Harvard Medical School, Boston, MA 02114, USA. e-mail:
[email protected] Lasers in Ophthalmology – Basic, Diagnostic and Surgical Aspects, pp. 91–98 edited by F. Fankhauser and S. Kwasniewska © 2003 Kugler Publications, The Hague, The Netherlands
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ruby laser pulses on the retina, in one of the very early papers written on the subject:4 “During the time (~300 µsec) when light energy is delivered to the fundus, very little energy can escape to the surrounding tissue either by convection within the capillary system or by conduction to surrounding tissue. Because of this, the temperature of the small amount of tissue which absorbs the energy becomes highly elevated. If a large amount of energy is delivered rapidly, tissue fluid is vaporized and the vapor erupts through the retina into the vitreous humor and a bubble is formed. A hemorrhage or a retinal hole can also be caused under these conditions. If more moderate amounts of energy are delivered, the absorbed tissue are not vaporized and some of the stored energy is passed on to the retina causing it to be ‘welded’ to the choroid.” In a well-known paper on selective photothermolysis, Anderson and Parrish gave a more in-depth account of the relationship between the optimum laser pulse duration and the size of the target to be treated.22 To confine the energy strictly to the ~1 µm melanin granules requires a pulse duration of less than 1 µsec. The 200-500-µsec pulse duration of the free running ruby laser, also known as normal mode or long-pulsed ruby laser (i.e., one that is not Qswitched), does indeed allow limited heat diffusion into the surrounding tissue (the distance of heat diffusion on this timescale is approximately 20 µm). Nevertheless, the description by Kapany et al. is essentially correct in that the pulsed ruby laser causes rapid, localized heating of a small amount of tissue, followed by vapor bubble formation leading to mechanical disruption if the local energy density exceeds a certain threshold value. Indeed, the ruby laser is not well-suited for creating thermal lesions (‘retinal weld’) because only a small volume of tissue is heated at low radiant exposures, whereas higher radiant exposures lead to vaporization, disruption, and hemorrhage. With Q-switched (nanosecond) pulses, photomechanical tissue damage becomes even more pronounced. Marshall and Mellerio7 used terms such as ‘blast forces’ and ‘microexplosions’ to describe the histological appearance of retina lesions created by Q-switched ruby laser pulses. Damage always begins at the RPE, but they noted that, “Q-switched laser lesions, even those just above threshold, show a subretinal hemorrhage, while massive preretinal hemorrhage are obtained with only slightly more energy.” Thus, the range between threshold effect and serious hemorrhage was very limited.7 The granular nature of melanin absorption becomes less important when longer interaction times are considered (e.g., thermal effects created by the argon laser) because heat diffusion tends to smooth out any irregularity in the initial temperature distributions. Much of the thermal modelling work has been carried out by treating the RPE as a uniform absorbing layer, but more sophisticated models are available that take into account the subcellular distribution of the absorbing melanin granules.16,20,21,23
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Microcavitation bubble formation around irradiated melanin granules Ophthalmologists are familiar with the cavitation bubble formation that accompanies optical breakdown and plasma generation from the Q-switched Nd:YAG lasers.12 Typically these bubbles expand and collapse so rapidly that they are not visible with the unaided eye. The maximum expansion diameter for a bubble created by 1 mJ of pulse energy is about 1 mm, and the collapse time for such a bubble is about 100 µsec.12 Due to their short lifetime, the cavitation bubbles can only be detected with special techniques, such as high-speed (strobe) photography.12 Because optical breakdown requires a threshold of about 1 mJ for nanosecond pulses, the bubble size cannot be significantly smaller than ~1 mm. (Except under optimum focusing conditions using a high numerical aperture objective lens, bubble size of the order of 100 µm can be produced with microjoule pulse energies.)24 Now consider a 1-µm diameter spherical particle irradiated by a short laser pulse. Assume that the particle absorbs 63% of the photons that fall within the 1 µm circular cross-section of the sphere (this would require an absorption coefficient in excess of 10,000 cm-1). Then the total energy absorbed by the particle, for a radiant exposure of 55 mJ/cm2, is about 0.27 nJ, which is more than one million times smaller than the energy needed to create optical breakdown in the vitreous. Amazingly, a radiant exposure of 55 mJ/cm2 (nanosecond laser pulses at 532 nm) is sufficient to initiate bubble formation around the melanin particles. In other words, with a fraction of a nanojoule of energy, the particle is heated to a temperature high enough to vaporize fluid on its surface. What are the sizes of these bubbles? Since the volume of a bubble is to a first approximation proportional to its energy E, the diameter of the bubble is proportional to E-1/3.12 Compared to a typical bubble induced by optical breakdown (1 mm at 1 mJ), a bubble created from melanin granular absorption of 1/106 the energy should have a diameter that is 102 times smaller, or about 10 µm in size. The lifetime of a 10-µm bubble is estimated to be about 1 µsec based on the Rayleigh formula.25 Experimental evidence of bubble formation around laser-irradiated melanin granules was obtained 26 using a microscope capable of very high-speed image capture with nanosecond time resolution (Fig. 1). The experiment was done as follows:26 A microscope was set up to image individual melanosomes isolated from fresh bovine RPE cells. A laser pulse was split into two pulses. The first one (532 nm) was delivered through the microscope to irradiate the particles. The second pulse was used to generate strobe light (~565 nm) that illuminates the sample at various delay times after the first pulse to produce a high-speed ‘stop-action’ image. The arrival time of the strobe pulse was determined by the length of
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Selective absorption by melanin granules and selective cell targeting
Fig. 1. Nanosecond time-resolved imaging of microcavitation bubble formation around individual laser-heated melanosomes.
an optical delay line and was variable from 160 nsec. This set-up eliminates timing jitter, and the time resolution of this imaging system is determined by the laser pulse duration – nanoseconds with Q-switched lasers and picoseconds with modelocked lasers. In a typical experiment, three images were taken in sequence, as shown in Figure 1. The first image on the left was taken before laser irradiation; three melanosomes were visible in this particular example. The second image (middle) was taken a few nanoseconds after the particles were heated by a 532 nm laser pulse. An expanding bubble can clearly be seen around each particle. Bubbles appear dark in the bright field (trans-illumination) image because the strobe light is scattered out of the collection optics of the microscope by the bubbles. The third image on the right was taken after the bubbles have collapsed, showing apparently intact particles. Bubble dynamics around a single particle were investigated by measuring the intensity of a lowpower helium-neon laser probe beam, focused to a small spot around the microparticle.26 As the bubble grows, the intensity of the forward-transmitted beam is attenuated, while the backscattered signal increases. By using a microscope with three different laser colors (532 nm for irradiation, 565 nm for strobe light, and 633 nm for the probe beam), we are able to create and image the cavitation bubble and detect the bubble dynamics simultaneously. Figure 2
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shows the detected signal in the forward (transmitted) direction when a single melanosome particle was irradiated. At radiant exposure level below cavitation threshold, no signal was detected. (The upward spike at the beginning of each trace was due to leakage of the strong pump pulse radiation into the photodiode detector.) For radiant exposures above cavitation threshold, a transient decrease in the transmitted intensity of the probe beam was observed as the result of attenuation by the bubble. The lifetimes of the bubble obtained from the lower two traces shown in Figure 3 were 580 nsec at 1.4 × threshold and 850 nsec at 2.2 × threshold. Corresponding bubble diameters, obtained from images taken with 125nsec strobe delay, were 5.5 and 7.5 µm, respectively. Both the bubble diameter and the bubble lifetime increased with increasing radiant exposure. The measured values were in the range of 1-10 µm (diameter) and 0.1-1 µsec (lifetime) for radiant exposures that ranged from just above threshold to a few times above threshold. Under these conditions, the particles appeared to stay intact after bubble collapse. A single particle can be irradiated repeatedly, producing a bubble each time, without being destroyed. This observation is consistent with the notion that the bubble comes from fluid vaporization at the surface of the particle (a ‘vapor blanket’) rather than vaporization of the particle itself.27 A bubble that contains hot vapor inside and surrounded by cooler fluid outside is an unstable bubble. The vapor rapidly condenses, leaving an ‘empty’ cavity with low internal pressure, and the bubble must collapse. This is the reason for the short bubble lifetime. The dynamics of bubble expansion and implosion is similar to the cavitation dynamics previously observed with laser-induced breakdown, except on a much smaller scale. This is a good example of cavitation by local energy deposition as defined by Lauterborn.28
Fig. 2. Transient attenuation of the probe beam by a single microbubble created by irradiating an isolated melanin particle. (Reprinted from Lin and Kelly26 by courtesy of the publisher.)
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Fig. 3. Intracellular cavitation bubble formation. A. Cultured TM cells with melanin particles that have been internalized. The ingested particles accumulated in lysosomal compartments in the perinuclear regions. B. Microbubble formation inside a single irradiated TM cell around the internalized particles (image taken 125 nsec after laser pulse). C. Cells regain their shape and appearance after bubble collapse. D. Calcein fluorescence image before laser irradiation. E. Calcein fluorescence image after laser irradiation. x marks the cell that was irradiated, cavitated, and lost viability. F. Nucleus of the dead cell is labelled by ethidium brimide (EB).
The threshold for bubble formation around bovine melanosomes is equal to 55 mJ/cm2 for nanosecond and shorter laser pulses at 532 nm.26 In this regime, the threshold is independent of the pulse duration as long as the pulse duration is shorter than the thermal relaxation time of the particles. As pulse duration increases, becoming comparable then exceeding the thermal relaxation time, the threshold increases, from ~300 mJ/cm2 for 1-µsec pulses, to ~400 mJ/cm2 for 2-µsec pulses, and ~550 mJ/cm2 for 3-µsec pulses.29 For these longer pulse durations, more energy escapes from the particle into the surrounding fluid. Therefore, higher radiant exposures are required to reach the temperature for vaporization. What is the temperature for vaporization at the surface of the melanin particle? Does this happen at 100°C, as with normal boiling of water? The answer is no, because bubble formation requires nucleation. In the absence of nucleation, liquid water can be superheated well beyond 100°C without turning into vapor. The reason is surface tension: excess energy is needed to overcome surface tension when
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creating a new interface in the fluid. Pre-existing bubbles or impurities (nucleation centers) lower this energy requirement by allowing bubbles to grow from the existing surface. The pressure Psur associated with the surface tension σ is inversely proportional to the radius of the surface r according to Psur=2*σ/r.30 In the case of melanin in water, the particle itself is acting both as a heat source and as a nucleation site. For a bubble to grow from an initial diameter of 1 µm (diameter of the nucleation center), it has to overcome a total pressure of about 3.5 atm (the hydrostatic pressure plus Psur). The boiling point of water increases with increasing pressure. At 3.5 atm the vaporization temperature is 137°C. Experimentally the particle temperature has been determined to be about 150°C.29 Intracellular cavitation bubble formation and selective cell killing The preceding section describes bubble formation around isolated melanin particles. What happens
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Selective absorption by melanin granules and selective cell targeting when they are produced in cells that contain these particles? Using cultured TM cells, Latina and Park31 first showed that nanosecond laser irradiation selectively induce lethality in cells containing pigment particles. Is bubble formation the mechanism for cell killing? To answer this question, we performed highspeed imaging of laser-irradiated TM cells to visualize bubble formation in these cells.32 We also added an argon ion laser for fluorescence excitation and an integrating CCD camera for fluorescence detection. This set-up allowed us to irradiate cells, image bubble formation (on the nanosecond timescale), and perform fluorescence cell viability assay all under the same microscope without moving the cells. With high-speed imaging, we observed intracellular microbubble formation around melanin particles that have been ingested by the TM cells. Moreover, in any given cell when bubble was created, the cell lost
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viability. Conversely, when bubble was not produced, the cell survived.32 Selective cell killing Selective cell killing was investigated using mixed TM cell culture, as described by Latina and Park.31 Pigmented and nonpigmented TM cells were grown separately, then mixed together before the experiment. Irradiation of the mixed population at radiant exposure above bubble formation threshold results in killing of every cell that contain melanin, while nearby nonpigmented cells remain viable (Fig. 4). Similar experiments were performed with primary RPE cells using tissue explants from freshly enucleated bovine eyes. Bubble formation in the RPE results in cell death (Fig. 5). For nanosecond and shorter pulses, the threshold for RPE cell death is 55 mJ/
Fig. 4. Selective cell killing experiment in a mixed culture of pigmented and nonpigmented TM cells. The pigmented cells were obtained by incubating TM cells with sepia ink melanin. The ingested particles accumulated in lysosomal compartments in the perinuclear regions. The three images on the left were taken: A. before laser irradiation; B. 225 nsec after irradiation with a 565 nm, 25 nsec laser pulse at 0.2 J/cm2, showing the formation of microbubbles in the perinuclear regions where the particles accumulated; and C. after cavitation bubble collapse. D and E. The fluorescence images were taken with FDA viability probe: D. before, and E. after laser exposure, showing the selective loss of viability in cells which underwent cavitation (marked by x in E). Adjacent cells without particles remained viable. F. Fluorescence image taken after the addition of EB, which stained the nuclei of nonviable cells and confirmed that the membranes of these cells became permeabilized by the cavitation process. All images were taken with the exact same field of view. Bar = 20 µm. (Reprinted from Lin et al.32 by courtesy of the publisher.)
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Fig. 5. Cavitation bubble formation in RPE cells. Left: before laser pulse; middle: high-speed image taken 500 nsec after the cells were irradiated by a nanosecond laser pulse; right: image taken after bubble collapse.
cm2, identical to the threshold for bubble formation around single melanosomes.33 Is this the same kind of bubbles as those implicated in the ruby laser induced retinal hemorrhage in the earlier studies? While Kapany et al. talked about vapor eruption through the retina and into the vitreous humor,4 our microbubbles expand and collapse entirely within the confines of the RPE or TM cells. These short-lived (microsecond) microbubbles are not visible by ophthalmoscopy and were most likely not noticed in the previous studies. In fact, Roider et al.34 investigated bubble formation as a damage mechanism for 200-nsec laser pulses, using ophthalmoscopically visible bubbles as an endpoint. They found the threshold for what they called macroscopic bubble formation (those visible by ophthalmoscopy) to be about ten times higher than the angiographic threshold for retinal damage; the threshold for hemorrhage was even higher. Histological examination of the lesions irradiated at the angiographic threshold show damage located primarily to the RPE, which were often physically separated from the Bruch’s membrane and from the photoreceptor outer segments.34 The radiant exposure at angiographic threshold was 45 mJ/cm2 (ten pulses at 532 nm, 200 nsec pulse duration), a value that is comparable to our measured threshold (55 mJ/cm2) for bubble formation around single melanin granules and for RPE cell death ex vivo. (It should be cautioned, however, that accurate comparison of radiant exposures is very difficult due to the uncertainty in determining the spot size accurately in vivo.) We conclude that the mechanism for RPE damage at the angiographic threshold level is microbubble expansion and collapse within the RPE cells. At this level, neither the lesion nor the bubbles themselves are ophthalmoscopically visible. When the radiant exposure was increased to about ten times the threshold for cavitation, stable gas bubbles were observed after the cavitation bubble collapse. These gas bubbles are the product of ‘rectified diffusion’ that occurs when the lifetime of the cavitation bubble becomes sufficiently long to allow dissolved gas to escape from the fluid and diffuse into the cavity. These stable bubbles are visible by ophthalmoscopy and are what Roider et al. identified as macroscopic bubbles.34 It is clear from the above discussion that different authors mean different things when using the term bubble formation, a term that has appeared in many
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papers on retinal damage.4-7,27,29,32-35 Readers are advised to check the definition and not to assume that all bubbles are the same. Implications for selective laser trabeculoplasty and selective retinal pigment epithelium laser treatment Perhaps the most interesting finding from these studies is the observation that cavitation damage (a kind of microscopic underwater explosion) can be confined to the scale of single cells. Precise localization is possible because of the minute amount energy absorbed by each melanin granule (a fraction of a nanojoule), producing cavitation bubbles that are only a few micrometers in size. Thus, even though the damage process is photomechanical in nature, selective cell targeting can be achieved using short pulsed lasers without causing gross tissue disruption. In principle, the size of the bubble should be ‘selflimiting’, meaning that once a bubble is created, it forms an insulating blanket around the particle (the thermal conductivity for water vapor being much lower than that for liquid water). The vapor blanket should both scatter incoming radiation and limit further heat transfer from the particle to the surrounding fluid. In practice, we have not seen evidence of such a self-limiting process. The bubble size always increased with increasing pulse energy. Consequently, proper setting of laser energy is essential for ensuring that the target cells are successfully treated without creating excess damage to collateral tissue. For SLT, unlike conventional argon laser trabeculoplasty (ALT), there is no visible endpoint. To find the proper energy setting, the current protocol for SLT calls for gradual increase of pulse energy until bubble formation is observed, then backing down until the energy is below the bubble formation threshold.36 What do they mean by bubble formation? This is an example why caution is warranted when encountering the term bubble formation. Surely what they mean here is the stable (gas) bubble formation, because conventional slit-lamp illumination is being used to observe the anterior chamber in this procedure. When the pulse energy is reduced below bubble formation threshold, as instructed by the protocol, we are in the regime of the transient in-
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Selective absorption by melanin granules and selective cell targeting tracellular cavitation bubble formation that is not visible by ophthalmoscopy. This is the proper treatment setting for SLT. An automatic method to monitor the transient cavitation bubble formation can make this process more reliable and less time consuming. It remains an interesting open question whether any biological response can be elicited if the radiant exposure is further reduced to below the threshold for transient intracellular bubble formation (for example, intracellular hot spots can be produced by heating the melanin granules to 120°C without bubble nucleation). For SRT, the requirement for selectivity is probably more stringent than for SLT if we aim to preserve the photoreceptors adjacent to the treated RPE cells. Microsecond pulses were originally chosen15,16 precisely to avoid photomechanical damage, but evidence now points to the possibility of microbubble formation inside the RPE as the potential damage mechanism even for pulse duration lasting a few microseconds.29 If so, monitoring bubble formation can be used as an online control for laser setting to ensure that the damage is confined to the RPE cells. Interesting remaining scientific questions are: (a) at what pulse duration does the crossover from photomechanical to photothermal damage takes place; and (b) what pulse duration offers the best therapeutic bandwidth, defined as the ratio of the threshold for photoreceptor damage to the threshold for selective RPE damage? Resolution of these questions will enable ophthalmologists to answer the most important medical question, that is, the true value of SRT in treating RPE-related disorders, particularly in macular degeneration.
Conclusions A very small amount of energy (less than 1 nJ) deposited in a very small volume (a single melanosome) can produce very high local temperature, which in turn causes a very small underwater explosion (microcavitation) with damage range on the scale of single cells. The subtle damage caused by the microcavitation bubbles are not visible under ophthalmoscopic examination. Had the existance of these transient microbubbles been recognized, the development of cell-selective laser therapies such as SLT and SRT could have come much earlier, perhaps even before the introduction of cw argon laser photocoagulation, using the pulse ruby laser. Acknowledgments I wish to thank Michael W. Kelly, whose PhD thesis research resulted in many of the results presented here, Dr Santiago Sibayan for providing cultured TM cells, and Jan Roegener, Clemens Alt, Ralf Brinkmann, Gereon Huettmann, Reginald Birngruber, Johanne Roider, Franz Hillenkamp, Mark Latina, Rox Anderson, and David Sliney for stimulating discussions. This work was supported by NIH EY12970 and AFOSR F4962000-1-0179.
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References 1. Maiman TH: Stimulated optical radiation in ruby. Nature 187:493-494, 1960 2. Zaret et al: Ocular lesions produced by an optical maser (laser). Science 134:1525, 1961 3. Cambell CJ, Rittler MC, Koester CJ: The optical maser as a retinal coagulator: an evaluation. Trans Am Acad Ophthalmol 67:58, 1963 4. Kapany NS, Peppers NA, Zweng HC, Flocks M: Retinal photocoagulation by lasers. Nature 199:146-149, 1963 5. Marshall J, Mellerio J: Histology of the formation of retinal laser lesions. Exp Eye Res 6:4-9, 1967 6. Marshall J, Mellerio J: Pathological development of retinal laser photocoagulations. Exp Eye Res 6:303-308, 1967 7. Marshall J, Mellerio J: Histology of retinal lesions produced with Q-switched lasers. Exp Eye Res 7:225-230, 1968 8. Fankhauser F, Roussel P, Steffen J, Van der Zypen E, Chrenkova A: Clinical studies on the efficiency of high power laser radiation upon some structures of the anterior segment of the eye: first experiences of the treatment of some pathological conditions of the anterior segment of the human eye by means of a Q-switched laser system. Int Ophthalmol 3:129-139, 1981 9. Fankhauser F, Lortscher H, Van der Zypen E. Clinical studies on high and low power laser radiation upon some structures of the anterior and posterior segments of the eye: experiences in the treatment of some pathological conditions of the anterior and posterior segments of the human eye by means of a Nd:YAG laser, driven at various power levels. Int Ophthalmol 5:15-32, 1982 10. Loertscher HP: Laser-induced breakdown for ophthalmic applications. In: Trokel SL (ed) YAG Laser Ophthalmic Microsurgery, pp 39-66. Norwalk, CN: Appleton-CenturyCrofts 1983 11. Docchio F, Dossi L, Sacchi CA: Q-switched Nd:YAG laser irradiation of the eye and related phenomena: an experimental study. I. Optical breakdown determination for liquids and membranes. Lasers Life Sci 1:87-103, 1986 12. Vogel A, Schweiger R, Frieser A, Asiyo M, Birngruber R: Intraocular Nd:YAG laser surgery: light-tissue interaction, damage range, and reduction of collateral effects. IEEE J Quantem Electr QE-26:2240-2260, 1990 13. Latina MA, Sibayan SA, Shin DH, Noecker RJ, Marcellino G: Q-switched 532-nm Nd:YAG laser trabeculoplasty (selective laser trabeculoplasty): a multicenter, pilot, clinical study. Ophthalmology. 105:2082-2088, 1998 14. Park CH, Latina MA, Schuman JS: Developments in laser trabeculoplasty. Ophthalmic Surg Lasers 31:315-322, 2000 15. Roider J, Michaud NA, Flotte TJ, Birngruber R: Response of the retinal pigment epithelium to selective photocoagulation. Arch Ophthalmol 110:1786-1792, 1992 16. Roider J, Hillenkamp F, Flotte TJ, Birngruber R: Microphotocoagulation: selective effects of repetitive short laser pulses. Proc Nat Acad Sci US 90:8463-8647, 1993 17. Roider J, Brinkmann R, Wirbelauer C, Laqua H, Birngruber R: Retinal sparing by selective retinal pigment epithelial photocoagulation. Arch Ophthalmol 117:1028-1034, 1999 18. Sliney DH, Marshall J: Tissue specific damage to the retinal pigment epithelium: mechanisms and therapeutic implications. Lasers Light Ophthalmol 5:17-28, 1992 19. Wolbarsht ML, Fligsten KE, Hayes JR: Retina: pathology of neodymium and ruby laser burns. Science 150:1453-1454, 1965 20. Hayes JR, Wolbarsht ML: Thermal model for retinal damage induced by pulsed lasers. Aerospace Med 39:474-480, 1968 21. Hansen WP, Fine S: Melanin granule models for pulsed laser induced retinal injury. Appl Opt 7:155-159, 1968
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22. Anderson RR, Parrish JA: Selective laser photothermolysis: precise microsurgery by selective absorption of laser radiation. Science 220:524-527, 1983 23. Thompson CR, Gerstman BS, Jacques SL Rogers ME: Melanin granule model for laser-induced thermal damage in the retina. Bull Math Biol 58:513-553, 1996 24. Venugopalan V, Guerra A 3rd, Nahen K, Vogel A: Role of laser-induced plasma formation in pulsed cellular microsurgery and micromanipulation. Phys Rev Lett 88:78-103, 2002 25. Rayleigh L: On the pressure developed in a liquid during the collapse of a spherical cavity. Phil Mag 34:94-98, 1917 26. Lin CP, Kelly MW: Cavitation and acoustic emission around laser-heated microparticles. Appl Phys Lett 72:2800, 1998 27. Pustovalov VK: Thermal processes under the action of laser radiation pulse on absorbing granules in heterogeneous biotissues. Int J Heat Mass Transfer 36:391-399, 1993 28. Lauterborn W: Cavitation and Inhomogeneities in Underwater Acoustics. Springer-Verlag 1980 29. Brinkmann R, Hüttmann G, Rögener J, Roider J, Birngruber R, Lin CP: Origin of RPE-cell damage by pulsed laser irradiance in the ns to µs time regime. Lasers Surg Med 27:451–464, 2000
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30. Young F: Cavitation. McGraw Hill 1989 31. Latina MA, Park C: Selective targeting of trabecular meshwork cells: in vitro studies of pulsed and CW laser interactions. Exp Eye Res 60:359-371, 1995 32. Lin CP, Kelly MW, Sibayan SA, Latina MA, Anderson RR: Selective cell killing by microparticle absorption of pulsed laser radiation. IEEE J Sel Top Quant Electron 5:963-968, 1999 33. Kelly WM, Lin CP: Microcavitation and cell injury in RPE cells following short-pulsed laser irradiation. Proc SPIE 2975:174-179, 1997 34. Roider J, El Hifnawi E, Birngruber R: Bubble formation as the primary interaction mechanism in retinal laser exposure with 200-nsec laser pulses. Lasers Surg Med 22:240248, 1998 35. Gerstman BS, Thompson CR, Jacques SL, Rogers ME: Laser Induced Bubble Formation in the Retina. Lasers Surg Med 18:10-21, 1996 36. Latina MA, Tumbocon JA: Selective laser trabeculoplasty: a new treatment option for open angle glaucoma. Curr Opin Ophthalmol 13:94-96, 2002
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Mechanisms of short-pulsed plasma-mediated laser ablation and disruption Alfred Vogel Medizinisches Laserzentrum Lübeck, Lübeck, Germany
Keywords: photodisruption, ablation, plasma, optical breakdown, intraocular surgery, Nd:Yag laser, femtosecond laser, ophthalmology
Introduction In the late 1970s, Krasnov introduced the use of focused Q-switched ruby laser pulses for goniopuncture and iridotomy in the treatment of glaucoma,1 and, in 1980, Aron-Rosa et al.2 reported performing posterior capsulotomies following extracapsular cataract surgery employing picosecond (psec) pulse trains generated by a mode-locked Nd:YAG laser. Shortly afterward, following extensive preliminary morphological studies,3 Fankhauser et al. published their first reports on laser surgery in the anterior and posterior sectors of the eye,4,5 which were conducted using pulses from a Q-switched Nd:YAG laser (pulse duration, ≈10 nsec). The possibility of operating surgically by means of ‘photodisruption’ on pigmented and nonpigmented structures without opening the eye aroused great enthusiasm among ophthalmologists,6-9 as well as concerns about safety because of the explosive character of the laser effects and the fact that part of the laser light is transmitted through the laser plasma onto the retina.10,11 For this reason, a vigorous debate developed about the advantages and disadvantages of Q-switched lasers as opposed to mode locked lasers, which produce pulse trains of 7-10 psec pulses separated by 5-8 nanosecond (nsec) intervals. This debate initially relied primarily on the available literature on fundamental physical effects,6,10,12,13 but specific studies were soon undertaken on the thresholds for optical breakdown,14-16 plasma transmission,14,17-20 and mechanical laser effects.21-26 The major inadequacy of many of these studies was that they only dealt with individual aspects of the complex lasertissue interaction. Later, a more consistent picture evolved on the basis of systematic studies that examined the complete sequence of different physical
mechanisms involved in intraocular photodisruption and ablation (plasma formation, shock wave production, cavitation), and the resulting tissue effects. The present article aims at presenting a concise but comprehensive portrait of this picture. Thereby, special emphasis is laid on explaining the physical basis for the fascinating new possibilities that emerged with the availability of compact femtosecond (fsec) laser sources as, for example, intrastromal corneal refractive surgery,27-29 and intracellular surgery.30,31 Kinetics of plasma formation in biological tissues Laser-induced, plasma-mediated ablation, also known as laser-induced breakdown, relies on nonlinear absorption in the target, which is achieved when a material-specific irradiance threshold is exceeded.32-34 In tissues with strong linear absorption, plasma formation can be initiated by thermionic emission of (quasi-)free electrons. In this case, the plasma formation usually ‘shields’ the underlying structures and impedes further energy deposition by linear absorption.35 However, plasma can also be formed in materials that are transparent at low irradiance. This occurs at high irradiance when seed electrons for an ionization avalanche are provided by multiphoton ionization. The process then progresses through an interplay of multiphoton ionization and avalanche ionization of target molecules. Thus, plasma formation provides a unique possibility for the achievement of a highly-localized energy deposition in transparent or low-absorbing materials. Localized energy deposition is achieved by using focused laser radiation because plasma formation is limited to regions where the irradiance is high enough to exceed the threshold for laser-induced breakdown.
Address for correspondence: Alfred Vogel, PhD, Medizinisches Laserzentrum Lübeck, Peter-Monnik-Weg 4, D-23562 Lübeck, Germany. e-mail:
[email protected] Lasers in Ophthalmology – Basic, Diagnostic and Surgical Aspects, pp. 99–113 edited by F. Fankhauser and S. Kwasniewska © 2003 Kugler Publications, The Hague, The Netherlands
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Experimental studies have demonstrated that the optical breakdown threshold in water is similar to that in transparent ocular and other biological media (cornea, vitreous fluid, and saline).15,35 For convenience, we shall therefore focus our attention on the physics of plasma formation in pure water. While optical breakdown in gases leads to the generation of free electrons and ions, electrons in liquids are either bound to a particular molecule or are ‘quasifree’ when they have sufficient kinetic energy to move without being captured by local molecular energy potentials. Thus, transitions between bound and quasi-free states are the equivalent of ionization of molecules in gases. In order to describe the breakdown process in water, we adopt Sacchi’s approach, who proposed to treat water as an amorphous semiconductor with an excitation energy of ∆E = 6.5 eV, corresponding to the transition from the molecular 1b1 orbital into an excitation band.37,38 For simplicity, we use the terms ‘free electrons’ and ‘ionization’ as abbreviations for ‘quasi-free electrons’ and ‘excitation into the conduction band’. The process of plasma formation is schematically depicted in Figure 1. It essentially consists of the formation of free electrons by an interplay of multiphoton ionization and avalanche ionization. The promotion of an electron from the ground state to the valence band requires the energy of two photons for UV wavelengths of up to λ = 383 nm, three photons for wavelengths of up to 574 nm, and four, five, and six photons for wavelengths of up to 766, 958, and 1153 nm, respectively. In pure water, this energy can only be provided when several photons interact simultaneously with a bound electron. The multiphoton ionization rate is proportional to Ik, where I is the laser light irradiance and k the number of photons required for ionization. Once a free electron exists in the medium, it can
absorb photons in a non-resonant process called ‘inverse Bremsstrahlung absorption’ (IBA) in the course of collisions with heavy charged particles (ions or atomic nuclei).39 A third particle (ion/atom) is necessary for conserving energy and momentum during optical absorption. Absorption of the photon increases the kinetic energy of the free electron. After k IBA events, the kinetic energy of the electron exceeds the band gap energy ∆E, and the electron can produce another free electron via impact ionization. After impact ionization, two free electrons with low kinetic energies are available which can again gain energy through IBA. The recurring sequences of IBA events and subsequent impact ionization lead to a rapid growth in the number of free electrons, if the irradiance is sufficient to overcome the losses of free electrons through diffusion out of the focal volume and through recombination. In addition, the energy gain through IBA must be more rapid than energy losses through collisions with heavy particles. Energy is lost because a fraction of the kinetic energy of the electron that is proportional to the ratio of the electron and ion masses is transferred to the ion during each collision. The process involving both IBA and impact ionization is called ‘avalanche’ or ‘cascade’ ionization. At high irradiances, the losses play a minor role, and the cascade ionization rate for a given number density of free electrons is proportional to the irradiance.32 Multiphoton ionization occurs on a time scale of a few femtoseconds, and the multiphoton ionization rate is independent of the number density of free electrons. In contrast, cascade ionization depends on the number density of free electrons at the laser focus and requires a longer time because several consecutive IBA events are necessary for a free electron to acquire the kinetic energy for impact ionization. With an ionization energy of 6.5 eV and a photon energy of, for example, 1.56 eV (corresponding to
Fig. 1. Interplay of multiphoton and avalanche ionization in the process of plasma formation. Avalanche ionization is based on sequences of inverse bremsstrahlung absorption events and impact ionization. (Reproduced from Vogel et al.47 by courtesy of the publisher.)
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Mechanisms of short-pulsed plasma-mediated laser ablation and disruption λ = 800 nm), an electron must undergo at least five
IBA events before it can produce another free electron through impact ionization. As mentioned above, IBA can only occur during collisions of the electrons with heavy particles. In condensed matter, the time τ between collisions is roughly 1 fsec.40 Thus, even at extremely high irradiance where most collisions involve IBA, every doubling of the number of free electrons requires at least 5 fsec. Due to this time constraint, avalanche ionization can contribute significantly to plasma formation for laser pulse durations in the fsec range only after a large number density of free electrons has been provided by multiphoton ionization. Several authors have used rate equations based on the Drude model to describe the temporal evolution of the volumetric density of free-electrons ρ under the influence of the laser radiation and to calculate breakdown thresholds for various laser parameters.32,41-46 The generic form of such a rate equation is dρ = ηmp+ ηcasc ρ – gρ – ηrec ρ 2 dt
(1)
The first two terms on the right hand side of Equation (1) represent the production of free electrons through multiphoton and cascade ionization, respectively. The last two terms describe losses through diffusion of electrons out of the focal volume and recombination. The cascade ionization rate ηcasc and the diffusion loss rate g are proportional to the density of free electrons, g, while the recombination rate ηrec is proportional to ρ 2, as it involves an interaction between two charged particles (an electron-hole pair). A detailed description of the individual terms of Equation (1) has been given by Kennedy41 and Noack and Vogel.46 Several investigations, based on the above rate equation, neglected either multiphoton ionization,32,42 recombination,41,45 or diffusion,44 and all four terms of the rate equation have only been considered in a few publications.43,46,47 While many of the early studies were focused on calculation of the breakdown thresholds, recent numerical simulations also included an analysis of the time evolution of the electron density during the laser pulse, the irradiance dependence of the free-electron density, plasma absorption, and volumetric energy density in the plasma.46,47 The description of the plasma formation process is complicated by the fact that both the refractive index and absorption coefficient depend on irradiance. This leads to a spatial phase modulation of the wave front of the laser beam that depends on the intensity distribution across the beam. This modulation brings about a change in the intensity distribution (self-focusing or defocusing) that, in turn, has an effect on the nonlinear absorption process. The degree of self-focusing increases with decreasing focusing angle and shorter laser pulse dura-
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tion.44,48-50 Self-focusing is more prominent in these conditions as it requires a critical power to be exceeded that is largely independent of the focusing angle or pulse duration used.34 In contrast, optical breakdown requires an irradiance threshold to be surpassed. The power necessary to provide this irradiance becomes larger with larger spot size (i.e., decreasing focusing angle) and decreasing laser pulse duration. Therefore, the optical breakdown threshold is at sufficiently small focusing angles and at short pulse durations larger than the critical power for self-focusing, and the breakdown process is influenced by self-focusing, leading to the formation of plasma filaments.48 Threshold for plasma formation The threshold radiant exposure for breakdown determines the minimum achievable extent of the laser effect used for laser ablation or dissection. From an experimental perspective, the threshold for nsec and psec laser-induced breakdown in aqueous media is defined by the irradiance or radiant exposure leading to the observation of a luminescent plasma at the laser focus.51 With shorter laser pulses, there is no plasma luminescence in the visible region of the spectrum, and breakdown is experimentally detected by the observation of a cavitation bubble in the liquid.46,52 From a theoretical point of view, optical breakdown is identified by the generation of a critical free-electron density between ρcr = 1018 cm-3 and 1021 cm-3.41,43,46,51 A good match between experimental threshold values and theoretical predictions for optical breakdown in water is obtained when critical electron densities of ρcr = 1020 cm-3 for nsec pulses and ρcr = 1021 cm-3 for psec and fsec pulses are assumed.46 The irradiance threshold for plasma generation increases by three orders of magnitude when the laser pulse duration is decreased by six orders of magnitude from the nsec range into the fsec range, as shown in Figure 2. The increase in irradiance is required to compensate for the reduced time available to reach the critical electron density. Remarkably, the threshold radiant exposure decreases by three orders of magnitude for the same decrease of pulse duration. Two reasons are responsible for this decrease: (a) the proportionality of the multiphoton ionization rate with Ik, and (b) the decrease in the plasma energy density with shorter laser pulse durations that is explained further below. For pulse durations in the nsec and psec range, the plasma formation thresholds are considerably reduced when the target has a high linear absorption coefficient, because the seed electrons for avalanche ionization are provided by thermionic emission of free electrons. The threshold for plasma formation in transparent media (water or cornea) using pulses of a few nsec is of the order 100-400 J/cm2.48,53 In contrast, plasma formation for ArF excimer laser
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Fig. 2. Threshold irradiance and radiant exposure for optical breakdown in water versus laser pulse duration. The data are from Noack and Vogel46 and Vogel et al.70 Note that all experimental data were obtained for a focusing angle of around 20° with almost diffraction-limited focusing. Aberrations in the optical delivery system lead to severe errors in the threshold values, even if the focal diameter is determined experimentally.67
Fig. 4. a. Temporal evolution of free electron density during laser irradiation and b. irradiance dependence of the maximal free electron density for λ = 532 nm and tp = 100 fsec. (Reproduced from Vogel et al.47 by courtesy of the publisher.)
Fig. 3. a. Temporal evolution of free electron density during laser irradiation, for 1064-nm wavelength and 6-nsec pulse duration. The time t is normalized with respect to the pulse duration tp. The contribution of multiphoton ionization to the total free-electron density is plotted as a dotted line. b. Irradiance dependence of the maximal free electron density ρmax for the same laser parameters. The irradiance is normalized with respect to the calculated threshold irradiance Irate. The threshold Irate and the corresponding value of ρmax are marked with dotted lines. (Reproduced from Vogel et al.47 by courtesy of the publisher.)
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ablation of skin (λ = 193 nm, tL = 22 nsec) where the linear absorption is very high (µa ≈ 40,000 cm-1),54 was reported for radiant exposures of as small as 0.25 J/cm2,55 and in TEA CO2 laser ablation of skin (λ = 10.6 µm tL = 100 nsec, µa ≈ 500 cm-1), the plasma formation threshold is 12-18 J/cm2.56,57 However, for ultrashort pulse durations of ≈100 fsec, linear absorption of the target is nearly irrelevant, even for linear absorption coefficients of as high as µa ≈ 1000 cm-1.53 Small impurities, as are found in tap water, are irrelevant for pulse durations in the psec range and shorter.41,46 This is because the irradiance necessary to complete the ionization avalanche during the short laser pulses is so high that the initial electrons are readily created by multiphoton ionization, and linear absorption does not result in a lowering of the threshold. The laser pulse duration does not only affect the threshold irradiance, but also the entire dynamics of plasma formation and the irradiance dependence of the free-electron density, as shown in Figures 3 and 4. With nsec pulses in the IR (Fig. 3), no free electrons are formed by impact ionization for irradiance values below the breakdown threshold because no seed electrons created by multiphoton ionization are available. Once the irradiance is sufficiently high to provide a seed electron, the ionization cascade proceeds very rapidly, due to the high irradiance. The electron density increases by nine
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Mechanisms of short-pulsed plasma-mediated laser ablation and disruption orders of magnitude within a small fraction of the laser pulse duration, and actually overshoots the critical electron density of ρcr = 1020 cm-3. This results in an extremely sharp breakdown threshold because either a highly ionized plasma is produced, or no plasma at all. It is important to note that this ‘sharpness’ does not exclude the possibility of pulse-topulse variations of the threshold irradiance. These variations are due to the probabilistic nature of the multiphoton-induced generation of seed electrons. With fsec pulses (Fig. 4), there is no lack of multiphoton-induced seed electrons for avalanche ionization, and the onset of plasma formation is, therefore, deterministic. An avalanche is initiated at irradiance values considerably lower than the breakdown threshold, and the free electron density varies continuously with irradiance. Therefore, it is possible to generate any desired free-electron density by an appropriate irradiance.47 It is interesting to note that, even for fsec plasmas, avalanche ionization is the mechanism that produces the majority of the free electrons during the laser pulse (Fig. 4). Multiphoton ionization dominates during the initial part of the pulse but avalanche ionization takes over later, as its rate depends on both the irradiance and free-electron density, whereas the multiphoton ionization rate only depends on irradiance (Equation 1).
Plasma formation above the breakdown threshold At the breakdown threshold, plasma formation is restricted to the focal region of the laser beam. In contrast, when the laser beam provides an energy in excess of the breakdown threshold and is focused within a transparent medium, the plasma formation is characterized by a growth of the plasma from the beam waist towards the incoming laser beam, as illustrated in Figure 5. Almost no plasma develops behind the laser focus since most of the laser light has already been absorbed prior to and in the beam waist. Thus, the region behind the focus is ‘shielded’ by the plasma absorption.34,48,58-60 A realistic explanation for the expansion of the plasma is provided by the ‘moving breakdown’ model originally proposed by Raizer,61 and further refined by Docchio and coworkers.58,59 In this model, it is assumed that optical breakdown is independent of the preceding plasma formation, and occurs at all locations where the irradiation exceeds the breakdown threshold. As the power increases during the laser pulse, the plasma front moves along the optical axis at the same velocity as the location where the breakdown threshold is exceeded. This process is illustrated in Figure 6. For a Gaussian beam, Docchio and coworkers derived the following prediction for the plasma length zmax from the beam waist towards the incoming laser beam that is reached at the intensity peak of the laser pulse.58
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β –1 zmax = zR cvvvvvvvv
103 (2)
Here zR is the Rayleigh range, and β the ratio between peak irradiance in the laser pulse and the threshold irradiance for laser-induced breakdown, i.e. β = I/Ith. The laser radiation incident on the plasma after the intensity peak of the laser pulse only serves to heat the plasma, but does not elongate it further. The validity of the moving breakdown model was shown experimentally for Nd:YAG laser pulses (λ = 1064 nm) with pulse durations in the nsec and psec range.48,58 The quantitative predictions were very good for psec pulses, but degrade for pulse durations in the nsec range as the assumption of a spatially and temporally constant breakdown threshold during the laser pulse is incorrect in these conditions.34,48 The reason for this is that the UV emission of the plasma contributes to the formation of additional free electrons in the plasma vicinity that act as seed electrons for cascade ionization. For nsec pulses, this leads to a lowering of the breakdown threshold during the laser pulse and, thus, to a much larger plasma size than predicted by the moving breakdown model. At the start of plasma formation, the threshold is determined by the high irradiance needed to generate the initial quasi-free electrons for the ionization avalanche through multiphoton ionization. Later, when seed electrons are provided by the UV plasma luminescence, the breakdown threshold decreases to the irradiance value necessary to reach the critical electron density by avalanche ionization. For shorter pulses in the psec or fsec range, a higher irradiance is required to achieve the critical free-electron density at the end of the ionization avalanche, and the creation of seed electrons for the avalanche through multiphoton ionization does not provide an additional barrier. Therefore, at superthreshold energy, the breakdown threshold remains constant throughout the laser pulse and the predictions of the moving breakdown model hold. For breakdown produced using fsec pulses focused at moderate angles, the plasma length observed above the breakdown threshold is considerably longer than the spatial length of the laser pulse (30 µm for a 100fsec pulse). This indicates that plasma formation starts before the pulse reaches the laser focus. The plasma front moves with the laser pulse toward the focus, so that free electrons remain in its wake.62 This is in contrast to psec and nsec breakdown, where the physical length of the pulse is much longer than the plasma and the plasma front moves from the focus towards the incoming laser beam. The plasma length β – 1, as for psec for fsec pulses is proportional to cvvvvvv pulses, even though the plasma front moves in opposite direction.34 In both cases, the extent of the plasma on the laser side is determined by the maximum axial distance from the beam waist at which the breakdown threshold is exceeded. The situation is completely different for plasmas
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Fig. 5. Plasma shape for different pulse energies above the optical breakdown threshold: a. for 6-nsec pulse duration; b. for 30-psec pulse duration. The plasmas were produced with Nd:YAG laser pulses at a 1064-nm wavelength, using a focusing angle of 22°. The pulse energies are indicated in the frames. The laser light is incident from the right. The scale bar represents a length of 100 µm. The shapes of the plasmas produced by the nsec and psec pulses are different because the laser beam profile for the nsec pulses had an annular structure, while it was Gaussian for the psec pulses. With increasing pulse energy, the plasma center is located further away from the laser focus. No plasma is formed in the region behind the focus because this region is ‘shielded’ by the light absorption within the plasma. (Reproduced from Vogel et al.48 by courtesy of the publisher.)
Fig. 6. a. Temporal evolution of the laser power and b. the plasma contours during the optical breakdown process at superthreshold energies (‘moving breakdown’). When the laser power exceeds the optical breakdown threshold Pth, plasma is formed in the beam waist (1). With increasing laser power, the threshold irradiance is exceeded further upstream in the incoming laser beam (2). The maximum plasma extension is reached when the laser power reaches its maximum (3). While upstream of the beam waist the plasma contours at times (1) to (3) correspond to iso-irradiance lines, this does not hold for the region behind the beam waist because of the light absorption within the plasma (plasma shielding).
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Mechanisms of short-pulsed plasma-mediated laser ablation and disruption formed at free surfaces in air. During the initial phase of the superthreshold laser pulse, plasma formation is restricted to the target material because the breakdown threshold is higher in air than in the tissue.34 When the electron density becomes high enough that the plasma frequency exceeds the frequency of the light (the critical value is 1021 cm-3 for λ = 1064 nm34), the plasma absorption coefficient increases drastically and the plasma becomes highly reflective up to a value of R = 0.9 for very high superthreshold exposures.63-65 This change has two consequences: (a) the plasma electron and energy densities increase very rapidly,66 and (b) plasma formation extends into the surrounding air because hot electrons ejected from the target start to ionize the air. The latter process leads to the development of a plasma plume that largely reduces the amount of laser light reaching the target.35 It is only for fsec laser pulses that the coupling of optical energy into the target is not impaired by plasma shielding, since the laser pulse is too short to allow the formation of a plasma plume during the laser pulse.66 Plasma absorption Plasma absorption determines how much energy is coupled into the target medium and how much is transmitted past the target volume. As well as influencing the efficiency of the laser surgical process, plasma absorption is important for its safety if surgery is performed near sensitive, strongly absorbing biological structures as, for example, the retina. Absorption coefficients of plasmas produced in bulk water have been determined experimentally by measuring the plasma transmission, scattering and reflection, together with the plasma length.34,58 The investigations covered a range of radiant exposures up to 50 times threshold, and yielded values of between 100 and about 400 cm-1, depending on pulse duration (6 nsec and 30 psec), wavelength, and radiant exposure. Slightly higher values of between 100 and about 1000 cm-1 were obtained by numerical calculations of the plasma absorption coefficients at the breakdown threshold that considered the time evolution of the free-electron density and the absorption cross section of free electrons for IBA. These calculations were performed for pulse durations of between 100 fsec and 100 nsec.46 For plasma formation at tissue surfaces in air, no experimental data for plasma absorption coefficients and the spatial distribution of energy deposition are available to date. Once the electron density at the surface starts to exceed the critical value of ≈1021 cm-3, the absorption coefficients will certainly be higher than the values for bulk media. Feit et al.66 assumed that plasma is formed in a layer with a thickness of only a few nanometers, but did not discuss how the plasma electron and energy densities corresponding to such a small absorption depth relate to the ablation thresholds of soft and hard tissues, respectively.
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Plasma energy density Plasma energy density is closely linked to the strength of the mechanical effects (shock waves and cavitation) associated with breakdown. It determines how strongly disruptive the breakdown event is, and how much mechanical damage is caused in the vicinity of the laser focus. The deposition of optical energy into the medium is mediated by the generation and subsequent acceleration of free electrons. The energy gained by the electrons is transferred to heavy plasma particles through collisions and recombination, resulting in a heating of the atomic and ionic plasma constituents. The number of collisions and recombination events and the resulting energy transfer to the medium are proportional to the laser pulse duration. Therefore, the plasma energy density must increase with increasing laser pulse duration. Theoretical predictions of the dependence of plasma energy density on laser pulse duration using Equation (1) are shown in Figure 7.46 For fsec exposures, the laser pulse duration is shorter than the electron cooling and recombination times. Thus, minimal energy is transferred during the pulse, and the energy density deposited into the breakdown region is simply given by the number of free electrons produced, multiplied by the mean energy gain of each electron. For pulse durations longer than the electron cooling time (several psec) and recombination time (several 10 psec), a dynamic equilibrium is established between the energy transfer through collision and recombination and the generation of free electrons by the incident radiation. For pulse durations in the nsec range, the calculated energy density is proportional to the laser pulse duration. Experimental values of the plasma energy density are 33-40 kJ/cm3 for 6-nsec pulses, ≈10 kJ/cm3 for 30-psec pulses, and less than 1 kJ/cm3 for 100fsec pulses.68,69 The model predictions in Figure 7
Fig. 7. Calculated plasma energy density at the optical breakdown threshold versus laser pulse duration. The calculations were performed for a wavelength of 580 nm and a critical electron density of 1020 cm-3, which is realistic for nsec-optical breakdown (see text). For psec and fsec pulses, the critical electron density is approximately 1021 cm-3, and thus the actual plasma energy density is approximately ten times higher than plotted. (Reproduced from Noack and Vogel46 by courtesy of the publisher.)
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agree qualitatively well with the experimental data. However, the calculated value for the 6-nsec pulse duration (150 kJ/cm3) is approximately four times greater than the experimental value because the plasma expansion during the laser pulse is not accounted for in the model. On the other hand, the energy density values predicted for 100-fsec and 30-psec pulses (150 J/cm3 and 550 J/cm3, respectively) are by about one order of magnitude smaller than the experimental values. This is due to the fact that all energy density values were calculated assuming an electron density of 1020 cm-3. A much better agreement with the experimental data is obtained assuming an electron density of 1021 cm-3 for psec and fsec breakdown, which also yields the best match between calculated and experimental values for the breakdown thresholds.46 The above data indicate that, while the threshold energy density for optical breakdown and plasmamediated ablation using nsec pulses is extremely high, for fsec pulses it is smaller than the vaporization enthalpy of water at constant pressure, and resembles the volumetric energy density threshold for ablation based on linear absorption. For plasma-mediated ablation in bulk media at superthreshold energies, the plasma volume grows during the laser pulse, and thus there is little variation in the plasma energy density at superthreshold energies. Details of the dependence of plasma energy density on the laser pulse energy, pulse duration and focusing angle have been analyzed by Vogel et al.34,38,70 The situation differs for plasmas produced with ultrashort laser pulses at surfaces in air. In that case, there is no shielding of the irradiance at the target surface until the breakdown threshold in air has been reached. Thus, the energy density in the superficial plasma layer will increase with growing radiant exposure, due to the increase in IBA events with growing electron density. Thermo-mechanical and chemical plasma effects During laser-induced plasma formation, an extraordinarily high energy density develops in the focal volume within a very short time, in particular for nsec and psec optical breakdown.22,48,68,70-74 Temperature and pressure rise rapidly to very high values, causing an explosive expansion of the laser plasma. This expansion of the plasma leads to the production of a shock wave68,75 and, if the application site is in a fluid environment, to the formation of a cavitation bubble,68,75 as shown in Figures 8 and 9. The high initial plasma pressure results in a very rapid bubble expansion that overshoots the equilibrium state where the internal bubble pressure equals the hydrostatic pressure. When both pressures are equal, the kinetic energy of the fluid proximal to the bubble has reached its maximum and, owing to inertia, the bubble continues to expand radially. The expansion leads to a drop in the internal bubble pressure, and
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the increasing difference between the hydrostatic pressure and the internal bubble pressure decelerates the expansion and brings it to a halt. At this point, all kinetic energy is transformed into the potential energy of the expanded bubble (Fig. 9b). The bubble energy is related to the radius of the bubble at its maximum expansion, Rmax, and the difference between the hydrostatic pressure p0 and the vapor pressure pv inside the bubble by:75 EB =
4π (p – pv)R3max 3 0
(3)
The expanded bubble collapses again, due to the static background fluid pressure. This collapse compresses the bubble content into a very small volume, thus generating a very high pressure that can exceed 1 Gpa.76 The rebound of the compressed bubble interior leads to the emission of a strong stress transient into the surrounding liquid that for approximately spherical bubbles evolves into a shock wave.22,76 While the events during bubble generation are strongly influenced by the laser parameters, the subsequent bubble dynamics are primarily influenced by the boundary conditions in the neighborhood of the laser focus. In a free fluid, a spherical bubble retains its spherical shape while oscillating, and the bubble collapse takes place at the site of bubble formation. Near material boundaries, the collapse is asymmetric and associated with the formation of one or two high-speed water jets, which concentrate the bubble energy at some distance from the locus of bubble generation.76-79 When the bubble collapses in the vicinity of a rigid boundary (such as, for example, an intraocular lens implant), the jet is directed towards this boundary, as illustrated in Figure 9c. During bubble collapse near elastic tissue-like boundaries, bubble splitting and formation of two liquid jets directed away from and towards the boundary have been observed, and velocities of as high as 960 m/sec have been measured for the jet directed towards the boundary.78 A summary of peak shock pressure amplitudes and transduction efficiencies of absorbed optical energy into cavitation bubble energy for different pulse durations is given in Table 1. Vogel et al. presented a complete energy balance for plasma formation in bulk water at nsec to fsec time scales.34,70 They found that the transduction of laser energy into mechanical energy (shock wave and cavitation bubble energy) for nsec pulses is as high as 90%, more than for any other laser-tissue interaction. The ratio of shock wave energy (not included in Table 1) to cavitation bubble energy was ≈2:1 for nsec pulses and ≈3:2 for psec pulses. The explosive expansion of the plasma produces disruptive tissue effects extending far beyond the vaporization and disintegration of tissue that occurs within the plasma volume. In the immediate vicinity of the plasma, the effects of the shock wave and
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Fig. 8. Shock wave emission and cavitation bubble expansion in the initial phase following optical breakdown with a. a 10-mJ, 6nsec Nd:YAG laser pulse (λ = 1064 nm), and b. a 1-mJ, 30-psec pulse. The laser light is incident from the right. The self-luminous plasma is visible in all pictures, regardless of the time at which the shock wave and cavitation bubble were illuminated. The scale bar represents a length of 100 µm. (Reproduced from Vogel et al.68 by courtesy of the publisher.)
Fig. 9. Cavitation bubble dynamics produced by focusing a 5-mJ, 6-nsec Nd:YAG laser pulse into water. The laser light was incident from the right. a. Photograph shows the self-luminous plasma, the shock wave and the emerging bubble 90 nsec after the laser pulse. b. Picture taken 130 µsec after the laser pulse when the cavitation bubble had reached its maximum size. c. During bubble collapse in the vicinity of a solid boundary (located just below the bottom of the picture), a high-speed liquid jet is formed. The jet becomes visible during the rebound oscillation of the bubble following its collapse. The picture was taken 50 µsec after the bubble had collapsed. Jet formation concentrates energy at some distance from the optical breakdown site and is thus a potential source of collateral damage. (Reproduced from Vogel et al.33,80 by courtesy of the publishers.)
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Table 1. Dependence of shock wave pressure and cavitation bubble energy on laser pulse duration Pulse duration
Wavelength (nm)
E/Eth
Pressure at plasma boundary (GPa)
Pressure at a distance of 12 mm (MPa)
Degree of conversion of absorbed light energy into bubble energy (%)
76 6 30 3 300 100
750 1064 532 580 580 580
6 60 60 60 60 60
10 7-30 1.7-10 2.2 1.8 0.9
4.0 3.0 0.65 0.23 0.11 0.06
22.0 22.5 13.5 11.0 3.0 3.0
nsec nsec psec psec fsec fsec
Data from Vogel34 and Vogel et al.68
cavitation bubble expansion can hardly be distinguished, but at a somewhat larger distance, cavitation effects are without doubt responsible for the creation of morphologically-identifiable tissue alterations.36,68,71-74,81-83 The effects of the plasma and bubble expansion strongly depend on the location of the plasma in the tissue. When the plasma is formed in the bulk of the tissue, all deposited energy above the vaporization threshold acts to deform the surrounding tissue.71,84 However, when the laser pulse is focused on a tissue surface in a liquid environment, a large fraction of the deposited energy is imparted to the surrounding fluid, and the hole created in the tissue is only slightly greater than the diameter of the laser focus.34,71 Nevertheless, the inertial confinement of the vaporized material by the surrounding fluid causes a distinct indentation of the tissue surface during the expansion of the cavitation bubble. Collateral effects are much less severe when the plasma is produced at a tissue surface in air where the plasma expansion is not mechanically confined.84,85 As well as the cavitation bubble expansion, the jet formation during bubble collapse is another potent cause of far-reaching tissue effects because the jet concentrates energy at locations away from the site of plasma formation. The jets have been shown to cause collateral damage in photodisruption71 and pulsed laser ablation,79 and to increase the amount of material removed.78,79,86 The reduction of plasma energy density with decreasing pulse duration explains the strong reduction of mechanical effects produced with ultrashort, as opposed to nsec, laser pulses (Table 1). The ratio of mechanical energy Emech to the energy fraction consumed for vaporization of the fluid within the plasma volume Evap can serve as a metric for the strength of the disruptive effects accompanying plasma-mediated ablation. The ratio (Emech/Evap) for breakdown in water was found to decrease from 12:1 to 1:2 when the pulse duration was reduced from 6 nsec to 100 fsec.70 The drop in plasma energy density with pulse duration also explains the decrease observed in plasma luminescence, which is no longer visible for pulse durations ≤ 3 psec.62,69 The pressure of the shock waves emitted from the optical breakdown site decreases with decreasing
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pulse duration34,68,69,72,73 (Table 1), but not as strongly as the reduction in plasma energy density. The reason for this phenomenon is that psec and fsec plasmas are always produced under stress confinement. While the pressure induced by a phase transition in the plasma volume is small for low energy densities, the thermoelastic stresses are still very high. Numerical simulations for plasmas with a free-electron density of ρ = 1021 cm-3 predicted a temperature rise of 274°C, which is accompanied by the generation of a thermoelastic stress wave with 2.4 kbar compressive amplitude outside the laser focus,47 which is much higher than the saturated vapor pressure for this temperature. The tensile component of the thermoelastic stress wave determines the threshold for bubble formation at the laser focus. We can conclude that the mechanical effects at the threshold for plasma-mediated soft-tissue ablation with fsec pulses are comparable with those accompanying stressconfined linear absorption. Although high temperatures of several thousand Kelvin are reached within nsec and psec plasmas,87,88 a thermally-modified zone less than 0.2 µm thick was found at the rim of ablation craters in corneal tissue, regardless of pulse duration.71,89 The sharp delineation of the heated zone is due to the sharpness of the plasma boundary, owing to the nonlinearity of its formation and to the short time available for heat diffusion out of the heated volume that is limited by rapid adiabatic cooling during the cavitation bubble expansion.71 Within the plasma, thermal degradation of the tissue is accompanied by chemical dissociation induced by the interaction of the free electrons with the biomolecules and water. Electrons with energies below 15 eV can initiate fragmentation of biomolecules via attachment of the incident electron. The electron attachment leads to the formation of a resonance, namely a transient molecular anion state.47,90,91 For a molecule XY, this process corresponds to XY + e- → XY*-, where the XY*- has a repulsive potential along the X-Y bond coordinate. The molecular anionic state can decay by electron detachment (leaving a vibrationally excited molecule) or by molecular dissociation along one or several specific bonds, such as XY*- → X• + Y-. Nikogosyan et al.92 describe the formation of OH*
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Mechanisms of short-pulsed plasma-mediated laser ablation and disruption and H2O2 through various pathways following ionization and dissociation of water molecules. Both oxygen species are highly reactive and known to cause cell damage.93 The dissociation of water molecules and free radical formation during nsec and fsec laser-induced plasma formation has been confirmed in several experimental studies.94,95 While the chemical processes within the breakdown region have little practical relevance for laser parameters producing a high plasma energy density, where the tissue effects are dominated by the thermomechanical effects, they are of major importance for fsec plasmas with energy densities below the threshold for bubble formation. Such low-density plasmas are possible because of the gradual increase of the free-electron density with irradiance at ultrashort pulse durations (Fig. 4). They enable highly-localized, chemically-mediated ablation or dissection processes with little or no thermomechanical contribution.47
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a
b
Implications for tissue ablation and intraocular photodisruption The use of plasma-mediated ablation and disruption that are based on nonlinear absorption enables the performance of surgery inside of transparent biological structures.7,8,27-31,34 Owing to the high energy density at nsec and psec durations, and the inertial confinement inside transparent structures, the precision of plasma-mediated effects is generally compromised by cavitation effects. Therefore, plasma-mediated effects in bulk tissue are better-suited for cutting and disruption than for ablation of large tissue volumes with sharply delineated boundaries.81 The precision of plasmamediated cutting can be optimized by various measures: 1. The use of a clean beam profile that provides the minimum possible spot size at any given focusing angle. 2. The use of the largest possible focusing angle for each application because a large focusing angle guarantees a small focal spot both in lateral and axial direction. 3. Minimization of the aberrations in the optical delivery system, including the contact lenses used for intraocular applications.67,96 4. The use of ultrashort pulse durations to exploit the decrease of the optical breakdown threshold and the minimization of mechanical side-effects observed with decreasing pulse duration. 5. When pulse series are applied, it must be prevented that plasma production is hindered by the cavitation pulses from previous pulses. Hindrance of the plasma production arises when the time between pulses is shorter than the bubble lifetime, for example, when highrepetition rate bursts are used for iridotomies,97 or in intrastromal corneal surgery for refractive corrections. In the latter case, this can be partially circumvented by the use of irradiation strategies that avoid fusion of the microbubbles produced by individual laser pulses to large bubbles that affect the beam path of subsequent laser pulses.29
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c Fig. 10. Refinement of plasma-mediated tissue effects by the use of shorter laser pulses. a. Cut in Descemet’s membrane in the cornea produced by 6-nsec laser pulses of 1064 nm wavelength. b. Cut produced by 30-psec laser pulses. The scale bars represent a length of 100 µm. (Reproduced from Vogel et al.36 by courtesy of the publisher.) c. Lenticule dissected out of the corneal stroma using 110-fsec pulses of 780 nm wavelength. (Reproduced from Lubatschowski et al.29 by courtesy of the publisher.)
The refinement of tissue effects with decreasing pulse duration has been demonstrated in various experimental studies,28,29,36,84,98,99 and some clinical studies,82,100 and is illustrated in Figure 10. When ultrashort laser pulses are applied at a very large numerical aperture, it is possible to perform intranuclear chromosome dissection in living cells.30 In this procedure, the material removal per pulse is less than 0.1 µm3. On the other hand, standard ophthalmic applications, such as iridotomies and capsulotomies, often take advantage of the disruptive mechanical effects accompanying optical breakdown, and are thus mostly performed with nsec pulses.7 Plasma-mediated processes have also been applied
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for material removal from soft and hard tissue surfaces in air.29,89,101-104 We showed above that the plasma absorption coefficients in bulk aqueous media are two orders of magnitude smaller than the linear tissue absorption coefficients for ArF excimer or Er:YAG laser radiation that are often used for tissue ablation at surfaces. Therefore, the desired precision of plasma-mediated ablation must be controlled by an appropriate choice of focusing angle and pulse energy. Both parameters determine the growth of the plasma from the beam waist into the cone angle of the incoming laser beam. The minimal ablation depth for fsec ablation of soft tissues was found to be relatively large (1.5-200 µm) when the laser radiation was focused to a spot size on the order of 20 µm.74,89,103-105 While the ablation depth could probably be reduced by using a smaller spot size (i.e., a larger focusing angle) and very small pulse energies, this would involve a considerably longer processing time. Moreover, such a procedure would require very precise tracking of the tissue surface in the axial direction of the laser beam, because the focal spot must be located exactly at the tissue surface in order to achieve precise ablation. When the laser beam is only moderately focused, a tissue layer of up to 200 µm in thickness is ejected by the expansion of the plasma formed in the beam waist.74,104,105 Because mechanical ejection is involved in the ablation process, the ablation depth depends on the ultimate tensile strength of the tissue. At equal radiant exposure and pulse duration, much smaller ablation rates have been observed for the mechanically resistant corneal stroma (1.5-10 µm with 30-psec pulses89 and 3-16 µm with 140-fsec pulses103) than for the much weaker neural tissue (50200 µm with 30-psec pulses105, 20-200 µm with 140fsec pulses104). The ablation efficiency achieved using fsec pulses is larger than when using psec and nsec pulses.103,104 This trend can be explained by the decrease of plasma energy density with shorter pulse durations. The nonlinear dependence of the plasma energy density with irradiance (Figs. 3 and 4) enables a large ablation depth to be achieved with little thermal damage to the residual tissue. Thermal side-effects are almost negligible with single laser exposures, regardless of the laser pulse duration.71,89,104 Even with nsec pulses, the thickness of the thermallydamaged layer in corneal tissue remains < 1 µm.71 However, when ultrashort laser pulses are applied at high repetition rates, the residual heat remaining in the nonablated tissue may accumulate and lead to larger zones of thermal damage if the laser beam is applied to a large spot and not laterally scanned during ablation.66,102,106 This phenomenon is due to the continuous dependence of free-electron density and energy density on irradiance in the fsec range (Fig. 4). Therefore, when using fsec pulses, the target is heated in regions where the ablation threshold is not exceeded. When nsec pulses are used, the sharp drop of electron density below the optical breakdown
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threshold prevents heating of the non-ablated tissue (Fig. 3). The dependence of the ablation depth on radiant exposure was found to be linear for neural tissue,104 but logarithmic for corneal tissue.89 The ablation efficiency for corneal tissue is thus highest close to the ablation threshold, and decreases for higher radiant exposures.53,103 This dependence is typical of a blow-off ablation process under conditions where the spatial energy density profile in the tissue corresponds to Beer’s law (exponential decrease of the energy density with depth). Similar conditions may arise in plasma-mediated ablation due to the shielding of deeper tissue layers by the plasma produced at the surface. The shielding-induced reduction in energy density with increasing depth is, to a certain degree, counteracted by self-focusing and self-channelling effects, leading to filament formation beyond the location of the laser focus. This phenomenon has been described theoretically first for laser beam propagation in air,107,108 but was also observed in water,70 gelatin,109 and in TEM pictures of fsec laser-irradiated cornea.29 Filament formation could possibly explain why the ablation depth for neural tissue is much larger than for corneal tissue, and exhibits a different dependence on radiant exposure. The thin filament, much smaller than the irradiated spot size, may only be able to remove mechanicallyweak material like neural tissue, but hardly influences strong materials like the corneal stroma.
Conclusions The most outstanding feature of plasma-mediated ablation is its ability to create ablation effects inside tissues that are transparent at low irradiance. On the other hand, plasma-mediated surface ablation of soft tissues enables the combination of relatively large etch depths with minimal thermal damage, and thus provides an additional unique feature. For both applications, minimization of mechanical collateral effects can be achieved by employing ultrashort laser pulses. However, if disruptive laser effects are desired, as, for example, in posterior capsulotomies, the use of nsec pulses is more appropriate. References 1. Krasnov MM: Q-switched laser iridectomy and Q-switched laser goniopuncture. Adv Ophthalmol 34:192-196, 1977 2. Aron-Rosa D, Aron J, Griesemann J, Thyzel R: Use of the neodymium:YAG laser to open the posterior capsule after lens implant surgery. J Am Interocul Implant Soc 6:352354, 1980 3. Van der Zypen E, Fankhauser F, Bebie H, Marshall J: Changes in the ultrastructure of the iris after irradiation with intense light. Adv Ophthalmol 39:59-180, 1979 4. Fankhauser F, Roussel P, Steffen J, Van der Zypen E, Chrenkova A: Clinical studies on the efficiency of high
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and transport for material processing. Appl Surf Sci 127/ 129:193-198, 1998 Feit MD, Rubenchik AM, Kim BM, Da Silva LB, Perry MD: Physical characterization of ultrashort laser pulse drilling of biological tissue. Appl Surf Sci 127/129:869874, 1998 Vogel A, Nahen K, Theisen D, Birngruber R, Thomas RJ, Rockwell BA: Influence of optical aberrations on laserinduced plasma formation in water and their consequences for intraocular photodisruption. Appl Opt 38:3636-3643, 1999 Vogel A, Busch S, Parlitz U: Shock wave emission and cavitation bubble generation by picosecond and nanosecond optical breakdown in water. J Acoust Soc Am 100:148165, 1996 Noack J, Hammer DX, Noojin GD, Rockwell BA, Vogel A: Influence of pulse duration on mechanical effects after laser-induced breakdown in water. J Appl Phys 83:74887495, 1998 Vogel A, Noack J, Nahen K, Theisen D, Busch S, Parlitz U, Hammer DX, Noojin GD, Rockwell BA, Birngruber R: Energy balance of optical breakdown in water at nanosecond to femtosecond time scales. Appl Phys B 68:271-280, 1999 Vogel A, Schweiger P, Frieser A, Asiyo M, Birngruber R: Intraocular Nd:YAG laser surgery: light-tissue interaction, damage range, and reduction of collateral effects. IEEE J Quant Electr 26:2240-2260, 1990 Juhasz T, Hu XH, Turi L, Bor Z: Dynamics of shock waves and cavitation bubbles generated by picosecond laser pulses in corneal tissue and water. Lasers Surg Med 15:91-98, 1994 Juhasz T, Kastis GA, Suarez C, Bor Z, Bron WE: Timeresolved observations of shock waves and cavitation bubbles generated by femtosecond laser pulses in corneal tissue and water. Lasers Surg Med 19:23-31, 1996 Fischer JP, Juhasz T, Bille JF: Time resolved imaging of the surface ablation of soft tissue with IR picosecond laser pulses. Appl Phys A 64:181-189, 1997 Cole RH: Underwater Explosions. Princeton, NJ: Princeton University Press 1948 Vogel A, Lauterborn W, Timm R: Optical and acoustic investigation of the dynamics of laser-produced cavitation bubbles near a solid boundary. J Fluid Mech 206:299-338, 1989 Lauterborn W, Bolle H: Experimental investigations of cavitation bubble collapse in the neighborhood of a solid boundary. J Fluid Mech 72:391-399, 1975 Brujan EA, Nahen K, Schmidt P, Vogel A: Dynamics of laser-induced cavitation bubbles near an elastic boundary. J Fluid Mech 433:251-281, 2001 Brujan EA, Nahen K, Schmidt P, Vogel A: Dynamics of laser-induced cavitation bubbles near elastic boundaries: influence of the elastic modulus. J Fluid Mech 433:283314, 2001 Vogel A, Busch S, Jungnickel K, Birngruber R: Mechanisms of intraocular photodisruption with picosecond and nanosecond laser pulses. Lasers Surg Med 15:32-43, 1994 Vogel A, Günther T, Asiyo-Vogel M, Birngruber R: Factors determining the refractive effects of intrastromal photorefractive keratectomy with the picosecond laser. J Cataract Refract Surg 23:1301-1310, 1997 Juhasz T, Kurtz R, Horvath C, Suarez C, Nordan LT, Slade SG: Femtosecond laser eye surgery: the first clinical experience. Proc SPIE 4633:1-10, 2002 Birngruber R, Hillenkamp F, Stefani FH, Gabel VP: Qswitched ruby laser damage of the rabbit eye lens. Adv Ophthalmol 34:158-163, 1977
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Mechanisms of short-pulsed plasma-mediated laser ablation and disruption 84. Stern D, Schoenlein RW, Puliafito CA, Dobi ET, Birngruber R, Fujimoto JG: Corneal ablation by nanosecond, picosecond, and femtosecond lasers at 532 and 625 nm. Arch Ophthalmol 107:587-592, 1989 85. Fabbro R, Fournier J, Ballard P, Devaux D, Virmont J: Physical study of laser-produced plasma in confined geometry. J Appl Phys 68:775-784, 1990 86. Chapyak EJ, Godwin RP: Simulations of laser thrombolysis. Proc SPIE 3590:328-335, 1999 87. Stolarski DJ, Hardman J, Bramlette CG, Noojin GD, Thomas RJ, Rockwell BA, Roach WP: Integrated light spectroscopy of laser-induced breakdown in aqueous media. SPIE Proc 2391:100-109, 1995 88. Chapyak EJ, Godwin RP, Vogel A: A comparison of numerical simulations and laboratory studies of shock waves and cavitation bubble growth produced by optical breakdown in water. Proc SPIE 2975:335-342, 1997 89. Niemz MH, Klancnik EG, Bille JF: Plasma-mediated ablation of corneal tissue at 1053 nm using a Nd:YLF oscillator/regenerative amplifier laser. Lasers Surg Med 11: 426-431, 1991 90. Boudaiffa B, Cloutier P, Hunting D, Huels MA, Sanche L: Resonant formation of DNA strand breaks by low-energy (3 to 20 eV) electrons. Science 287:1658-1660, 2000 91. Hotop H: Dynamics of low energy electron collisions with molecules and clusters. In: Christophorou LG, Olthoff JK (eds) Proceedings International Symposium on Gaseous Dielectrics IX, May 22-25, 2001, Ellicott City, MD, pp 314. New York, NY: Kluwer Academic/Plenum Press 2001 92. Nikogosyan DN, Oraevsky AA, Rupasov V: Two-photon ionization and dissociation of liquid water by powerful laser UV radiation. Chem Phys 77:131-143, 1983 93. Tirlapur UK, König K, Peuckert C, Krieg R, Halbhuber KJ: Femtosecond near-infrared laser pulses elicit generation of reactive oxygen species in mammalian cells leading to apoptosis-like death. Exp Cell Res 263:88-97, 2001 94. Timberlake GT, Gemperli AW, Larive CK, Warren KA, Mainster MA: Free radical production by Nd:YAG laser photodisruption. Ophthalmic Surg Lasers 28:582-589, 1997 95. Heisterkamp A, Ripken T, Mamom T, Drommer W, Welling H, Ertmer W, Lubatschowski H: Nonlinear side effects of fs pulses inside corneal tissue during photodisruption. Appl Phys B 74:419-425, 2002 96. Rol P, Fankhauser F, Kwasniewska S: Evaluation of contact lenses for laser therapy. Part 1. Lasers Ophthalmol 1:120, 1986
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97. Jungnickel K, Vogel A: Efficiency of bursts in intraocular Nd:YAG laser surgery. Lasers Light Ophthalmol 5:95-99, 1992 98. Zysset B, Fujimoto JG, Puliafito CA, Birngruber R, Deutsch TF: Picosecond optical breakdown: tissue effects and reduction of collateral damage. Lasers Surg Med 9:183-204, 1989 99. Lin CP, Weaver YK, Birngruber R, Fujimoto JG, Puliafito CA: Intraocular microsurgery with a picosecond Nd:YAG laser. Lasers Light Ophthalmol 15:44-53, 1994 100. Geerling G, Roider J, Schmidt-Erfurth U, Nahen K, ElHifnawi ES, Laqua H, Vogel A: Initial clinical experience with the picosecond Nd:YAG laser for intraocular therapeutic applications. Br J Ophthalmol 82:504-509, 1998 101. Niemz MH: Laser-Tissue Interactions: Fundamentals and Applications, 2nd edn. Berlin: Springer 2002 102. Neev J, Da Silva LB, Feit MD, Perry MD, Rubenchik AM, Stuart BC: Ultrashort pulse lasers for hard tissue ablation. IEEE J Sel Top Quant Electr 2:790-800, 1996 103. Kurtz RM, Elner V, Liu X, Juhasz T: Photodisruption in the cornea as a function of pulse width. J Refract Surg 13:653-658, 1997 104. Loesel FH, Fischer JP, Götz MH, Horvath C, Juhasz T, Noack F, Suhm N, Bille JF: Non-thermal ablation of neural tissue with femtosecond laser pulses. Appl Phys B 66:121-128, 1998 105. Fischer JP, Dams J, Götz MH, Kerker E, Loesel FH, Messer CJ; Niemz MH, Suhm N, Bille JF: Plasma-mediated ablation of brain tissue with picosecond laser pulses. Appl Phys B 58:493-499, 1994 106. Kim BM, Feit MD, Rubenchik AM, Joslin EJ, Eichler J, Stoller P, Da Silva LB: Effects on high repetition rate and beam size on hard tissue damage due to subpicosecond laser pulses. Appl Phys Lett 76:4001-4003, 2000 107. Braun A, Korn G, Liu X, Du D, Squier J, Mourou G: Selfchanneling of high-peak-power femtosecond laser pulses in air. Opt Lett 20:73-75, 1995 108. Couairon A, Tzortzakis S, Berge L, Franco M, Prade B, Mysyrowicz A: Infrared femtosecond light filaments in air: simulations and experiments. J Opt Soc Am B 19:11171131, 2002 109. Maatz G, Heisterkamp A, Lubatschowski H, Barcikowski S, Fallnich C, Welling H, Ertmer W: Chemical and physical side effects at application of ultrashort laser pulses for intrastromal refractive surgery. J Opt A: Pure Appl Opt 2:59-64, 2000
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The first clinical application of the laser
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The first clinical application of the laser Charles J. Koester and Charles J. Campbell Department of Ophthalmology, Columbia University, New York, NY, USA
Keywords: ruby laser, xenon arc photocoagulator, retinal blood vessels, angiomatosis retinae
Abstract In 1960, the established method of photocoagulation was the xenon arc photocoagulator. The ruby laser, invented that year by Maiman, had very different characteristics: a single short pulse, deep red wavelength, and a highly collimated beam. An experimental ruby laser photocoagulator was built, tested on rabbits, and successfully used to photocoagulate the retina of a patient in 1961. This was the first clinical application of the laser.
After the invention of the ruby laser by Maiman in 1960,1 there was considerable interest in possible medical applications. In 1961, Zaret et al. reported on experimental laser photocoagulation of rabbits.2 Investigations of the pathological effects on skin were conducted by Goldman et al., and reported in 1963.3 At that time, photocoagulation of the retina using a xenon arc source was an established procedure, following the pioneering work of Meyer-Schwickerath.4 Treatment required a 1000-W xenon arc lamp, and an optical delivery system that was impressive in its design, size, and articulation. The patient was typically seated in a recliner chair, and the eyepiece of the delivery system was brought to the patient. Exposure times were approximately one second, and were frequently a source of discomfort. The ruby laser also required a large power supply plus a powerful flash lamp. The laser output was about 1 msec in duration, so that the laser light itself could not be used for aiming. Furthermore, the energy output from the first ruby lasers was only about one quarter of the energy that was required by the xenon arc lamp to produce a photocoagulation on the human retina. However, the short duration of the laser output more than compensated for the low energy level, as was soon demonstrated in experiments on rabbits.
Aiming the laser beam required a different approach from that used in the xenon arc instrument, since the ruby laser emitted a single pulse, followed by a significant pause for recharging the bank of capacitors. Aiming the beam into the eye was accomplished by using a unique property of the laser: the light output emerges from a laser rod in a direction precisely perpendicular to the mirror end surface of the rod. Aiming was achieved by collimating white light from a point source so that the beam was perpendicular to the end surface of the laser. A portion of the collimated white light was directed into the eye using the mirror of an indirect ophthalmoscope. The laser output followed the same path through the pupil and to the retina. Experiments on rabbits were carried out at the research laboratory of American Optical Co. by Charles J. Campbell, MD. They served to define the range of power output that would produce appropriate photocoagulations. When the instrument and the aiming system were ready for clinical application, the instrument, with its power supply and large bank of capacitors, was transported to the Harkness Eye Institute at Columbia University. Dr. Campbell treated the first patient on November 22nd, 1961. Other participants in the project were Charles J. Koester, Elias Snitzer, Stephen M. MacNeille, Vonda Curtis, and M. Catherine Rittler. The New York Times and Wall Street Journal reported the event during the following weeks, and the first presentation was at a meeting of the Optical Society of America in March 1962.5 At the same meeting, Zaret et al. reported ruby laser photocoagulation experiments on rabbits.6 The first human subject had a diagnosis of advanced angiomatosis retinae that had destroyed virtually all vision in the affected eye. Xenon arc photoco-
Address for correspondence: Charles J. Koester, PhD, 60 Kent Road, Glen Rock, NJ 07452, USA. e-mail:
[email protected] Lasers in Ophthalmology – Basic, Diagnostic and Surgical Aspects, pp. 115–117 edited by F. Fankhauser and S. Kwasniewska © 2003 Kugler Publications, The Hague, The Netherlands
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Fig. 1. The patient’s fundus photograph taken immediately after the laser exposures. Two laser photocoagulations can be seen on or near blood vessels just above the center of the photograph.
agulation had been proposed to destroy the abnormal vessels, but because of the patient’s youth, that procedure would have required general anesthesia. The ruby laser held the promise of a more readily tolerated therapeutic approach, even by a young patient. Therefore, ruby laser photocoagulation was explored for the first human patient, even though the wavelength of the ruby laser (0.694 µm) was not optimum for absorption by the blood and blood vessels. Figure 1 is the fundus photograph taken immediately after the laser exposure. An area of the retina was selected, away from the pathology, in order to determine the power required for retinal coagulations. Normal vessels were selected because they posed a lesser risk of fragility. The areas of pathology were then subjected to photocoagulation intensities, based on the calibration exposures. Figure 2 is the same area as Figure 1, photographed six days after exposure, showing scars and pigmentation in the treated areas, with excellent localized coagulation. However, the blood vessels appear to be intact. Even though the treatment could be performed comfortably, it was concluded that the single pulse ruby laser exposures were not effective in destroying blood vessels. These and other early results were reported at the American Academy of Ophthalmology and Otolaryngology in 19637 and in the Archives of Ophthalmology in 1964.8 A water-cooled, rapidly pulsed (60 Hz) ruby laser photocoagulator was later developed by American Optical Co. Campbell and Rittler reported that the appearance of retinal lesions was virtually identical to lesions produced with an argon laser.9 An indirect ophthalmoscope was used for viewing the fundus and aiming the beam. Absorption of the laser light by blood or blood vessels was limited. Photophobia from the ruby laser light was not marked.
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Fig. 2. The same area as in Figure 1, photographed six days after laser exposure.
Ross et al. developed a ruby laser pulsed at 80 Hz.10 This laser, effectively a continuous source, was combined with an adaptation of the Goldmann threemirror contact lens and provided a precise aiming system for photocoagulation.11 Thus, as the characteristics of the ruby laser were developed and improved for the purpose of photocoagulation, the aiming systems were adapted to provide optimum viewing of the fundus and aiming of the laser beam.
Conclusions The first laser photocoagulator demonstrated the feasibility of the laser as a light source for clinical photocoagulation. The ruby photocoagulator also identified the limitations of a very short pulse, deep red wavelength light source. These limitations have since been removed by improved systems that employ ruby and other laser sources. References 1. Maiman TH: Stimulated optical radiation in ruby. Nature 187:493-494, 1960 2. Zaret MM, Breinin GM, Schmidt H, Ripps H, Siegel IM: Ocular lesions produced by an optical maser (Laser). Science 134:1525-1526, 1961 3. Goldman L, Blaney DJ, Kindel DJ, Richfield D, Frank EK: Pathology of the effect of the laser beam on the skin. Nature 197:912-914, 1963 4. Meyer-Schwickerath G: Light coagulation. St Louis, MO: CV Mosby 1960 5. Koester CJ, Snitzer E, Campbell CJ, Rittler MC: Experimental laser retina coagulator. J Opt Soc Am 52:607, 1962 6. Zaret MM, Ripps H, Siegel IM, Breinin GM: Biomedical experimentation with optical masers. J Opt Soc Am 52:607, 1962 7. Campbell CJ, Rittler MC, Koester CJ: The optical maser as a retinal coagulator: an evaluation. Trans Am Acad Ophthalmol Otol 67:58-67, 1963
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The first clinical application of the laser 8. Noyori KS, Campbell CJ, Rittler MC, Koester CJ: The characteristics of experimental laser coagulations of the retina. Arch Ophthalmol 72:254-263, 1964 9. Campbell CJ, Rittler MC: Effects of lasers on the eye. New York Academy of Science: Laser Conference, February 1970, p 36. 1970
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117 10. Röss D et al: Pumping new life into ruby lasers. Electronics 39:115-118, 1966 11. Fankhauser F, Lotmar W: Photocoagulation though the Goldmann contact glass. Ophthalmology 77:320-330, 1967
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Selective retinal pigment epithelium laser treatment Theoretical and clinical aspects
Johann Roider1, Ralf Brinkmann2 and Reginald Birngruber2 1 Department of Ophthalmology, University of Kiel, Kiel, Germany; 2Medical Laser Centre, Lübeck, Germany
Keywords: pigment epithelium, selective absorption, diabetic retinopathy
Introduction
Healing response after conventional laser photocoagulation
Retinal photocoagulation has been performed for more than 30 years and its value in various macular diseases has been established. Typically, the power of the laser is adjusted to produce white or gray retinal lesion depending on the depth of the coagulation. Heat conduction from the irradiated retinal pigment epithelium (RPE) into the retina may lead to irreversible thermal denaturation of the outer and inner segments.1-3 However, the benefit of retinal laser treatment has traditionally been attributed to the destruction of retinal tissue. So, the threshold energy required to effect most treatments remains unknown. Heat conduction from the RPE into the retina in a typical laser lesion leads to irreversible thermal denaturation of the outer and inner segments.1-3 The exact biological mechanism of how such damage produces a therapeutic effect is poorly understood. However, for a variety of retinal diseases, which are thought to be associated with a degradation of the RPE, it might be sufficient to selectively damage the malfunctioning RPE, but to spare the overlying photoreceptors in order to avoid scotoma, which is especially useful within the macula. If the damaged RPE is rejuvenated in the healing process, due to migration and proliferation of the adjoining RPE, minimal, destructive, selective RPE treatment might be optimal. Various aspects of this new selective RPE treatment will be discussed in the following paragraphs and will be evaluated with respect to conventional laser photocoagulation.
The most common explanation for the beneficial effect of photocoagulation in diabetic retinopathy is the destruction of oxygen consuming photoreceptors.4 Another theory suggests that the beneficial effect results from the restoration of a new RPE barrier and the subsequent production of a variety of growth factors.5-8 While the exact interaction mechanisms of the different growth factors is not well understood, the vascular endothelium growth factor (VEGF) seems to play a major role in regulation of the neovascularization associated with ischemic retina.9,10 In the treatment of diabetic macular edema, the beneficial effect is thought to be mediated by restoration of a new RPE barrier.11 A similar effect can be postulated in the treatment of drusen. Drusen are located within Bruch’s membrane or beneath the RPE, and often disappear after photocoagulation of the surrounding tissues. The value of prophylactic treatment of drusen is now being studied by several investigators.12-14 Based on the fact that drusen are located deep in the retina, there is no clear rationale routinely to include the neural retina in photocoagulation. For instance, in central serous retinopathy (CSR), the rationale of therapy is the photocoagulation and subsequent formation of a new RPE barrier, replacing the old diseased RPE cells. Destruction of the photoreceptors would only appear to be an unwanted side-effect of such a disease, and the same may be true in the treatment of macular edema. Following laser photocoagulation, the targeted tissue undergoes a healing process which is inde-
Address for correspondence: Johann Roider, MD, Klinik für Ophthalmologie, Universitätsklinikum Schleswig-Holstein, Campus Kiel, Hegewischstr. 2, 24105 Kiel. e-mail:
[email protected] Lasers in Ophthalmology – Basic, Diagnostic and Surgical Aspects, pp. 119–129 edited by F. Fankhauser and S. Kwasniewska © 2003 Kugler Publications, The Hague, The Netherlands
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pendent of the laser used. Typically, the tissue in the laser target area will be replaced by proliferating glial tissue originating from the surrounding retina and choroid. In addition, RPE cells contribute a significant part to this healing process,15 and the outer blood retinal barrier is reformed after about seven days.16 In animal experiments, it has been shown that the RPE may respond in several ways after injury. Adjacent RPE cells may spread out and the defect will be filled by hypertrophy of the neighboring RPE cells.17,18 Another mechanism is cell division of the RPE cells. This has been shown after photocoagulation in rabbits,19 in monkeys after retinal detachment,19 and in rabbits after surgically-induced RPE defects.20 Glaser et al.21 have shown that RPE cells produce inhibitors for neovascularization, suggesting that these cells may play a role in regulation of the neovascularization process. In addition, Boulton et al.22 found a significant change in growth factors in the vitreous after panretinal photocoagulation. Yoshimura et al.8 showed that photocoagulated RPE cells produce inhibitors of endothelial cells. The molecular and immunological characteristics of these inhibitors correlate with TGF-ß2. Regenerating RPE cells are known to produce more TGF-ß2 compared to normal RPE cells.23 Based on the experimental findings elucidated above, photocoagulation and stimulation of the RPE layer alone might theoretically be enough to mediate a biological response in the treatment of retinal diseases, without destroying the photoreceptors. That is, rather than targeting the sensory retina, we should consider the very important role of the RPE cells. It may be that these, rather than the sensory retina, are of primary importance when producing an effect from photocoaguation. Temperature and histology after laser photocoagulation When laser light strikes the retina, a high percentage of the energy is absorbed in the RPE layer, and this percentage is wavelength-dependent. With green light, about 50% of the laser light is absorbed by the RPE.24 The typical exposure time of conventional argon laser treatment is 100 msec or more. While the laser energy is being absorbed, it is converted to heat which diffuses out of the absorbing RPE layer in three dimensions; i.e., heat is conducted posteriorly into the choroid and anteriorly in the direction of the neural retina. Heat is also transferred horizontally, along the RPE cells and retinal layers. Heat diffuses out of an absorbing structure at a speed of roughly 1 µsec per µm (1 µs = 10-6 s). Hence, a laser exposure of 100 ms results in considerable heat conduction. The spatial and temporal temperature distribution can be calculated by mathematical models, and can be verified experimentally.25,26 Figure 1 shows the calculated temperature profile inside the RPE and inside the neural retina. It is obvious that
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Temperature [°C]
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Retina
Fig. 1. Computed spatial temperature profile in the RPE and retina after exposure to a 100 ms argon (514 nm) laser pulse with various exposure times (50 and 100 ms, 1 s) and various threshold powers (spot size 110 µm). There is only a small temperature gradient between the RPE and the neural retina.
only a small temperature difference exists between the RPE and neural retina after a 100-ms exposure. This difference is about 18% from the RPE to 5 µm into the retina. Therefore, lowering the laser power, while keeping the exposure time constant, will not protect the retina from thermal damage. Instead, the exposure duration must be reduced. Figure 2 shows the histology of the retina of an argon laser exposure with 20 mW and 100 ms on a 110-µm spot diameter. The power was chosen so that the lesion was not clinically visible, but it was detectable by fluorescein angiography. Despite the subtle nature of the lesion, the photoreceptors have been irreversibly damaged. Over time, the damaged photoreceptors will be replaced by scar tissue. This drop out of photoreceptors, after even mild photocoagulation very close to the damage threshold, is obvious. Concept of selective retinal pigment epithelium treatment Sparing of adjacent structures is only possible if the laser pulse duration is matched with the target’s physical characteristics. The pulse duration needed to spare the neural retina can be estimated by the thermal relaxation time or the time interval required for the heat to diffuse out of an absorbing tissue.25,27,28 If we consider that the size of an RPE cell is about 10 µm, high temperatures can be confined to the RPE cell itself, if the exposure time is of the order of a few microseconds rather than the customary millisecond settings. Figure 3 shows the spatial temperature profile as calculated after a argon laser exposure of 1 µsec. The difference in the temperature profile compared
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µm Fig. 2. Light micrograph of an ophthalmoscopically non-visible laser lesion obtained 14 days after exposure to a 100 ms argon laser pulse (power: 20 mW, spot diameter: 110 µm). A significant dropout and damage of photoreceptors is seen. Magnification 500 ×.
to Figure 1 is obvious. It is conceivable that most of the heat is concentrated within the RPE at the end of the laser pulse, and the temperature elevation inside the retina can be minimized. Hot spots do occur, but these are situated around single melanin granules. Since no more laser energy is delivered at the end of the laser pulse, the temperature profile quickly smoothes out and leads to only a very low temperature increase at the retina. This effect is especially pronounced if repetitive laser pulses are applied in such a manner that the following laser pulse is applied only after the retinal tissue has had sufficient time to completely cool down to baseline. Thus, the pulse repetition rate has to be matched to the irradiation in order to achieve high temperatures inside the RPE and low temperatures at the adjoining photoreceptors. If the melanosome temperatures become high enough, thermal denaturation of cell proteins close to the granules, as well as coagulation of the whole cell, can take place. However, the temperatures required for denaturation in the microsecond time domain are still unknown. If the vaporization temperature around the melanin granules is exceeded, i.e., about 140°C, microbubble formation occurs.29 The expansion and collapse of the these microbubbles can cause thermomechnical disruption in the RPE cells.25 Figure 4 shows the histological effect after exposure to a chain of 500 repetitive 5-µs laser exposures. It is obvious that the RPE is heavily damaged, but most of the photoreceptors have been spared. Similar histological effects have been shown after repetitive 200-nsec laser pulses with a Nd: YAG (532 nm) laser.30 RPE defects after selective RPE treatment18,31 may be covered by a new population of RPE cells. Cells close to and far away from the treatment site react, and a new RPE barrier is quickly restored. This effect on the RPE is not different from the reaction of the RPE after conventional laser photocoagulation.2,15 Since the RPE cells do not divide during their lifetime, apart from special circumstances, therapeutical laser effects could be explained by direct or indirect stimuli originating from RPE alterations.
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Fig. 3. Computed spatial temperature profile in the RPE and retina after exposure to a 1 µs argon (514 nm) laser pulse (2 µJ, 110 µm). A significant temperature gradient between the RPE and retina is conceivable. The hottest spots are conceivable around single melanin granules, which are absorbers of the laser light inside the RPE.
Fig. 4. Transmission electron micrograph of an ophthalmoscopically non-visible laser lesion obtained two hours days after exposure to 500 repetitive 5-µs argon laser pulses (3 µJ, 110 µm, 500 Hz). Most of the photoreceptors have been spared and the outer segments look normal. Magnification 3000 ×.
Therapeutical effects after selective retinal pigment epithelium treatment A salient question is whether, and in which diseases, selective RPE treatment leads to a positive therapeutic effect. In a first clinical pilot study, we focused on three pathological conditions: diabetic macular edema, central serous retinopathy, and drusen in age related macular degeneration (AMD). Treatment was performed either using a chain of repetitive laser pulses with a frequency-doubled Nd:YLF laser (wavelength 527 nm, pulse duration 1.7 µs, 100 and 500 Hz repetition rate, 30 and 100 pulses and a retinal spot diameter of 160 µm) or a pulsed Nd:YAG laser (wavelength 532 nm, pulse duration 800 ns, 500 Hz repetition rate, 100 pulses and a retinal spot diameter of 200 µm). Figure 5 shows the appearance of the fundus of a patient who had been treated for drusen with a
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Table 1. Number of relative defects detected by microperimetry after the application of test exposures (160 µm, 527 nm) in the lower macula during the follow-up period. RPE defects were obtained by a chain of short 1.7-µsec laser pulses at a repetition rate of 500 Hz Number of defects 500 pulses Energy per single pulse
70 µJ n
1 1 1 3 6 1
day week month months months year
28 25 25 21 18 2
100 µJ n
18 7 9 1 2 0
11 10 10 8 6 3
70 µJ n
8 8 5 0 2 0
9 6 6
0 0 0
6
0
100 pulses 100 µJ n 17 11 4 7
0 0 1 0
130 µJ n 8 1 2
3 0 0
n: number of laser lesions tested
3) were treated and followed up for one year.32 Treatment was performed using an Nd:YLF laser. Laser energy was based on the visibility of test lesions on fluorescein angiography (50-130 µJ). Patients were examined at various times by ophthalmoscopy, fluorescein- and ICG-angiography, as well as by infrared imaging. After six months, hard exudates disappeared in six of nine patients in group 1, and leakage had disappeared in six of 12 diabetic patients. In group 2, there were less drusen in seven of ten patients. In group 3, serous detachment had disappeared in three of four cases. Visual acuity was stable in all cases. Selectivity of retinal pigment epithelium treatment (sparing the retina)
Fig. 5. Fundus photograph in a patient: (A) before, and (B) 12 months after laser treatment for drusen. A significant reduction in the number of drusen is visible.
Nd:YAG laser. After test exposures, selective RPE spots were placed in a horseshoe pattern around the fovea. Drusen slowly began to disappear after three months.32 In another study, 12 patients with diabetic maculopathy (group 1), ten with soft drusen (group 2), and four with central serous retinopathy (CSR) (group
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In a first pilot study, test laser lesions were used to investigate whether sparing of the retina was possible in the human retina.33 In order to investigate whether selective RPE effects were really selective from a microperimetry point of view, microperimetry was performed directly on top of laser lesions during a follow-up period of up to one year. A repetitive pulsed Nd:YLF laser was applied in 17 patients, using pulse energies of from 20-130 µJ. To establish the necessary energy, test exposures were performed in the inferior macular region (see, for example, Fig. 6A). Of 179 test lesions, 73 were followed at various times by performing microperimetry directly on top of the laser lesions. For testing, laser lesion threshold stimuli were determined before laser exposure. The threshold sensitivity values were defined as the minimal contrast at which a response was obtained. This threshold value was used to evaluate the test lesions during the follow-up period. Probit analysis showed that all the test lesions were at the threshold of RPE disruption and that none of the laser effects were visible by means of ophthalmoscopy during photocoagulation, and were only detectable by fluorescein angiography. After exposure with 500 pulses, retinal defects could be detected in up to 73% of patients (100 µJ) after the first day. Most of these defects were no longer detectable after three months. After exposure with 100
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* V
r
Fig. 6. (A) Fluorescein angiogram of a test area one day after application of several test exposures with different laser parameters (arrowhead: 100 pulses, 100 µJ, asterisk: 500 pulses, 70 µJ, straight arrow: 500 pulses, 100 µJ). Microperimetry image (B) (Rodenstock scanning laser ophthalmoscope) performed over the test lesions with a threshold stimulus (Goldmann II) performed over areas which had been treated with various energies and laser pulses. The threshold sensitivity value of the stimulus was evaluated before photocoagulation and was defined as the minimal contrast at which a response was obtained. No retinal defect due to laser photocoagulation could be detected over most laser effects. All stimuli were recognized. Only small relative defects (black bars) could be found over lesions with higher energies (500 pulses and 100 µJ). The other letters correspond to stimuli, which had been recognized by the patient.
Fig. 7. A: Fundus photograph of a diabetic patient treated both for diabetic macular edema and proliferative diabetic retinopathy one year earlier. B: Microperimetry of laser scars achieved after continuous wave irradiation. Absolute defects (black bars) can be detected over each laser scar visible on ophthalmoscopy.
pulses, no defects could be detected with 70 and 100 mJ after one day, and the neural retina remained undamaged during the follow-up period. Table 1 summarizes the results of relative defects after selective RPE treatment.33 Figure 6A shows the fluorescein angiogram of a patient in whom laser effects have been applied to the part below the macula. This patient was scheduled for treatment for soft drusen. Figure 6B shows the corresponding infrared image with the threshold stimuli recognized by the patient superimposed. No retinal defects could be detected after selective RPE treatment with pulse energies below 100 µJ and 100 or 500 pulses. Our findings are in contrast to microperimetry findings after conventional laser photocoagulation, which is associated with thermal damage to the outer and inner nuclear layer and replacement by scar
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tissue.2,3,15,31,34 It is not surprising that such lesions lead to absolute scotomas on microperimetry. Figure 7 shows an example of such a patient treated with conventional laser techniques for diabetic macular edema and proliferative diabetic retinopathy. Absolute scotomas can be seen over each argon laser exposure. If such lesions are located temporal to the fovea, the patient may have problems with reading, despite good visual acuity. Autofluorescent behavior after conventional and selective laser treatment (imaging the retinal pigment epithelium) Photoreceptors degenerate as soon as they are devoid of RPE cells for an extended period. The question of whether RPE cells regenerate is crucial for
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retinal sparing laser treatment. The RPE contains high amounts of age-accumulated lipofuscein, which is also a strong fluorophore. Lipofuscin can be divided into ten subcomponents and is considered to be a storage granule within the pigment epithelium.35 The excitation spectrum of fundus autofluorescence has a broad band between 450 and 540 nm, with a maximum around 500 nm. The emission spectrum of lipofuscin ranges around 500780 nm, with a maximum at 620 nm.35 In selective RPE treatment, most of the energy is absorbed in the RPE, and thus a change of its autofluorescence due to the thermal effects of lipofuscein can be expected. Since changes in the RPE occur both during and after laser treatment in the healing phase, monitoring its autofluorescence might be a helpful tool both for verifying as well as characterizing the laser effects. After mild, barely visible continuous wave photocoagulation (100-200 ms, 100 µm, power equivalent to a barely visible laser burn), there is a change in autofluorescent behavior over a period of one year, which is much longer than we would have expected from animal experiments. Autofluorescence has been observed by retina angiography (HRA) in the fluorescence mode, using excitation and barrier filters, as used during fluorescein angiography, excitation wavelength 488 nm, without using fluorescein. This was examined in 13 patients, in whom focal diabetic macular edema had been treated by conventional laser parameters, as described above. In 12 of the 13 patients treated by conventional focal laser photocoagulation, the laser spots appeared as hypofluorescent areas one hour after treatment. The autofluorescent behavior of different laser exposures did not vary between patients, but did change dramatically with time. Figure 8A shows fluorescent behavior one hour after focal laser treatment. One month after focal laser treatment, all laser spots examined appeared as hyperfluorescent areas (Fig. 8B). This situation remained unchanged for three months. After six months, most of the spots (94%) were highly hyperfluorescent. Some had a mixed form with a central white hyperfluorescent island
surrounded by a dark hypofluorescent area. After a period of more than 12 months, most of the laser spots once again appeared hypofluorescent (85%) (see Fig. 8C), while 15% showed a mixed appearance. The changes in autofluorescent behavior were significant between one hour and one month after treatment (P = 0.0001), and also between six months and one year (P = 0.001). These findings show that changes within the RPE cells occur after laser treatment over a long period of time. These autofluorescence findings can partially be explained by the histological results of animal experiments. After one hour, autofluorescence photographs show that all cells are damaged and that the irradiated area appears to be hypofluorescent. These findings show that all fluorophores are completely destroyed. After laser treatment, the cell debris from damaged RPE cells and photoreceptors is phagocitized by RPE cells sliding in from the surrounding area or by macrophages originating from the choriocapillaris.18 Abundant phagosomes can be found within these cells during this highly active period. These storage granules accumulate within the cells and could be responsible for the hyperfluorescent signal. In animal experiments, the damaged irradiated area is repopulated by RPE cells or RPEderived cells. After conventional laser photocoagulation, all photoreceptors are irreversibly destroyed and scarring can be seen. Lipofuscein is a storage granule. Several observations have confirmed that most of the material accumulating in RPE lipofuscin granules in AMD originates from phagocytosis of photoreceptor material.36,37 When the RPE is destroyed, the newly formed RPE does not accumulate the same amount of lipofuscin as seen in animals reared in dim light.38 This would suggest that no storage granules develop. This may explain why the irradiated area appears to be hypofluorescent after one year, when scarring has occurred. In the meantime, transition from a highly active period to a quiet period occurs. One year after conventional laser treatment, the RPE has obviously obtained a different status of viability compared to the initial one.
Fig. 8. Autofluorescent images as obtained with the Heidelberg angiograph: A: one hour; B: three months; and C: one year after treatment, after conventional mild laser treatment. After one hour, the laser spots appear hypofluorescent; after three months, highly hyperfluorescent; and after one year, they appear hypofluorescent again.
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Fig. 9. A: Fluorescein angiogram and corresponding autofluorescence image, B: one hour and C: one year after selective RPE treatment. One year later, autofluorescence of the laser spot between the two vessels is similar to the surrounding area, suggesting complete recovery of the RPE.
Future studies will show whether the RPE is capable of completely regenerating after selective treatment. Figure 9 shows the autofluorescent image one hour (Fig. 9B) and one year (Fig. 9C) after selective RPE treatment in a 36-year-old patient. Images of the laser spots between the two large vessels one year after selective RPE treatment are very similar to the autofluorescence image of the surrounding area, suggesting the complete recovery from damage after selective RPE treatment. Monitoring retinal pigment epithelium damage for dosimetry by means of autofluorescence Selective RPE treatment requires a refinement of laser protocol, since it is a problem that RPE laser lesions are not clinically visible after placement. This is in contrast to conventional retinal photocoagulation where the laser spots can be seen as a whitening, since the coagulated retina scatters the illumination light of the slit lamp in all directions. Fluorescein angiography is even more sen-sitive in detecting coagulated areas, compared to conventional observation or photography. In fact, after conventional cw laser exposure, fluorescein angiography is twice as sensitive as ordinary black and white or color fundus photographs.36 In selective RPE treatment, the heat is confined to the RPE, which shows that, on its own, RPE damage cannot be observed by white light illumination. However, in animal experiments, using repetitive µs or 200 nsec laser exposures,18,30 fluorescein angiography shows RPE damage quite well, by indirectly demarking a tight junction defect in the treated RPE areas. However, noninvasive dosimetry control is needed. Preferably, this should serve directly during treatment as an on-line method of control. In order to monitor the acute changes in autofluorescence during selective RPE treatment, the laser beam itself was simultaneously used for fluorescence excitation, since its wavelength of 527 nm is well within the excitation spectrum of lipofuscein. The
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fluorescent light originating from the fundus was detected by a photodiode coupled to the slit lamp. A dichroic beam splitter was mounted in the pathway of the laser beam within the slit lamp, as shown in Figure 10. The excitation spectrum of the fundus autofluorescence and the wavelength of the laser is described on the previous page first paragraph and the laser wavelength a few lines ahead. Figure 11 shows its spectral transmission. The photodiode current was recorded by a PC and was promptly integrated over each laser pulse applied. By means of this set-up, the fluorescence signal is detected during selective RPE treatment. The fluorescence intensity was strong enough for a photodiode instead of a more sensitive photomultiplier to be used. In the four patients treated for diabetic macular edema, decay of autofluorescence over the laser pulses applied was detected at each laser spot. Figure 12A shows the typical time course of the signal for different pulse energies of single spots in one patient. Time and spectral integrated fluorescence was normalized to the fluorescence of the first pulse. For all pulse energies, fluorescent intensity decreases with an increasing number of pulses. In order to quantify fluorescent decay, a monoexponential decay according to I(n) ~ e-n/τ, was fitted to the data; I(n) represents the fluorescent intensity after n pulses within the chain, τ is the decay constant. Figure 12B shows the mean decay constants and their statistical variations at various pulse energies in four different patients. For 50 and 70 µJ, two spots per patient were applied (test exposures), while many more spots were applied and averaged in Figure 12B at 100 and 130 µJ (treatment pulses). In two patients, the decay constant slightly decreased with pulse energy, however, this was not statistically relevant. In the other two patients, strong variations with pulse energy were found. In order to use the decay of the autofluorescence as an on-line dosimetry criterium, a threshold value has to be defined. Thus, the decay constant has to be significantly different with radiant exposures below and above RPE damage. This could be proved
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treatment laser Nd:YLF* retinal fluorescence light photodiode
laser block filter dichroic beamspliter
[τ]
Fig. 10. Schematic drawing of the modified slit lamp. Using a beamsplitter and a photodiode as shown in Figure 11, it is possible to collect the fluorescence light during selective RPE treatment. *Neodym:Yttrium-lithium floride.
Fig. 11. A: Transmission spectrum of the beam splitter, as used with the slit lamp to separate the fluorescent light (B) from the laser beam. (The excitation and emission spectra is taken from Delori et al.35)
in vitro using porcine RPE samples, in which a high correlation between autofluorescence and cell damage, demarked with a viability stain, was found. Clinically, correspondence has to be found with respect to fluorescein leakage in angiography. In the four patients, threshold energy for RPE disruption was between 70 and 130 µJ. However, the decay constants from 50-130 µJ do not differ significantly,
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Fig. 12. A: Autofluorescent decay during laser irradiation with a chain of 93 repetitive laser pulses at a rate of 500 Hz, obtained from patient A. The fluorescent intensity decreases on average for all pulses, however strong variations between pulses were detected. B: Autofluorescent decay constants (see text) of four patients treated for diabetic macular edema with various pulse energies. Error bars indicate 1 SD.
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Fig. 13. B: Two-dimensional autofluorescent image obtained with the Heidelberg device one hour after treatment with selective laser pulses. A: The corresponding fluorescein angiogram shows the laser spots as hyperfluorescent areas.
the first clinical results presented here show high fluctuations of the decay constant, and so far no clearcut threshold. We are reporting the first results. It should be taken into account that micromovement of the fundus during irradiation, as well as the nonuniform lipofuscin distribution across the RPE, influence autoflourescence. Furthermore, fluorescein angiography and a change in autofluorescence signify very different events, which are not necessarily correlated. With fluorescein angiography, a break-up of tight junctions between the RPE cells can be detected. After laser photocoagulation, this happens if the cell is totally destroyed as shown by histology after laser treatment.18,31 The decay of RPE autofluorescence may reflect the thermal alteration of lipofuscin, or at least, the temperature dependence of its fluorescence. However, up to now, it is unclear whether the on-line detection of autofluorescence can be used for dosimetry in order to monitor RPE damage during laser treatment. Imaging of RPE autofluorescence after treatment by retinal angiography in the autofluorescent mode (Heidelberg Engineering, Heidelberg, HRA) typically showed a faint hypofluorescent spot at the site of exposure ten minutes after treatment. This hypofluorescent spot became more pronounced one hour after treatment (Fig. 13B), as discussed above. In conclusion, two-dimensional imaging of autofluorescence one hour after laser exposure has been proved to be reliable in detecting RPE damage, as long as the signal is not weakened by fluid originating from retinal edema.
tion. From a theoretical point of view, this technique may require different strategies in order to achieve efficacy; i.e., more RPE changes may be required compared to conventional laser photocoagulation. Macrophages, which are one of the cells that also react and migrate after continuous wave photocoagulation, are less likely to be as stimulated after selective RPE treatment compared to customary treatment.18 It may be that macrophages play an important role in the resolution of certain retinal diseases. These pilot studies on selective RPE treatment, as well as studies on subthreshold diode laser burns,39,40 suggest that actual blanching of the retina, as for example that recommended by early treatment diabetic retinopathy (ETDR) studies, is not necessary. Despite the undebatable advantages of laser photocoagulation with current laser techniques, local defects cannot be avoided with these techniques. As shown above, the use of selective RPE treatment can avoid the formation of local scotomas. Another advantage of solely RPE-related laser effects is that retreatment may be easier to implement while still achieving a therapeutic effect. Acknowledgements The authors wish to thank Georg Schüle, Carsten Framme, Hann Elsner and Norman Michaud for their collaboration and contribution to the joint project on selective RPE treatment.
References Conclusions Both experimental and initial clinical results are encouraging and support the concept of selective RPE treatment, which seems to be effective for certain diseases, while being less harmful to visual func-
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1. Birngruber R, Gabel VP, Hillenkamp F: Experimental studies of laser thermal retinal injury. Hlth Phys 44:519-531, 1983 2. Wallow IH, Tso MOM, Fine BSF: Retinal repair after experimental xenon arc photocoagulation. 1. A comparison between rhesus monkey and rabbit. Am J Ophthalmol 75:3252, 1973
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3. Wallow IH, Birngruber R, Gabel VP, Hillenkamp F, Lund OE: Netzhautreaktion nach Intensivlichtbestrahlung. Adv Ophthalmol 31:159-232, 1975 4. Wolbarsht ML, Landers MB III: The rationale of photocoagulation therapy for proliferative diabetic retinopathy: a review and a model. Ophthalmic Surg 11:235-245, 1980 5. Marshall J, Clover G, Rothery S: Some new findings of retinal irradiation by krypton and argon lasers. In: Birngruber R, Gabel VP (eds) Laser Treatment and Photocoagulation of the Eye. Doc Ophthalmol Proc Series 36, pp 21-37. The Hague: Junk Publ 1984 6. Marshall J: Cell biology and mechanisms in panretinal photocoagulation. Laser in der Ophthalmologie, 14th laser seminar, Medical Laser Center Lübeck, November 29, 1997, Medizinische Laserzentrum Lübeck, D 23562 Lübeck 1997 7. Yamamoto C, Ogata N, Yi X, Takahashi K, Miyashiro M, Yamada H, Uyama M, Matsuzaki K: Immunolocalization of transforming growth factor ß during wound repair after laser photocoagulation. Graefe’s Arch Clin Exp Ophthalmol 236:41-46, 1998 8. Yoshimura N, Matsumoto M, Shimizu H, Mandai M, Hata Y, Ishibashi T: Photocoagulated human retinal pigment epithelial cells produce an inhibitor of vascular endothelial cell proliferation. Invest Ophthalmol Vis Sci 36:16861691, 1995 9. Aiello LP, Pierce EA, Foley ED, Sulliva R, Chen H, Ferrara N, King GL, Smith LEH: Inhibition of vascular endothelial growth factor (VEGF) reduces retinal neovascularization in the mouse. Invest Ophthalmol Vis Sci (Suppl) 36:401, 1995 10. Miller JW, Shima DT, Tolentino M, Gragoudas ES, Ferrara N, Connolly EJ, Folkman J, D’Amore PA, Adamis AP: Inhibition of VEGF prevents ocular neovascularization in a monkey model. Invest Ophthalmol Vis Sci (Suppl) 36:401, 1995 11. Bresnick GH: Diabetic maculopathy: a critical review highlighting diffuse macular edema. Ophthalmology 90:13011317, 1983 12. Figueroa MS, Regueras A, Bertrand J, Aparicio MJ, Manrique MG: Laser photocoagulation for macular soft drusen. Retina 17(5):378-384, 1997 13. Fine SL, Maguire MG, Ho Allen, Javornik NB, CNVPT Research Group: Short-term effects of light laser treatment for eyes at high risk of choroidal neovascularisation from AMD. American Academy of Ophthalmology, Annual Meeting, San Francisco, CA, October 26-29, 1997, final program, 142 (abstract) 1997 14. Frennesson C, Nilsson: Significant decrease in exudative complications after prophylactic laser treatment of soft drusen maculopathy in a randomized study. Invest Ophthalmol Vis Sci (Suppl) 38(4):18, 1997 15. Wallow IH: Repair of the pigment epithelial barrier following photocoagulation. Arch Ophthalmol 102:126-135, 1984 16. Johnson RN, McNaught EI, Foulds WS: Effect of photocoagulation on the barrier function of the pigment epithelium. II. A study by electron microscopy. Trans Ophthalmol Soc UK 97:640-651, 1977 17. Bülow N: The process of wound healing of the avascular outer layers of the retina: light- and electron microscopic studies on laser lesions of monkey eyes. Acta Ophthalmol (Kbh) (Suppl) 139:7-60, 1978 18. Roider J, Michaud N, Flotte T, Birngruber R: Response of the RPE to selective photocoagulation of the RPE by repetitive short laser pulses. Arch Ophthalmol 110:1786-1792, 1992 19. Inomata H: Wound healing after xenon arc photocoagulation in the rabbit retina. Ophthalmologica 170:462-474, 1975
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20. Heriot WJ, Machemar R: Pigment epithelial repair. Graefe’s Arch Clin Exp Ophthalmol 230:91-100, 1992 21. Glaser BM, Campochiaro PA, Davis JL, Jerdan JA: Retinal pigment epithelial cells release inhibitors of neovascularization. Ophthalmology 94:780-784, 1987 22. Boulton ME, Xiao M, Khaliq A, Moriarty P, Cranley J, McLeod D: Changes in growth factor expression in pig eyes following scatter laser photocoagulation. Invest Ophthalmol Vis Sci (Suppl) 36:95, 1995 23. Matsumoto M, Yoshimura N, Honda Y: Increased production of transforming growth factor-ß2 from cultured human retinal pigment epithelial cells by photocoagulation. Invest Ophthalmol Vis Sci 35:4645-4652, 1994 24. Gabel VP, Birngruber R, Hillenkamp F: Visible and near infrared light absorption in pigment epithelium and choroid. In: Shimizu K, Oosterhuis JA (eds) International Congress Series No. 450, pp 658-662. XXIII Concilium Ophthalmologicum, Kyoto. Amsterdam/Oxford: Excerpta Medica 1978 25. Birngruber R: Thermal modeling in biological tissues. In: Hillenkamp F, Pratesi R, Sacchi CA (eds) Lasers in Biology and Medicine, pp 77-97. New York, NY: Plenum Publ Corp 1980 26. Roider J, Birngruber R: Solution of the heat conduction equation. In: Welch AJ, Van Gemert M (eds) Optical-Thermal Response of Laser Irradiated Tissue, pp 385-409. New York, NY: Plenum Press 1995 27. Anderson RR, Parrish JA: Selective photothermolysis: precise microsurgery by selective absorption of pulsed radiation. Science 220: 524-527, 1983 28. Roider J, Hillenkamp F, Flotte T, Birngruber R: Microphotocoagulation: selective effects in biological tissue using repetitive short laser pulses. Proc Nat Acad Sci US 90:8643-8647, 1993 29. Brinkmann R, Hüttmann G, Rögener J, Lin C, Roider J, Birngruber R: Origin of RPE-cell damage by pulsed laser irradiance in the ns to µs time regime. Lasers Surg Med 27:451-464, 2000 30. Roider J, El-Hifnawi E, Birngruber R: Bubble formation as primary interaction mechanism in retinal laser exposure with 200 ns laser pulses. Laser Surg Med 22:240-248, 1998 31. Roider J, Michaud N, Flotte T, Birngruber R: Histologie von Netzhautläsionen nach kontinuierlicher Bestrahlung und nach selektiver Mikrokoagulation des retinalen Pigmentepithels. Ophthalmologe 90:274-278, 1993 32. Roider J, Brinkmann R, Wirbelauer C, Laqua H, Birngruber R: Subthreshold (RPE) photocoagulation in macular diseases: a pilot study. Br J Ophthalmol 84:40-47, 2000 33. Roider JR, Brinkmann R, Wirbelauer C, Laqua H, Birngruber R: Retinal sparing by selective RPE photocoagulation. Arch Ophthalmol 117:1028-1034, 1999 34. Wallow IH, Tso MOM: Repair after xenon arc photocoagulation. 2. A clinical and light microscopic study of the evolution of retinal lesions in the rhesus monkey. Am J Ophthalmol 75:610-626, 1973 35. Delori FC, Dorey CK, Staurenghi G, Arend O, Goger DG, Weiter JJ: In vivo fluorescence of the ocular fundus exhibits retinal pigment epithelium lipofuscein characteristics. Invest Ophthalmol Vis Sci 36(3) 718-729, 1995 36. Borland RG, Brennan DH, Marshall J, Viveash JP: The role of fluorescein angiography in the detection of laser-induced damage to the retina: a threshold study for Q-switched, neodymium and ruby lasers. Exp Eye Res 27:471-493, 1978 37. Machemer R, Laqua H: Pigment epithelium proliferation in retinal detachment (massive periretinal proliferation). Am J Ophthalmol 80:1-23, 1975 38. Katz ML, Eldred GE: Retinal light damage reduces auto-
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Selective retinal pigment epithelium laser treatment fluorescent pigment deposition in the retinal pigment epithelium. Invest Ophthalmol Vis Sci 30:37-43, 1989 39. Friberg TR, Karatza EC: The treatment of macular disease using a micropulsed and continuous wave 810-nm diode laser. Ophthalmology 104:2030-2038, 1997 40. Olk RJ, Friberg TR, Stickney KL, Akduman L, Wong KL,
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129 Chen M, Morse L, Garcia CA, Mush DC: A randomized pilot study comparing subthreshold vs threshold diode laser photocoagulation (PC) in the reduction of drusen associated with age-related macular degeneration. Invest Ophthalmol Vis Sci (Suppl) 38(4):18, 1997
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Confocal microscopy of the eye Charles J. Koester Department of Ophthalmology, Columbia University, New York, NY, USA
Keywords: optical sectioning, pinholes, confocal slits, confocal scanning laser ophthalmoscope, image acquisition, enriched adaptive optics
Abstract Confocal microscopy of the living eye is becoming a valuable tool for diagnosis of retinal and corneal disease, as well as for monitoring and measuring corneal changes post-operatively. It is also useful for studying variations with time, such as nerve growth, keratocyte activation, and endothelial cell changes. The instrumentation is evolving, in response to challenges such as involuntary eye motion, the need for accurate depth and thickness measurements, and the desire for high quality images under a large range of specimen thicknesses, and variations in optical properties: index of refraction, absorption and light scattering.
Introduction Confocal microscopy of the living eye is becoming a valuable tool for diagnosis, and for monitoring and measuring corneal changes post-operatively. It is also useful for studying variations with time, such as nerve growth, keratocyte activation, and endothelial cell changes. The instrumentation is evolving, in response to challenges such as eye motion, the need for accurate depth and thickness measurements, and the desire for high quality images under a large range of specimen thicknesses, and variations in optical properties: index of refraction, absorption, and light scattering. The development and applications of confocal microscopy have been motivated by the ability of the instrument to perform optical sectioning. That is, to illuminate and image one layer of the specimen at a time, by excluding from the final image most of the light scattered or reflected from other layers. In the eye, it has proved useful in the cornea, conjunctiva, iris, lens, and retina, for its ability to visualize cells and subcellular details in tissue that produces significant backscattering of the incident light. The
thickness of the optical section, the layer that is free of light scattered or reflected from out-of-focus regions, is the measure of the effectiveness of the system. The initial work in this area was done by Marvin Minsky,1 who wanted to see the interconnections between closely packed, interwoven cells in tissue of the central nervous system. The principle he employed is one that is still used today. In the illumination system of a light microscope, a pinhole is placed at a point that is confocal to the focal plane, so that a bright point of focused light is formed at the desired depth in the specimen. Light transmitted and scattered at this point is re-imaged at a second pinhole by the objective lens of the microscope. The two pinholes are ‘confocal’ because they share a common focal point. Light that passes through the second pinhole is primarily light from the first pinhole that has travelled directly to its image in the specimen and then directly to the second pinhole, without significant scattering in the specimen. Light that is scattered from other points in the specimen will not be imaged at the second pinhole, so that little or none will pass through. This is the essence of optical sectioning: eliminating scattered light from planes other than the object plane. However, it is necessary to add a method of scanning in order to produce a two-dimensional image. Minsky chose to use an x,y scan of the object, and to use a detector behind the second pinhole to generate a signal to the display. The first display was an available radar screen (early 1950s). Minsky, who at the time was a medical student, described this configuration as “an elegant, symmetrical geometry: a pinhole and an objective lens on each side of the specimen”. He went on to design a system of confocal imaging for reflected light.
Address for correspondence: Charles J. Koester, PhD, 60 Kent Road, Glen Rock, NJ 07452, USA. e-mail:
[email protected] Lasers in Ophthalmology – Basic, Diagnostic and Surgical Aspects, pp. 131–141 edited by F. Fankhauser and S. Kwasniewska © 2003 Kugler Publications, The Hague, The Netherlands
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Fig. 1. The Tandem scanning microscope. Light from the source passes through a group of pinholes (indicated by the shaded area on the insert diagram), is transmitted through the beam-splitting mirror (oval), and is focused by the objective lens to the plane of focus. Light returning from the plane of focus is reflected from the beam-splitter mirror and forms an image at the under surface of the disc. Light that has been reflected or scattered from the plane of focus will form a pattern of bright spots that will be transmitted by the corresponding holes in the disc. As the disc rotates, the spiral array of pinholes will scan the full length and width of the image, which is then transmitted to the eye or camera. Light that is scattered from out-of-focus planes in the specimen will not be in focus when it reaches the disc on the return trip, and only a small fraction will pass through the pinhole array. (Reproduced from PetráÍ M et al.2 by courtesy of the publisher.)
The next major step was taken by PetráÍ and Hadravský2 when they incorporated a Nipkov disc3 in an incident light microscope (Fig. 1). This not only eliminated the need to physically scan the specimen, but it also produced an image that could be viewed directly by eye and recorded in color. With precision fabrication of the disc, the image can be effectively free of visible scan lines. The system is quite effective because of the inverse square law, i.e., for a region of the object that is not exactly at the focal plane, light that passes through the image of a pinhole will be reduced in illuminance as the inverse square of the distance from that image. When the light is reflected or scattered back into the imaging system, the receiving pinhole will also be at a distance from the image of the outof-focus specimen, and the light passing through the receiving pinhole will again be reduced by the inverse square law. This design was adapted for the eye by the Tandem Scanning Corp., and more recently by Advanced Scanning, Ltd. Cavanagh, Jester, Petroll, and col-
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Fig. 2. Corneal images from the Tandem scanning confocal microscope in the through-focus mode. A: Epithelial cells, corresponding to peak A in the graph. B: Basal-epithelial nerve plexus corresponding to peak B. C: Anterior layer of keratocyte nuclei (peak C). D: Stromal nerve corresponding to position D. E: Endothelial cell layer, corresponding to peak E. F: Three-dimensional reconstruction. G: Image intensity as a function of depth in the cornea. (Reproduced from Li et al.4 by courtesy of Swets & Zeitlinger)
leagues4,5 have developed further methods and instrumentation for quantifying image data and for improved imaging. Their through-focus imaging technique facilitates the recording of cells, nerves, haze, and pathology at all depths in the cornea (Fig. 2). As illustrated at G in Figure 2, a plot of the intensity variation during the scan provides a quantitative comparison of backscattered light at various depths. An alternative to pinholes in a confocal system
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Fig. 3. The slit scanning confocal microscope. The illuminated slit S1 is imaged by objective lens L3 to the plane of focus. The Aperture with its central divider strip defines the shape of the incident light beam from slit S1 through the lower half of the objective lens L3. The rotation of mirror M2 causes the illuminated image of the slit to scan across the plane of focus in an oscillatory motion. (Reproduced from Koester12 by courtesy of the publisher.) Light returning from the plane of focus passes through the upper half of lens L3 and reflects from the second facet of mirror M2. This second reflection from the oscillating mirror stops the scanning motion of the light beam; the beam then passes through slit S2. The function of slit S2 is to block light that has scattered or reflected from planes proximal or distal to the plane of focus. After passing through S2 and reflecting from the third facet of mirror M2, the light is focused by lens L5 at the focal plane of the camera. The final image is formed as the slit image scans back and forth across the film or CCD array. (Adapted from: Pauley JB (ed) Handbook of Biological Confocal Microscopy, 2nd edn, 1995, by courtesy of Plenum.)
is to illuminate and image through narrow slits. Gullstrand,6 in 1911, introduced slit illumination of the cornea. Oblique illumination with the narrow beam from the slit provided a form of optical sectioning as the thin sheet of light passed obliquely through the cornea. Under optimum conditions, it was possible to visualize endothelial cells because the oblique illumination separated the bright reflex of the anterior cornea surface from the much weaker image of the endothelial cell layer. In 1968, David Maurice7 extended this principle to the microscope by imaging an illuminated slit onto the endothelial cell layer through one half of a microscope. The other half of the microscope was used to view the portion of the endothelial cell layer that was not obscured by the light reflected from the anterior cornea surface. First used in vitro, the principle later became widely used clinically as the specular microscope. Not satisfied with the narrow image that could be obtained in this manner, in 1974 David Maurice8 developed a double slit scanning microscope for the in vitro examination of tissue. First the specimen was mounted on a moving stage, under an illuminated slit. Secondly, the light that was reflected specularly from the surface was imaged at the plane
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of a stationary slit. Light passing through the second slit was focused on photographic film that was moving past the slit synchronously with the movement of the specimen. The jaws of the second slit served to block the reflex from the front surface of the cornea, but passed the light reflected from deeper layers. The specimen and the film were synchronously scanned so that the image of the specimen was laid down on the film continuously. The system produced excellent wide field images of glarefree images. This demonstrated that a double slit system could also be used for confocal imaging. It was not, however, suitable for in vivo studies. Other confocal designs using slits were developed by Svishchev in 1969,9 Baer in 1970,10 and Brakenhoff and Visscher in 1992.11 None of these developments resulted in instrumentation for the eye, in vivo, but they contributed to the advancement of confocal microscopy in general. The motivation for the slit scanning confocal microscope12 was to enable the specular microscope to become a wide field device (Fig. 3). As in the standard specular microscope, an illuminated (stationary) slit, S1 is imaged on the endothelial cell layer at the plane of focus, using slightly less than one half of the objective aperture. An oscillating mir-
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Fig. 4. Stromal nerve junction, with filaments to keratocytes. Slit scanning confocal microscope, NA 0.8. (Photo by James Auran.)
ror, M2, is used to scan the illuminated slit image across a one-millimeter width of endothelium. Light reflected from this illuminated strip is directed through the other half of the microscope objective, to the second facet of the oscillating mirror. The second reflection reverses the action of the first reflection, stabilizing the light so that it can pass through the second slit, S2. The other function of the second slit is to block light that has been reflected or scattered from portions of the specimen that are not at the focal plane. Finally, the light that has passed through the second slit is sent to the third facet of the oscillating mirror M2. Reflection from this rotating facet causes the beam to scan across the image plane of the camera (film or CCD array) and to lay down the image strip by strip. The thickness of the optical section is determined by the numerical aperture (NA) of the objective lens used, and the width of the confocal slits. When a very thin optical section is required, to visualize small details in the presence of scattered light from nearby structures, narrow slits will produce a thinner optical section, i.e., greater rejection of scattered light from adjacent layers. However, if the desired details are weakly scattering and there is some degree of eye movement, wider slits will provide greater illumination, thereby minimizing the exposure time. Thus, the variable slit width can be utilized to allow the examiner to select the best section thickness, depending on the degree of detail in the image and the amount of eye movement of the subject. Figure 4 is an image from the stroma of a normal subject. The nerve junction exhibits filaments extending to keratocytes. Figure 5 depicts another area where optical sectioning microscopy can be of value: the conjunctiva. Blood vessels, with associated nerves and white cells, are clearly imaged in spite of the density of the surrounding tissue and the close proximity of the sclera.13 Figure 6 illustrates the ante-
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Fig. 5. Conjunctiva: blood vessels and associated white cells. Bar: 100 µm. The slit scanning confocal microscope, NA 0.75. (Photo by Aryan Shayegani et al.13)
Fig. 6. Bovine lens, ex vivo. Anterior lens fiber cells. Slit scanning confocal image, NA 0.75. (Photo by Norman Kleiman.14)
rior region of a bovine lens, ex vivo utilizing the slit scanning confocal microscope, NA 0.80.14 Figure 7 is a human iris, in vivo, taken with the extended range objective, NA 0.80, on the slit scanning confocal microscope. A widely used, important confocal system for the eye is the scanning laser ophthalmoscope, applied in the confocal mode.15 (See Fig. 8.) A focused, low power laser beam is reflected from a scanning mirror and projected into the eye and focused on the retina. On returning from the retina and emerging from the eye, it reflects again from the scanning mirror. This stabilizes the beam, so that it passes through a pinhole to a detector. The scanned image is then displayed on a monitor. The pinhole serves
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Fig. 7. Human iris, in vivo. The pupillary margin, including the pigment ruff and anterior border layer, in a 37-year-old individual with a green iris. Slit scanning confocal microscope, NA 0.75. (Reproduced from Kummer et al.14 by courtesy of the publisher.)
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Fig. 9. Human fundus, imaged by the scanning laser ophthalmoscope in the confocal mode. (Reproduced from Webb15 by courtesy of the publisher.)
Fig. 10. Schematic diagram of the optical system of the Confoscan3 confocal microscope.
Fig. 8. The scanning laser ophthalmoscope. Light from laser, L, passes through an acousto-optical modulator, AO, and a beam expander, B, then reflects from a multifaceted rotating mirror, H, which imparts a horizontal scan. The vertical scan is added by mirror, V. The laser beam is focused by the optics of the eye to the retina and is scanned over a small area. Light returning from the retina (dashed lines) follows the reverse path, bypasses the small mirror, T, passes through pinhole aperture, A, and is detected at D. The signal then goes to the display. The system is confocal because the laser beam behaves as if it had come from a point source. It can also be used in a non-confocal mode by removing the pinhole aperture, A. (Reproduced from Koester12 by courtesy of the publisher.)
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to eliminate or greatly reduce the light that is scattered from regions in the eye other than at the focus of the laser beam. Figure 9 illustrates the human fundus as imaged in the confocal mode of the scanning laser ophthalmoscope. Because the numerical aperture of the system is limited by the eye’s pupil diameter (NA = 0.08 for a 3-mm pupil, or 0.22 for an 8-mm pupil), the resolution will not be as high as that of a high NA microscope objective that is used on the cornea, for example. Also, the optical section thickness will be greater than that from a high NA objective. Nevertheless, the confocal design improves the image when there is light scattering material in the paths of the incident and reflected light and delivers sharply defined images of retinal structures. Masters and Bohnke16 utilized confocal microscopy to construct three-dimensional, high resolution images of the cornea, in vivo. They have also utilized confocal microscopy to demonstrate that long-term contact lens wear can produce degenerative microdot deposits in the corneal stroma.17 A confocal system based on synchronously scanned slits has been described by Wiegand et al.16 Figure 10 illustrates the principle, as adapted to the Nidek Confoscan microscope. In the microscope objective, the illumination and imaging light paths are separated, and as they pass into the cornea. By synchronizing the scanning of the slits and the
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Fig. 11. Post-LASIK deep stromal nerve. Confoscan3. (Reproduced with permission.)
Fig. 12. LASIK interface (very low keratocyte density). Confoscan3. (Reproduced with permission of Nidek.)
video image cycle it was possible to eliminate effectively any blur due to eye movements. The Nidek ‘Confoscan3’ utilizes these principles. The objective lens does not touch the cornea, but utilizes the optical immersion principle with gel and the tear layer. Image recording is continuous on a digital camera, with instant storage of several hundred frames. The width of the slits, together with the NA of the objective, determines the thickness of the optical section. It is estimated to be 5-10 µm in the captured images. Figure 11 shows a post-LASIK deep stromal nerve. Figure 12 shows a LASIK interface. With the Confoscan3 system, the depth of any particular feature in the cornea can be deduced from a series of 350 images obtained in a fast scan through the cornea. If eye movement is detected, or suspected, the depth scan can be repeated. Applications of confocal microscopy in refractive surgery Specific applications of confocal microscopy in refractive surgery are listed and described in a comprehensive review paper by Ambrosio and Harrison.19 Wound healing has been observed and quantified to
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determine thickness of the flap, total corneal thickness, and changes over time. Vesaluoma et al.20 documented flap alignment and followed the wound healing at the margin of the LASIK flap. The authors were also able to follow keratocyte myofibroblast transformation and activation, and proliferating fibroplastic cells. These wound healing events had been documented previously in primate models by Jester et al.21 Complications in wound healing with LASIK were studied with confocal microscopy by Vesaluoma et al.22 Post PRK wound healing and cornea nerve growth were studied by Linna et al.23 Complications following LASIK, e.g. particles at the interface, were documented by Kaufman et al.24 The high resolution and optical sectioning capabilities of confocal microscopy facilitate the study of keratocytes and their behavior following laser surgery. Patel et al.25 quantified the density of keratocytes as a function of depth in normal eyes, a statistically significant decrease with age, and a concentration of keratocytes in the anterior stroma in most but not all subjects. Moller-Pedersen et al.26 observed large numbers of wound healing keratocytes with higher reflectivity and haze development, all related to refractive instability. Sub-epithelial and stromal nerves are particularly well imaged by high numerical aperture confocal microscopy. Linna et al.27 studied re-innervation after LASIK and correlated changes in neurotropic epitheliopathy with nerve regeneration into the flap. The growth pattern of basal epithelial nerves has been documented by Auran et al.28 Basal epithelial cell and unidentified cellular elements were also observed to move relative to established landmarks. Other applications of confocal microscopy, including detection and diagnosis of corneal disease have been presented by Cavanagh et al.,29,30 by Auran et al.31 and by Florakis et al.32 Kleiman and Auran33 recently reported on the dynamics of Langerhans cells in traumatized cornea. The cells were identified by their morphology. What future developments and refinements can be expected in confocal microscopy of the eye? A number of improvements and new developments are possible. The list should include the following: 1. Objectives with higher numerical aperture, for better resolution, improved optical sectioning, and greater light gathering power, particularly for weakly scattering details. Higher NA can improve both the optical sectioning and resolution of the system. One difficulty is that higher NA usually is accompanied by: a. shorter working distance; and b. smaller range of depth in the specimen within which the image quality is ‘diffraction limited’ (i.e., the image is essentially as perfect as diffraction and the wavelength of the light will permit).
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Confocal microscopy of the eye The problem of smaller range of depth requires a change in the usual design of an objective lens. A well corrected objective designed for water immersion, NA 0.7, will give a diffraction limited image through the full 0.5-mm thickness of the cornea, but the same is not true for NA 0.8. (This problem is the same as that for cover glasses with high NA objectives.) For examination of the lens, one problem is that the thickness of the lens is too great to obtain a good image at all depths using a standard design, high NA objective. In addition, there is variation in the depth of the anterior chamber between patients, so that an even greater range of focal depth is needed. In order to utilize objectives with NA 0.8 and greater, a variable focus design can be utilized. The basic design is analogous to the familiar zoom lens system in that one optical element is moved axially relative to the other. But, with the zoom system, the magnification is changed while the image and object foci remain fixed. In the variable focus design, the position of the focal plane in the object is changed, while the image focal plane remains fixed. With the variable focus system, NA 0.8, the full thickness of the cornea can be examined without loss of resolution. And, for patients with various anterior chamber depths, the full thickness of the lens can be examined without loss of resolution.34 There is a small change in the magnification as the focus is changed, just as there is with a conventional lens used for the same purpose. 2. Improved (thinner) optical sectioning. Measured values of optical section thickness, provided by manufacturers. The confocal slit, divided aperture systems for the eye have generally used equal, crescent shaped apertures for illumination and imaging (Figs. 3 and 10). For non-laser sources, e.g., arc lamps, this configuration provided the maximum efficiency for utilization of the available light. However, when a laser can be used, the full power of the laser can be injected through a small portion of the objective aperture. When that portion is close to the edge of the objective aperture, two benefits are available. First, a larger portion of the objective aperture is then available for the imaging rays. This leads to better resolution in the image, as well as a brighter image and/or shorter exposure time. At the same time, the optical sectioning capability can be increased, because the angular spacing between the illumination and imaging ray bundles is greater (due to the small diameter of the laser beam). This improvement could also be feasible with the combination of a bright non-laser source and a very sensitive detector system. The optical section thickness provided by each microscope needs to be well defined, so that comparisons can be made between various methods. With the confocal slit system there is a calculable depth (relative to the focal plane) beyond which no incident light can be scattered or reflected (by the specimen) into the imaging light beam. The calculation requires
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137 knowledge of the locations of focal planes and the aperture stop of the objective. However, this calculation does not include the effects of stray light that may arise in the optics of the system due to reflections from lenses or scattering from surfaces. With pinhole systems, a precise value for the optical section ‘thickness’ cannot be calculated. The intensity of the incident light decreases as the inverse square of the distance from the image of the pinhole, so the optical section does not have a well-defined thickness. For light scattered in the specimen, and travelling back through the objective, the fraction passing through the pinhole on the imaging side, the same inverse square law applies. If the optical section profile is to be calculated, the design of the optics needs to be known, and the precision of the pinhole locations and diameters are important. Therefore, it would be useful for manufacturers to measure (and publish) the optical sectioning characteristics of the instruments. Not only is it more practical than calculating a theoretical value, but also it is more realistic, since the measurement will include light reflected and scattered at lens and other internal surfaces in such a way that a fraction of the light finds its way back to the image plane. The method for measurement is simple, in principle. A reflector, preferably diffusely reflective, is positioned at the focal plane, mounted on a platform that can be moved along the axis of the microscope.35 The light that returns to the focal plane is then measured as a function of the distance that the reflector is moved away from the focal plane. For immersed systems, the fluid layer must of course be maintained between the objective and the reflecting target. This technique has also been useful in locating and identifying sources of stray light in the microscope, other than back scattering from the specimen. For example, when the measured light intensity no longer decreases as the distance to the target increases, it becomes apparent that there is stray light in the system coming from a source other that the reflecting surface. 3. Enriched image acquisition, in the presence of the inevitable eye movement. Involuntary eye movements provide a major challenge for the user of a high magnification device such as the confocal microscope. Contacting the cornea with a polished glass surface similar to that of the applanation tonometer can help to stabilize the eye, but it does not prevent eye movement. For many patients, non-contact, using a gel or a fluid, would probably be their choice. In both cases the problems for the examiner are substantial. At high magnification, locating and then focusing on the detail of interest, the task is somewhere between challenging and impossible. One approach is to turn on the video camera or the recording CCD camera, and guide the microscope slowly and repeatedly through the volume of interest, to provide the opportunity to catch an image that is not blurred by motion. Reviewing stored CCD images or video tape in a search for usable images can be done at one’s leisure.
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Fig. 13. The basal epithelial nerve plexus. The pachymeter reading of 0.040 mm represents the depth of the (in focus) nerves from the surface of the cornea. The depth measurement is derived from the focus setting of the objective lens. Slit scanning confocal microscope. (Photo by C. Koester.)
Digital cameras and image recording may offer another possibility. When the examination begins, the camera can be turned on so that each image is stored in temporary storage (T). When the examiner sees a good image, the SAVE button (or foot switch) is pressed. That causes the images that have been stored in the previous few seconds to be transferred to a longer term digital storage (L). As long as the SAVE button is being pressed the new images will be stored in (L) as well. When the SAVE button is released, the new images are again sent to the temporary storage (T). When this temporary storage is filled, the next images can be written over the older images in (T), and the cycle can be repeated. Variations on this theme are also possible. Other information could be added to the saved images: the examiner’s spoken comments at the time the SAVE button is activated, data available from the instrument such as focus setting, time and date, objective being used, patient ID, etc. The saved file would then be an enriched sample of images from the examination, not totally dependent on the reaction time of the examiner who is trying to catch the sharp image before the next eye movement. For both contact and non-contact instrumentation it will be helpful to have image recording systems that do not require reviewing the entire examination on tape (or on digital media) in order to find the acceptable images.
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4. Improvement in the quality of acquired images by optimizing the alignment of the microscope objective to the cornea. The instrument will, of course, record the optically sectioned image of whatever is at the focal plane. But the quality of that image will depend on the homogeneity, indices of refraction and thicknesses of the media between the objective tip and the focal plane, as well as any movement of the specimen that occurs during the exposure. Contact with the cornea eliminates one source of image degradation: the astigmatism that is introduced when a diverging bundle of rays passes at an angle through a surface separating two media of different indices of refraction. When the microscope is used in the non-contact mode (with tear layer or gel), it is not easy to assure that the axis of the microscope is aligned perpendicular to the surface. For example, if the microscope axis was first aligned to the visual axis, then moved without rotation 2 mm to the right, the tilt of the cornea surface at that point would introduce a 50% increase in the blur circle diameter for a point located at the endothelial surface, and the image would no longer be diffraction limited. The magnitude of this effect increases with the NA of the objective. It also increases in proportion to the difference in index of refraction between the cornea and the medium separating the objective tip and the cornea. Hence, one procedure is to be meticulous about aligning the microscope perpendicular to the cornea surface. Having the subject rotate the eye to the new area to be examined is not sufficient; the radius of the cornea is less than the distance from cornea to the center of rotation of the eye, so that the angle between the cornea surface and the front surface of the contact element increases as the eye is rotated. Another procedure is to align (approximately) the microscope to the cornea by eye, and then applanate with the tip of the microscope to assure contact and therefore alignment. Another possible method is discussed in section 6, below. 5. Recording the location of the image. The location of the image: the depth and to the extent possible, the x,y position on the cornea is potentially important information: for returning to the same location at a later date, for reporting results, for documenting any change in thickness of the cornea, etc. In instruments using the contact mode, the depth information can be made available from the focus setting. An example of a pachymeter reading that is recorded on photographic film beside the cornea image is shown in Figure 13. For systems used in the non-contact mode, a fast axial scan can locate the focus relative to two reference depths, e.g., the endothelial and epithelial surfaces. (See Petroll et al.4) Because of the frequency of eye movements, the measurement of depth should be based on two such scans. If the two fast scans show a difference in depth of the image, this may indicate that there has been eye motion between the two scans,
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Confocal microscopy of the eye and the depth measurement should be flagged as being unreliable. The measurement of x,y position is open to good ideas. An accurate determination of x,y location would be useful, but just the ability to return to the general area would be satisfactory in many cases. In one investigation, it was found that distinctive configurations of nerve junctions in the anterior cornea could be recognized and used to return to the same portion of the cornea a few days later.19 However, there were other, stationary landmarks in the cornea, and the nerve structures were observed to move during a period of several days relative to these stationary landmarks. Conclusion: Nerves in the anterior cornea, while quite visible and distinguishable, can be used as landmarks for relocating a position on the cornea for only a few days. Stationary landmarks that were found useful in this study were: the corneoscleral limbus, anterior stromal nerves (including the subepithelial nerve plexus), entry points where nerves emerged from Bowman’s layer into the epithelium, stromal banding and basal lamina ridges induced by pressure from the flat contact element, and the nuclei of nonneural stromal cellular elements.28,36 An approximate position of the contact element on the cornea could be recorded by a simple, small photographic or digital camera mounted on the head rest. It could image the contact element, the lids and the visible portion of the limbus, which would probably be satisfactory for purposes of relocating the area of interest. 6. A gel that matches the index of refraction of the cornea (n = 1.376). Such a gel would eliminate not only the astigmatism that is generated by a tilted cornea surface, but also the focusing power and spherical aberration caused by a spherical surface between two media with different refractive indices, including the cornea. If the dispersion (change of index with wavelength) also matched that of the cornea, chromatic aberration at that surface would also be eliminated. Of course, the gel should be viscous enough to stay in place on the cornea, but not so viscous that it would change the shape of the cornea during a movement of the eye or of the contact element of the objective lens. A gel used by Wiegand et al.18 has an index of 1.350, and is an improvement over the tear layer which has an index of about 1.34. 7. Image processing, particularly for motion blurred images. Motion blurring in the plane of focus can now be corrected, post-exposure, by deconvolution methods. Deconvolution is a mathematical procedure that calculates the most likely object light distribution that would form the observed image, given the blur (the point spread function) produced by the lens when a point source is imaged. If the point spread function of the microscope is known, then deconvolution is
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139 most effective. If not, a method called ‘maximum likelihood deblurring’ can be used.37 Confocal images generated by microscopes with thin optical sections should be excellent candidates for deconvolution, since they are relatively free from overlapping, out-of-focus images. 8. Reduction of stray light in the microscope. Confocal microscopy of nearly transparent tissue is particularly susceptible to stray light. The source of the stray light is often illumination light that is reflected from lens surfaces, from beam splitter surfaces, or scattered from imperfections or dust on mirror and lens surfaces, or inner walls and lens mounts. If this reflected/scattered light cannot be prevented from reaching the image plane of the detector (film, CCD, or eye), the contrast in the image and the optical sectioning capability will be compromised. The measurement of optical section thickness, as described above, is an effective technique for identifying sources of stray light. Another is to view the back aperture of the objective when the illumination light is on but there is no reflecting/scattering object at the focal plane. The back aperture of the objective should be appear black, as well as the rim of the aperture. 9. Confocal microscopy in other fields. Confocal microscopy in other fields, particularly in other in vivo applications, can be a source of advanced techniques and technology. Corcuff et al.38 and Rajadhyaksha39 used the technology to study the skin. Khanna et al.40,41 combined laser heterodyne interferometry and confocal microscopy to measure vibration of hair cells and membranes in the cochlea of animals, in vivo. 10. Contact versus non-contact in confocal microscopy. Normal eye motion is obviously a concern for high magnification by confocal microscopy. Contacting the cornea with a flat (or curved) end of the objective lens does not stop all eye motion, but it does establish the axial location of the epithelial surface. It also stops the motion of the cornea surface for periods of tenths of a second to several seconds, in most subjects. With non-contact, the eye is free to move in any of its normal modes, which have been characterized by Charman42 as tremor, drift, and microsaccades and saccades. Tremors have high frequencies, but the amplitudes are small. Saccades have a velocity of about 700E/sec or more. This means that the corneal surface can move at a linear velocity of 153 mm/sec, or about 150 µm during a 0.001-second exposure. During the period between large eye movements, nearly continuous eye movement is observed in the eyepiece or viewfinder of the camera, which makes it difficult to select the detail to be photographed and to focus on it before it moves. The use of objectives with higher NA can benefit the optical sectioning capability as well as resolution and image intensity.43 However, when the NA is in-
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creased, the requirements on the rest of the optical system, including the eye being examined become more stringent. For example, an objective, NA 0.7, with gel or saline between the objective and the surface of the cornea can focus through the full thickness of the cornea and give diffraction-limited (optically perfect) images. However, if the NA is increased to 0.8, the performance begins to drop at a depth of about 0.4 mm. (This problem is the same as the cover glass thickness problem in laboratory microscopy.) One way to overcome this obstacle is to use objectives that have been designed as variable focus lenses. These objectives have an internal moving component that allows the system to focus sharply over a range of thicknesses in the specimen. The first design for the eye allows the anterior surface of the lens to be examined in patients with anterior chamber depths from 2.4-4.0 mm, i.e., a range of variable focus of 1.6 mm, at NA 0.8.34 The above discussion assumes that the objective axis is aligned perpendicular to the surface of the cornea, as it will be when the objective lens is in contact with the cornea surface. If non-contact microscopy is used, there are several potential optical problems, in addition to eye movement. As discussed in section 4, degradation of image quality occurs when the surface of the cornea is tilted relative to the axis of the microscope. Furthermore, the quality of the image will depend on the separation between the cornea and the tip of the microscope. One way to overcome this problem would be to develop a compatible gel whose index of refraction (and dispersion) match that of the cornea (n = 1.376). The objective lenses would also need to be designed for immersion in this medium. 11. Adaptive optics. Pioneered by astronomers to improve the images of stars that are blurred by fluctuations in the atmosphere, adaptive optics are now being utilized to improve the resolution of retinal elements when observed through the subject’s cornea, aqueous, lens and vitreous.44 The aberrations of the eye’s optical system are first measured by sending a laser beam through the pupil, to the retina. The light returning through the pupil is then captured on an array detector. The data from this sensor are utilized to deform a mirror in such a way that the aberrations introduced by the eye can be significantly reduced. When the retina is to be photographed, the deformed mirror is incorporated into the imaging system, enabling a much improved image of the retina to be recorded.44 When the confocal microscope is used for examination of the cornea or lens, adaptive optics do not have the same advantages that they have in astronomy and retinal examination. Firstly, there is not a strongly reflective surface such as the retina that is close to the focal plane. Secondly, the surfaces that produce the aberrations (the corneal surfaces and lens) are relatively close to the detail that is to be photographed. Therefore, the correction that must be applied to the wavefront will depend on the location of the detail of
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interest relative to the aberrating regions. In this case however, the aberrating surfaces are relatively close to the details of interest, so that their effect on resolution in the cornea and lens will be less than their effect on retinal imaging. In summary, it appears to be likely that adaptive optics combined with high NA confocal microscopy will provide superb imaging of the retina. But the advantage of the combination for imaging the cornea and the lens is not so obvious. However, the image quality attainable with confocal imaging in the cornea and the lens is already excellent, and can become even better with objectives having higher NA, reduced stray light, and a greater range of focus. Conclusions Research studies using confocal microscopy have demonstrated that detailed images of normal and pathological corneas, lenses and retinas can be obtained. Changes over time, for example healing and nerve fiber growth, can be documented. Depths and thicknesses in the cornea can be measured. For clinical applications some important instrument capabilities are: the efficiency and accuracy of obtaining sharp images, acquiring depth information, and having the ability to return to the same location. These functions need to be optimized for the clinical situation. Just as the slit lamp went through many refinements before reaching its present mature design, the confocal microscope may require additional effort on the part of practitioners, instrument designers, and manufacturers in order to realize a fully clinical, stateof-the-art ophthalmic instrument. References 1. Minsky M: US Patent No. 3,013,467 (1961); Memoir on inventing the confocal scanning microscope. Scanning 10:128-138, 1988 2. PetráÍ M, Hadravský M, Egger D, Galambos, R: Tandemscanning reflected-light microscope. J Opt Soc Am 58:661664, 1968 3. Nipkow P: German Patent 30,105, 1884 4. Petroll WM, Cavanagh HD, Jester JV: Clinical confocal microscopy. Curr Opin Ophthalmol 9(4):59-65, 1998. Original figure in: Li HF, Petroll WM, Moller-Pederson T, Maurer JK, Cavanagh HD, Jester JV: Epithelial and corneal thickness measurements by in-vivo confocal microscopy through focusing (CMTF). Curr Eye Res 16:214-221, 1997 5. Gokmen F, Jester JV, Petroll WM, McCulley JP, Cavanagh HD: In vivo confocal microscopy through-focusing to measure corneal flap thickness after laser in situ keratomileusis. J Cataract Surg 28:962-970, 2002 6. Gullstrand A: Demonstration der Nerstspaltlampe. Heidelberg: Heidelberger Bericht 1911 7. Maurice DM: Cellular membrane activity in the corneal endothelium of the intact eye. Experientia 24:1094-1095, 1968 8. Maurice D: A scanning slit optical microscope. Invest Ophthalmol Vis Sci 13:1033-1037, 1974 9. Svishchev GM: Microscope for the study of transparent light-
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scattering objects in incident light. Opt Spectrosc 26:171172, 1969 Baer SC: US Patent 3,547,512 (1970). Proc SPIE 1139:99101, 1989 Brakenhoff GJ, Visscher K: Confocal imaging with bilateral scanning and array detectors. J Microsc 165:139-146, 1992 Koester CJ: Scanning mirror microscope with optical sectioning characteristics: applications in ophthalmology. Appl Opt 19:1749-1757, 1980 Shayagani A, Auran JD, Koester CJ: Clinical color confocal microscopy of the conjunctiva. Invest Ophthalmol Vis Sci 37(3):S317 (ARVO Abstract nr 1457), 1996 Kummer A, Kleiman NJ, Koester CJ: Scanning slit confocal microscopy of crystalline keratitis. Invest Ophthalmol Vis Sci 37(3):S1025 (ARVO Abstract nr 4709), 1996 Webb RH: Scanning laser ophthalmoscope. In: Masters BR (ed) Noninvasive Diagnostic Techniques in Ophthalmology, pp 438-450. New York, NY: Springer-Verlag 1990 Masters BR, Bohnke M: Three-dimensional confocal microscopy of the human cornea in vivo. Ophthalmic Res 33:125135, 2001 Bohnke M, Masters BR: Long-term contact lens wear induces a corneal degeneration with microdot deposits in the corneal stroma. Ophthalmology 104:1887-1896, 1997 Wiegand W, Thaer AA, Kroll P, Geyer OC, Garcia AJ: Optical sectioning of the cornea with a new confocal in vivo slit-scanning videomicroscope. Ophthalmology 102:568-575, 1995 Ambrosio R Jr, Harrison DA: Confocal microscopy in refractive surgery. Review of Refractive Surgery 2001:1-5, 2001 Vesaluoma M, Perez-Santonja J, Linna T, Alio J, Tervo T: Corneal stromal changes induced by myopic LASIK. Invest Ophthalmol Vis Sci 41:369-376, 2000 Jester JV, Petroll MW, Cavanagh HD: Corneal stromal wound healing in refractive surgery: the role of myofibroblasts. Progr Ret Eye Res 18:311-356, 1998 Vesaluoma MH, Petroll WM, Perez-Santonja JJ, Valle TU, Alio JL, Tervo TM: Laser in situ keratomileusis flap margin: wound healing and complications imaged by in vivo confocal microscopy. Am J Ophthalmol 130:564-573, 2000 Linna TU, Vesaluoma MH, Perez-Santonja JJ, Petroll WM, Alio JL, Tervo TM. Effect of myopic LASIK on corneal sensitivity and morphology of subbasal nerves. Invest Ophthalmol Vis Sci 41:393-397, 2000 Kaufman SC, Maitchouk DY, Beuerman RW: Interface inflammation after LASIK: Sands of the Sahara syndrome. J Cataract Refract Surg 24:1589-1593, 1998 Patel S, McLaren J, Hodge D, Bourne W: Normal human keratocyte density and corneal thickness measurement by using confocal microscopy in vivo. Invest. Ophthalmol Vis Sci 42:333-339, 2001 Møller-Pedersen T, Li HF, Petroll WM, Cavanagh HD, Jester HV: Confocal microscopic characterization of wound repair after photorefractive keratectomy. Invest Ophthalmol Vis Sci 39:487-501, 1998 Linna T, Tervo T. Real time confocal microscopical observations on human corneal nerves and wound healing after excimer laser photorefractive keratectomy. Curr Eye Res 16:640-649, 1997 Auran JD, Koester CJ, Kleiman NJ, Rapaport R, Bomann JS, Wirotsko BM, Florakis GJ, Koniarek JP. Scanning slit confocal microscopic observation of cell morphology and movement within the normal human anterior cornea. Ophthalmology 10:33-41, 1995 Cavanagh HD, Petroll WM, Alizadeh H, et al: Clinical and diagnostic use of in vivo confocal microscopy in patients with corneal disease. Ophthalmology 100:1444-1454, 1993
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141 30. Cavanagh HD, El-Agha MS, Petroll WM, Jester JV: Specular microscopy, confocal microscopy, and ultrasound biomicroscopy: diagnostic tools of the past quarter century. Cornea 19:712-722, 2000 31. Auran JD, Starr MB, Koester CJ, LaBombardi VJ: In vivo scanning slit confocal microscopy of Acanthamoeba keratitis. Cornea 13:183-185, 1994 32. Florakis GJ, Moazami G, Schubert H, Koester CJ, Auran JD: Scanning slit confocal microscopy of fungal keratitis. Arch Ophthalmol 115:1461-1463, 1997 33. Kleiman NJ, Auran JD: High resolution in-vivo confocal microscopy of Langerhans cells in the ocular mucosa, skin, and cornea. Invest Ophthalmol Vis Sci 42(4):S891 (ARVO Abstract nr 1415), 2001 34. Koester CJ, Kleiman NJ, Kummer AE, Auran JD, Roberts JE: High resolution, extended range optical system for confocal microscopy of the human iris. Invest Ophthalmol Vis Sci (Suppl) 5496, 1990 35. Khanna SM, Koester CJ: Optical sectioning characteristics of the heterodyne interferometer. Acta Otolaryngol (Stockh) Suppl 467:61-67, 1989 36. Auran JD, Koester CJ, Rappaport R, Florakis, GJ: Wide field scanning slit in vivo confocal microscopy of flattening-induced corneal bands and ridges. Scanning 16:182-186, 1993 37. Holmes TJ: Light microscopic image reconstructed by maximum likelihood deconvolution. In: Pawley JB (ed) Handbook of Biological Confocal Microscopy, 2nd edn, ch 24, pp 389-402. Plenum, New York, 1989 38. Corcuff P, Gonnord G, Pierard GE, Leveque JL: In vivo confocal microscopy of human skin: a new design for cosmetology and dermatology. Scanning 18:351-355, 1996 39. Rajadhyaksha M: In vivo confocal scanning laser microscopy of human skin: melanin provides strong contrast. J Invest Dermatol 104:946-952, 1995 40. Khanna SM, Koester CJ, Van Netten SM: Integration of the optical sectioning microscope and heterodyne interferometer for vibration measurements. Acta Otolaryngol (Stockh) Suppl 467:43-49, 1989 41. Willemin JF, Khanna SM, Dandliker R: Heterodyne interferometer for cellular vibration measurement. Acta Otolaryngol (Stockh) Suppl 467:35-42, 1989 42. Charman WN: Optics of the eye. In: Bass M (ed) Handbook of Optics, 2nd edn, vol 1, ch 24, p 35. Optical Society of America. New York, NY: McGraw-Hill 1995 43. Koester CJ, Auran JD, Rosskothen JD, Florakis GJ: Clinical microscopy of the cornea utilizing optical sectioning and a high-numerical-aperture objective. J Opt Soc Am A 10:1670-1679, 1993 44. Liang J, Williams DR, Miller DT: Supernormal vision and high resolution retinal imaging through adaptive optics. J Opt Soc Am A 14:2884-2892, 1997
Addendum Status of instruments cited • The Confoscan3 microscope is manufactured by Nidek. • The scanning laser ophthalmoscope is manufactured by Heidelberg Instruments. • The slit scanning confocal microscope is a research model at Columbia University, Department of Ophthalmology. The optical sectioning microscope is a research instrument at the Department of Otolaryngology, Columbia University. • The Tandem scanning microscope is not presently available.
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Imaging in ophthalmology Joel S. Schuman, Zinaria Y. Williams, James G. Fujimoto and Lelia Adelina Paunescu New England Eye Center, New England Medical Center, Boston, MA, USA
Keywords: confocal imaging, OCT, Heidelberg tomograph
Introduction
Heidelberg Retina Tomograph
Since Helmholtz first peered inside the human eye more than 150 years ago, the subjective evaluation of the fundus has played a central role in the assessment of eye health and disease. Over the past two decades, technologies have been developing for the objective evaluation of ocular structures. These devices promise quantitative measurement of the retina and optic nerve with high degrees of precision and reproducibility. Optical coherence tomography (OCT) and confocal scanning laser ophthalmoscopy (CSLO) are two methods that have emerged to permit accurate analysis of the internal structures of the eye.1 This chapter will focus on these two imaging techniques. Both are non-contact, non-invasive imaging systems.
The Heidelberg Retina Tomograph (HRT; Heidelberg Engineering GmbH, Heidelberg, Germany) has two confocal scanning laser ophthalmoscope systems that are commercially available. The original HRT I is an optical scanning system for acquisition and analysis of the posterior pole. HRT is typically used to assess the ONH by using a diode laser of 670 nm to scan a three-dimensional image from a series of optical sections at 32 consecutive focal planes. The confocal scanning optical microscope reconstructs an ONH image by bringing a series of two-dimensional digitized images into registration. The registration process corrects for microsaccades that may occur during image acquisition.2 The topography image is composed of 256 × 256 pixel elements, each of which is a measurement of height at its corresponding location. The optical transverse resolution is approximately 10 µm, whereas the axial resolution is about 300 µm. In HRT I, the transverse field of view can be 10 × 10°, 15 × 15°, or 20 × 20°. In current clinical practice, three scans of each eye are taken and then averaged to create a mean topography image. Images can be obtained through undilated pupil, but dilation will improve image quality in patients with small pupils and cataracts. Reproducibility is best in undilated eyes.3 The printed report displays a topographic image and a reflectivity image of the ONH and its contour line, ONH stereometric parameters, and a mean-height contour of the peripapillary retina (Figs. 1A and B). The HRT was reported to be more sensitive than clinical assessment in detecting early glaucomatous disc changes in a study that evaluated 72 normal
Confocal scanning laser ophthalmoscopy CSLO allows real-time, three-dimensional imaging of the retinal nerve fiber layer (RNFL) and optic nerve head (ONH). The confocal scanning system is based on the principle of spot illumination and spot detection. It is designed to allow only a ‘thin’ slice or spot of the retina to be in focus on the image plane. Light rays reflected from higher or lower focal planes are blocked, thus creating highresolution tomographic images. Retinal tissue is illuminated and imaged point by point through a pinhole. The illumination pinhole and the imaging pinhole correspond to the same focal point on the tissue, thereby making the system confocal. A three-dimensional image may be acquired by adjusting the x, y, and z coordinates.
Address for correspondence: Joel S. Schuman, MD, New England Eye Center, Tufts-New England Medical Center, Tufts University School of Medicine, 750 Washington Street, Box 450, Boston, MA 02111, USA. e-mail:
[email protected] Lasers in Ophthalmology – Basic, Diagnostic and Surgical Aspects, pp. 143–151 edited by F. Fankhauser and S. Kwasniewska © 2003 Kugler Publications, The Hague, The Netherlands
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Fig. 1. CSLO printed report: HRT I, subject with glaucoma. A. The topographic (left) and reflectivity image (right) illustrates the ONH. In the contour graph (below), the white line represents the reference plane at which there is a height of zero. The red line represents the height of the reference line between the cup and disc. The green line is the retinal height of the subject eye at the contour line showing the typical double hump feature at the superior and inferior poles. B. The topographic image is shown with the cup represented in red, the sloping neural tissue in blue and the rim in green. The ONH parameters and subject classification are shown on the right. The classification number for the HRT I is determined by an automated algorithm devised by Mikelberg, based on the ONH and retinal parameters.7
patients and 51 patients with early glaucoma, using qualitative assessment of stereoscopic optic disc photographs and CSLO imaging.4 In another comparative study, clinicians analyzed 13 ocular hypertensive eyes that subsequently developed reproducible visual field defects and 13 normal eyes that had undergone sequential optic disc images. HRT
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was found to detect glaucomatous changes in the ONH before visual field changes occurred.5 The recently-developed HRT II is also designed for topographic ONH analysis; however, the device is small, lightweight, portable, and almost completely automatic. All parameters for image acquisition are fixed or predetermined. The HRT II au-
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Fig. 2. CSLO printed report: HRT II, different subject with glaucoma. The topographic image (left) is shown with the cup represented in red, the sloping neural tissue in blue and the rim in green. The reflectivity image (right) illustrates the classification of the six ONH sector. Each sector is marked with a green check mark, a red cross, or a yellow exclamation mark to illustrate being within normal limits, outside normal limits, and borderline, respectively. A bar graph represents this further in the right middle panel. The stereometric parameters are displayed in the left middle panel. The classification number for HRT II is derived from an algorithm developed by Wollstein et al. at Moorfields Eye Hospital. (Reproduced by courtesy of Heidelberg Engineering, Inc., Carlsbad, CA.)
tomatically acquires 16-64 image planes covering a field of view fixed at 15 × 15° using 384 × 384 pixels per plane. Utilizing an internal fixation target, this system routinely acquires three images with the use of a quality control system that will obtain additional images if one or more of the images cannot be used (e.g., fixation loss). The printed report illustrates a topographic and reflectivity image of the ONH and its contour line, details of the ONH stereometric parameters and classification (Fig. 2). The HRT I is a research-oriented tool with a wide range of applications. It can be utilized to measure retinal circulation when combined with the Heidel-
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berg Retina Flowmeter. The HRT II, however, is restricted to ONH analysis. An inherent limitation of the HRT technology lies in the reference plane that is required to calculate cup area, rim area, rim volume, cup volume, cup-to-disc ratio, retinal nerve fiber layer thickness, and retinal nerve fiber layer cross-sectional area. The anatomical reference plane used by the current software may change over time, especially in patients with glaucoma who have changing topography.3 Another obstacle is the manual delineation of the optic disc margin by the operator, and the influence of this on ONH parameters. Published data regarding analysis of ONH pa-
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rameters using CSLO show that the slope of the cup (‘the third central moment of the depth distribution’) may be the most significant parameter in the prediction of glaucoma status (Fig. 2).6,7 CSLO appears to have an ability to discriminate between normal and glaucomatous eyes with a sensitivity and specificity of about 85%. However, considerable overlap exists between normal, ocular hypertensive, and glaucomatous eyes.7,8 Authors have claimed the ability to determine RNFL thickness or cross-sectional area using CSLO, by using a reference point in the nasal retina or the macula.9 This indirect method of measuring RNFL thickness is almost certainly not the best technique for RNFL analysis, nor is it particularly accurate given the low axial resolution of CSLO (approximately 300 µm) and that superior technologies exist for the direct measurement of RNFL thickness.10 Chauhan and colleagues recently developed software to detect topographic changes in the optic disc and peripapillary retina that appears to provide the best longitudinal data analysis to date using HRT II. Scanning laser tomography and conventional perimetry were used to follow 77 subjects with early glaucomatous visual field damage. Glaucomatous disc changes revealed with scanning laser tomography were found to occur more frequently than visual field changes. This result implies that glaucomatous damage and progression can be detected earlier using scanning laser polarimetry.11,12 Optical coherence tomography OCT is an optical diagnostic technology that permits high-resolution cross-sectional imaging of the human retina using low-coherence light that is backscattered by the sample boundaries.13 By performing multiple longitudinal scans at different transverse locations, a two-dimensional scanned image is obtained. Compared with the confocal scanning laser ophthalmoscope, which has a transverse resolution of about 20 µm and an axial resolution of the order of 300 µm, the first commercial OCT (OCT 1) produced by Zeiss-Humphrey has a significantly improved resolution: transverse of 20 µm and axial of about 10 µm in the human eye. Non-contact, non-invasive human eye imaging using OCT proved to advance ophthalmic diagnostics.14 Recently, a third generation commercial OCT (OCT 3) was introduced, which provides high-resolution images with 8-µm resolution and 512 A scans per image, acquired in about one second. A comparison between a scanned image obtained with a commercial 10-µm resolution OCT and one obtained with a pre-production 8-µm resolution OCT of a normal eye is shown in Figure 3. Research studies on ultrahigh resolution OCT (UHR OCT) demonstrate an axial resolution of 3 µm, using a broad bandwidth femtosecond laser light source.15 UHR OCT has image quality mak-
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Fig. 3. OCT 1 and OCT 3 images for comparison. The images are obtained with circular scans around the ONH and depict the RNFL of a normal eye. Both images were acquired in the right eye of the patient on the same day. The 8-µm resolution image obtained with the OCT 3 (B) reveals enhanced retina layer demarcation compared to OCT 1 (A). Layers such as the outer plexiform, outer nuclear, and choriocapillaris, and choroid are discernable in the OCT 3 image. These layers cannot be discriminated in the OCT 1 image.
ing possible identification and quantification of the intermediate layers of the retina, such as the ganglion cell, inner plexiform, inner nuclear, outer plexiform, and outer nuclear layers, as well as enabling differentiation of the choriocapillaris and choroid from the retinal pigment epithelium (RPE); these discriminations were not possible with any previous device. Recent UHR OCT images of the retina compared well with histology, as shown in Figure 4. UHR OCT images of the macula to optic nerve area in a normal eye showed a significant improvement compared to images obtained with OCT 1 or 2 and OCT 3, as seen in Figure 5. OCT has been used to detect glaucoma, retinal diseases such as macular holes and pseudoholes, non-proliferative and proliferative diabetic retinopathy, macular degeneration, retinal and RPE detachment, chorioretinal inflammatory disease, retinal dystrophies, retinal trauma, and diseases of the optic nerve such as optic disc pitting, optic disc swelling.16 Retinal imaging In the retina, OCT provides a cross-sectional image of optical reflectivity. Retinal thickness is automatically calculated from the anterior and posterior boundaries of the retina. Retinal thickness increases with edema, which can be located in the macula. This is an important and common process in diabetic retinopathy, retinal vein occlusion, and uveitis. A decrease in retinal thickness is associated with atrophy or scarring, as well as with glaucoma. Neurosensory detachments of the retina or RPE, macular lesions such as holes, and fibrosis can be
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Fig. 4. UHR OCT macular scan and macular histology. Comparison of an in vivo UHR OCT image (top) of a normal human macula and normal macular histology (bottom, reproduced by courtesy of Drexler et al.15). Several layers are seen and labelled in the image: inner limiting membrane (ILM), retinal nerve fiber layer (RNFL), ganglion cell layer (GCL), inner plexiform layer (IPL), inner nuclear layer (INL), outer nuclear layer (ONL), external limiting membrane (ELM), retinal pigment epithelium (RPE). The foveola, fovea centralis, as well as the parafoveal region, are also indicated. The UHR OCT image was obtained with a linear scan through the macula.
Fig. 5. OCT 1, OCT 3, and UHR OCT images for comparison. The images depict a linear scan from the macula to the optic nerve in a normal eye. They were acquired with standard OCT 1 with ~10 µm (top), OCT 3 with ~8 µm (middle), and UHR OCT with ~3 µm (bottom) axial image resolution, respectively. UHR OCT significantly improves the ability to distinguish retinal features.
detected in an OCT image. Recent improvements in resolution facilitate distinguishing RPE and choriocapillaris, providing useful information on agerelated macular degeneration and choroidal neovascularization. OCT can be useful in the diagnosis and monitoring of macular holes, macular edema, and retinal detachment.17 In eyes with epiretinal membranes, OCT can provide structural assessment of the macula pre- and postoperatively.18 OCT may be useful
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for screening and monitoring patients with diabetic retinopathy.19 Enhancement of the quality of the images obtained with the 8-µm resolution afforded by OCT 3 make clear differentiation of the affected retinal layers and cyst-like spaces in cystoid macular edema possible, as seen in Figure 6. OCT calculates both retinal and RNFL thickness. A significant correlation exists between macular and RNFL thickness in glaucoma.20 Using the 8-
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Fig. 6. OCT 3: normal macula and macular edema. A. OCT 3 image of a normal macula. B. OCT 3 macular image demonstrating cystoid macular edema. Images were taken with linear scans through the macula. Image A shows the right eye, image B the left eye of the same patient. Note the ability to discriminate the affected retinal layers and cyst-like spaces in the eye with cystoid macular edema.
Fig. 7. OCT 3: macular thickness and macular map. OCT 3 image showing macular thickness and a macular thickness map calculated from the six radial foveally-centered scans through a normal macula. The image depicts a printed report: macular thickness delimited by the white lines (left), fundus view (right), and macular thickness map (bottom).
µm resolution OCT 3, retinal disease imaging and diagnosis may be enhanced; UHR OCT may provide further improvements in reproducibility, sensitivity, and specificity of measurements, due specifically to the higher resolution possible with UHR OCT. Figure 7 demonstrates a case in which a clear demarcation between retinal layers is shown in a normal eye using OCT 3. Optic nerve head imaging Alterations in the thickness of the RNFL and the ONH contour are important in the diagnosis and follow-up of glaucoma; these parameters are also
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of interest in papilledema and papillitis. OCT calculation of ONH parameters gives useful quantitative information about the degree of ONH injury. There is a strong correlation between OCT measurements of RNFL thickness in glaucomatous eyes and stereoscopic ONH photography, ophthalmic examination, and functional status, as measured by Humphrey 24-2 visual fields.21-23 OCT measures demonstrated thinning of the RNFL with age, RNFL thinning in glaucoma corresponding to visual field defects, and RNFL focal defects and cupping.22 A study in normal eyes, glaucoma suspects, and early and advanced glaucoma patients demonstrated
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Fig. 8. OCT 3: normal ONH. Normal ONH OCT 3 image and calculation of ONH parameters such as cup, disc, and rim areas, cup/ disc ratio. The image is obtained using six radial scans centered on the ONH. Note the red-shaded areas in the left image representing the rim areas. The image on the right depicts the disc and cup contours and a single radial ONH scan is shown on the left.
Fig. 9. OCT 3: glaucomatous ONH. Glaucomatous ONH OCT 3 image and calculation of ONH parameters such as cup, disc, and rim areas, cup/disc ratio. The image is obtained using six radial scans centered on the ONH. Note the red-shaded areas in the left image representing the rim areas. The image on the right depicts the disc and cup contours and a single radial ONH scan is shown on the left. Note the large cup attributed to glaucoma.
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that RNFL thickness calculated from the OCT image was the best parameter for discrimination, compared to confocal scanning laser ophthalmoscopy24 or to CSLO, scanning laser polarimetry, short wavelength automated perimetry, or frequency doubling perimetry.22 Improvement in resolution offered by OCT 3 may enhance OCT ONH imaging (Fig. 8), and may allow more accurate quantification of parameters such as rim area, cup-to-disc ratio, and volumetric measurements (Fig. 9). Summary The goal of imaging in ophthalmology is to enable the accurate and early diagnosis or tracking of eye disease. The HRT is used to assess the contour of the ONH and posterior pole. OCT imaging provides high (or ultrahigh) resolution images of the retina, permitting reproducible cross-sectional measurements of retinal and RNFL thickness and ONH topography. CSLO and OCT can contribute to earlier and more accurate diagnosis in ophthalmic disease, and may serve to measure change over time. These technologies approach the problem of structural ophthalmic imaging in different ways, and each provides reproducible quantitative information. OCT has the added benefit of utility in multiple areas of ophthalmic imaging; UHR OCT is analogous to in vivo histology. Conclusions In conclusion, imaging in ophthalmology is an important tool for the diagnosis and tracking of disease. Imaging techniques such as Heidelberg Retina Tomography and Optical Coherence Tomography provide high (or ultrahigh) resolution images that assess the structure and substructure of the retina. These imaging techniques allow quantitative assessment of retinal tissue and are expected to detect disease and its progression earlier than could be done in any other way. References 1. Schuman JS, Kim J: Imaging of the optic nerve head and nerve fiber layer in glaucoma. Ophthalmol Clin N Am 8(2):259-279, 1995 2. Echelman DA, Shields MB: Optic nerve imaging. In: Albert DM, Jakobiec FA (eds) Principles and Practice in Ophthalmology, Vol 3, pp 1310-1329. Philadelphia, PA: WB Saunders 1994 3. Mikelberg F, Wijsman K, Schulzer M: Reproducibility of topographic parameters obtained with the Heidelberg Retina Tomograph. J Glaucoma 2:101-103, 1993 4. Zangwill L, Shakiba S, Caprioli J, Weinreb RN: Agreement between clinicians and a confocal scanning laser ophthalmoscope in estimating cup/disk ratios. Am J Ophthalmol 119:415-421, 1994 5. Weinreb RN, Luski M, Bartsch DU, Morsman D: Effect of
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repetitive imaging on topographic measurements of the optic nerve head. Arch Ophthalmol 111:636-638, 1993 6. Brigatti L, Caprioli J. Correlation of visual field with scanning confocal laser optic disc measurements in glaucoma. Arch Ophthalmol 113(9):1191-1194, 1995 7. Mikelberg FS, Prafitt CM, Swindale NV et al: Ability of the Heidelberg Retina Tomograph to detect early glaucomatous visual field loss. J Glaucoma 4:242-247, 1995 8. Zangwill L, Horn SV, Lima MDS, Sample PA, Weinreb RN: Optic nerve head topography in ocular hypertensive eyes using confocal scanning laser ophthalmoscopy. Am J Ophthalmol 122:520-525, 1996 9. Weinreb RN, Shakiba S, Sample PA et al: Association between quantitative nerve fiber layer measurement and visual field loss in glaucoma. Am J Ophthalmol 120:732738, 1995 10. Schuman JS, Noecker RJ: Imaging of the optic nerve head and nerve fiber layer in glaucoma. Ophthalmol Clin N Am 8(2):259-279, 1995 11. Chauhan BC, Blanchard JW, Hamilton DC, LeBlanc RP: Technique for detecting serial topographic changes in the optic disc and peripapillary retina using scanning laser tomography. Invest Ophthalmol Vis Sci 41(3):775-782, 2000 12. Chauhan BC, McCormick TA, Nicolela MT, LeBlanc RP: Optic disc and visual field changes in a prospective longitudinal study of patients with glaucoma: comparison of scanning laser tomography with conventional perimetry and optic disc photography. Arch Ophthalmol 119(10):14921499, 2001 13. Huang D, Swanson EA, Lin CP, Schuman JS, Stinson WG, Chang W, Hee MR, Flotte T, Gregory K, Puliafito CA, Fujimoto JG: Optical coherence tomography. Science 254(5035):1178-1181, 1991 14. Swanson E, Izatt, Hee M et al: In vivo retinal imaging by optical coherence tomography. Optics Lett 18:1864-1866, 1993 15. Drexler W, Morgner U, Ghanta RK, Kartner FX, Schuman JS, Fujimoto JG: Ultrahigh-resolution ophthalmic optical coherence tomography. Nature Med 7(4):502-507, 2001 16. Puliafito CA, Hee MR, Lin CP, Reichel E, Schuman JS, Duker JS, Izatt JA, Swanson EA, Fujimoto JG: Imaging of macular diseases with optical coherence tomography. Ophthalmology 102(2):218-229, 1995 17. Wilkins JR, Puliafito CA, Hee MR, Duker JS, Reichel E, Coker JG, Schuman JS, Swanson EA, Fujimoto JS: Characterization of epiretinal membranes using optical coherence tomography. Ophthalmology 103(12):2142-2151, 1996 18. Hee MR, Puliafito CA, Duker JS, Reichel E, Coker JG, Wilkins JR, Schuman JS, Swanson EA, Fujimoto JG: Ophthalmology 105(2):360-369, 1998 19. Guedes V, Schuman JS, Hertzmark E, Correnti A, Mancini R, Wollstein G, Lederer D, Voskanian S, Velazquez L, Parker HM, Pedut-Kloizman T, Fujimoto JG, Matox C: Optical coherence tomography measurement of macular and nerve fiber layer thickness in normal and glaucomatous human eye. Ophthalmology (in press) 20. Schuman JS, Hee MR, Puliafito CA, Wong C, PedutKloizman T, Lin CP, Hertzmark E, Izatt JA, Swanson EA, Fujimoto JG: Quantification of nerve fiber layer thickness in normal and glaucomatous eyes using optical coherence tomography: a pilot study. Arch Ophthalmol 113:586-596, 1995 21. Pieroth L, Schuman JS, Hertzmark E, Hee MR, Wilkins JR, Coker J, Mattox C, Pedut-Kloizman T, Puliafito CA, Fujimoto JG, Swanson E: Evaluation of focal defects of the nerve fiber layer using optical coherence tomography. Ophthalmology 106(3):570-579, 1999
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Imaging in ophthalmology 22. Bowd C, Zangwill LM, Berry CC, Blumenthal EZ, Vasile C, Sanchez-Galeana C, Bosworth CF, Sample PA, Weinreb RN: Detecting early glaucoma by assessment of retinal nerve fiber layer thickness and visual function. Invest Ophthalmol Vis Sci 42(9):1993-2003, 2001 23. Soliman MA, Van Den Berg TJ, Ismaeil AA, De Jong LA, De Smet MD: Retinal nerve fiber layer analysis: relationship between optical coherence tomography and red-free photography. Am J Ophthalmol 133(2):187-195, 2002 24. Parker HM, Schuman JS, Hertzmark E, Pedut-Kloizman
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151 Tpeiris ID, MacNutt J, Miller VM, So S, Ghanta RK, Drexler W, Fujimoto JG, Puliafito CA, Mattox C, Rasheed ES, Guedes VRF: Optical coherence tomography of the retinal nerve fiber layer, with comparison to Heidelberg retina tomography optic nerve head measurements, in normal and glaucomatous human eyes. In: Lemij HG, Schuman JS (eds) The Shape of Glaucoma, Quantitative Neural Imaging Techniques, pp 149-181. The Hague: Kugler Publications, 2000
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Different methods of refractive surgery The advantages and risks, and their relationship to professional ethics and morals
Björn M. Tengroth Stockholm, Sweden
Keywords: refractive errors, aberration, myopia, hyperopia, astigmatism, presbyopia, cornea, RK, PRK, LASIK, INTAC, LTK, LASEK, ELSA, IOL, ethics, morals, risks
In earlier times, having a refractive error in one’s vision was tantamount to having a life-threatening handicap. In prehistoric days, when hunting tribes were dominant, a nearsighted person was at an enormous disadvantage. His possibilities of obtaining food and defending himself were small. By necessity, he became dependent on associated with better vision. In a more civilized world with an increased differentiation in people’s daily lives, even the individual with a refractive error could find an occupation, doing primarily different types of craft work. Obviously, presbyopia was a handicap for older people. Close, detailed work was impossible to deal with. Reading and writing in poor lighting could be entrusted to nearsighted people. Various aids such as reading stones (magnifying glasses) were used, but these were of such poor quality that it was still impossible to carry out detailed close work. It was not until the twelfth century that eyeglasses appeared, but they were excessively expensive, and few could afford them. Giving nearsighted people qualified jobs involving fine details in writing, reading, and arts and crafts was common practice. As optical aids became more advanced, particularly during the twentieth century, with more accurately adapted glasses and contact lenses, individuals with refractive errors began to have the same opportunities for most kinds of work as those with normal vision. There are, however, still some jobs where eyeglasses and contact lenses are a disadvantage. Freedom from the inconvenience of using optical aids has long been a desire of those who nevertheless need them. When contact lenses began to be used, they were hailed by many as liberating, even though users’ vision did not always become optimal. The risks involved in using them were not
negligible either. Complicated infections and vascular ingrowth in the cornea are not unusual. Soft (hydrophillic) lenses have now been developed for a short period of use, and 24-hour lenses are currently the most popular kind. These has reduced the risks dramatically, but have not eliminated them. However, it has been determined that standard eyeglasses free the user from the risks of both injury and infection. An otherwise healthy person with nearsightedness (myopia), farsightedness (hyperopia), or astigmatism can be regarded as having a normal variation. Being able surgically to normalize a person’s refraction has long been desired, but it was not until the end of the twentieth century that such an intervention was seriously discussed. The risks involved often presented an obstacle to the development of refractive surgery. When Sato’s idea1 was taken up by Fyodorov2 in the Soviet Union, in what came to be called radial keratotomy (RK), the risks of injury and infection were regarded as relatively minor. Experience has been gained from several decades of using this method, during which millions of patients have undergone RK surgery. Even though most of those with low-grade myopia have had an acceptable visual improvement, there have been a few unfortunate cases in which perforations occurred, and postoperative infections followed. There have also been reported of hyperopia.3 It is also know that people who have had this operation are at higher risk of injury from blows to the eye. Considering all the disadvantages now known, it is understandable that RK surgery has been abandoned in most places. Replacing RK with surgery involving less risk and more precision led to the development of the excimer laser for this purpose. At first, attempts were made to replace the diamond knives then being used to
Address for correspondence: Professor Björn M. Tengroth, City Varden, Nya Ogon, Apelbergsgatan 48, SE 11137 Stockholm, Sweden. Lasers in Ophthalmology – Basic, Diagnostic and Surgical Aspects, pp. 153–157 edited by F. Fankhauser and S. Kwasniewska © 2003 Kugler Publications, The Hague, The Netherlands
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make the radial incisions with the excimer laser, in order to eliminate perforations. However, the incisions were too broad. The idea of improving the technique of using radial incisions to flatten the cornea, and being able to ‘plane away’ a portion of the corneal stroma over a larger area appeared at the beginning of the 1980s.4-8 However, this method involved working in the optical zone, which many ophthalmologists considered hazardous. Several companies in Germany and the USA developed excimer lasers for this purpose. They treated surfaces of between 3 and 4 mm in diameter with excimer pulses that removed between 0.2 and 0.3 µm per pulse. Generally, treatment was done to a depth of less than 100 µm. The larger the diameter of the treated area, the deeper you had to go into the stroma. This technique, photorefractive keratectomy (PRK) quickly became popular in Europe, where there were no major restrictions against carrying out a procedure of this nature. In the USA, the procedure was not accepted until after seeing the results of a multicenter study with a few cases in each study, and this was not until November 1995.8 In Europe, and especially in Sweden, thousands of people underwent PRK surgery in the early 1990s.9 The advantages of this method were that there was no risk of perforations, of reduction of stability in the cornea, or of infection. The disadvantages were that surgery was still being performed in the optical zone, and that there was postoperative pain and a slow healing process in ten percent of the cases. There could also be corneal haze in the postoperative period, which could nevertheless be regarded as a temporary scarring that always disappeared with time. Some regression also occurred in a number of eyes, and the precision was not perfect. Even though this method involved less risk when compared to RK, and produced good results (20/40 UCVA in 95% of cases, and stability after about three months), some patients had to undergo re-operation. Most patients had problems with night vision (halo phenomenon) and reduced sensitivity to contrast, primarily during the first year after surgery. The halo phenomenon was caused by the small diameter of the operated area, often smaller than the diameter of the pupil in weak light. By improving the technique with larger operation zones (>7 mm), and by using different treatments with antiphlogistic medications such as steroids, some of the postoperative complications could be reduced.9 In the USA, PRK was not permitted at first. Before the appearance of the excimer laser, a number of refractive procedures had been done on corneas, primarily in South America. Epikeratophakia and keratomileusis has been performed in Bogota.10 The latter technique could be modified so that, instead of implanting small discs made from donor corneas, the excimer laser was used.11,12 This was how laserassisted in situ keratomileusis (LASIK) was developed. LASIK became very popular, especially in the USA.
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One procedure which became popular is the one in which an intraocular lens is implanted after extraction of the clear lens. However, this method has a very long history.13,14 Another method with a more modern background, which became equally popular, involves implantation of anterior chamber lenses in front of the regular lens.15,16 In LASIK, a flap of the cornea with a diameter of about 9 mm and a thickness of about 150 µm is cut from the corneal surface. Excimer laser treatment removes tissue under this flap, which is normally fastened by a hinge at the 12 o’clock position in the corneal incision. After the operation, the flap is replaced, and a bandage lens is placed over it. The advantages of LASIK surgery include a painless postoperative period, rapid healing, and functional vision almost immediately after the procedure. Disadvantages are that, in a very few cases, the incision can go askew or be made too deeply, which can cause permanent injury. By cutting through the epithelium and into the cornea, it can happen that there is epithelial growth in under the flap. This must be removed by re-lifting the flap. This is not a problem, since the flap never grows together with the base. In the same manner, folds can appear in the flap, and re-operation is necessitated. When the results are not satisfactory, the flap must be lifted and a new treatment performed. Since the flap never reattaches, i.e., heals, an injury to the eye with a sharp object can cause the entire flap to detach. This is, of course, an injury that can be difficult to repair. When the flap is cut, a considerable reduction in the cornea’s stability is unavoidable. Keratectasia has been reported in a fair number of these cases.17 An increased hyperopia with irregular astigmatism is the consequence. The LASIK technique’s popularity has decreased somewhat in recent years, to be replaced by the ELSA technique described below. Experiments have been made with the implantation of flexible plastic bands into the corneal stroma.18,19 A plastic band is inserted into the corneal stroma about 4 mm from the center of the cornea, and by tightening this band, a certain amount of alteration can be made in the shape of the cornea. The procedure is not performed near the optical zone, and the band can be removed at any time. However, the precision of this procedure is not good, and, at present, it does not appear that this method has much of a future. Some infection has also been reported. Since cataract surgery is performed frequently, eye surgeons from all over the world have extensive experience in operating on cataracts by implanting an intraocular lens. Removing a completely clear lens and replacing it with an appropriate intraocular lens has become a popular technique. Since the risk of infection is very small, although still present, it has been recommended that this procedure be reserved for patients with higher degrees of refractive error. Another method involving the placement of an intraocular lens in front of the existing lens has been used successfully in several countries. There is a risk
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Different methods of refractive surgery of very serious infection here, as well as the likelihood of the patient’s own lens rubbing against the implanted lens, which can lead to cataract formation over time. The precision of the two last-mentioned procedures is not as advanced as that of LASIK. Laser thermokeratoplasty (LTK) is a technique in which an attempt is made to increase the curvature of the cornea. It has been used in cases of farsightedness and presbyopia. In this procedure, a special laser beam is used to heat up a spot deep down in the corneal stroma.20 Repeated burns are made around the entire circumference, sometimes in one circle and sometimes in several concentric circles. This method is completely painless, and results in an immediate correction. However, after three to six months, regression sets in and within 12 to 18 months, the refraction returns to its preoperative condition. This technique is still used in some places, but its use can hardly be defended when there are better techniques currently in existence. In recent years, a new method has been developed and used. This is called excimer laser subepithelial ablation (ELSA) or laser subepithelial keratectomy (LASEK).21,22 The latter name should be avoided, as it can easily be confused with LASIK. In this procedure, the epithelial layer is completely removed, but about 90° of the circumference is allowed to remain as a short of hinge. After the laser treatment, which is equivalent to PRK, the epithelium is replaced. In order to loosen the epithelium effectively, a solution of 20% alcohol is applied to the cornea for 25-35 seconds before the dissection of the epithelial flap is made. The results of this procedure are very good, and show ELSA’s advantages without any of its disadvantages. In general, the patient experiences very little postoperative pain, and healing usually takes place rapidly, so that the patient has useable vision within a day or so. There is no haze, and no long-term healing process. Since the technique is more or less the same as PRK, there is great precision. ELSA can in principle be regarded as riskfree in comparison with all the other methods. We have 15 years of experience with PRK, and the two methods can be seen as so similar that we expect no long-term reactions. There are several excimer laser apparatuses available at present. In principle, there are two different methods. The first involves placing various kinds of apertures in the excimer laser’s beam, and with the second, a small spot of 1-2 mm in diameter is programmed to move in a certain pattern over the area to be treated. The laser’s computer is programmed to treat the specific refraction. There has been no evidence of any long-term differences in results. There are several kinds of lasers on the market, which operate on somewhat different principles. No evaluation has yet been made, and no-one has been able to show any clear advantages with any one type of laser. This is a serious drawback. All the articles that tout the advantages of one method over another either contain statistics that are not well-qualified,
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155 or are written by authors who are associated with the companies that procedure the new instruments. A number of the additions mentioned above have not been adequately documented either. There is a danger that both patients and doctors will find distorted and biased information with Internet web searches. Many doctors use the various new instruments and auxiliary equipment for advertising purposes, which is hardly beneficial to refractive surgery in general. A lens system is always impaired by certain optical errors, such as spherical and chromatic aberrations, as well as coma.23,24 With the aid of a device called an aberrometer, these errors can be calculated. Aberrations located both in the cornea and in the eye’s lens vary with the size of the pupil and the accommodation status of the lens, and also with the composition of the light. In calculation, this is always done in the new technology with a dilated pupil and with accommodation paralyzed. A monochromatic light is also used. Thus, these calculations give information about a very special situation that never occurs in everyday life. The numbers obtained are only to be used in the laser algorithms, and the value of this can be discussed. So far, no satisfactory investigations have been done that allow us to say that these calculations are sufficiently significant to be used in treatment. Some of the lasers can be partially programmed with a topographical image of the cornea, and this ought to be better results. No scientific report has yet been published that shows the advantages. The topography of the cornea can change postoperatively in such a way that an irregular astigmatism appears. In a case like this, a treatment controlled by a topographical image would have had a great advantage. No such construction has yet appeared on the horizon; this is probably due to the laser algorithm’s inability to handle a more complicated topographic image. The use of refractive surgery has now spread over most of the industrial world. Millions of people have had operations. The development and manufacture of instruments for refractive surgery have become a major industry; the investment for each clinic is about $500,000 to $1,000,000. At conventions and meetings, as well as in professional publications, the subject of refractive treatment claims more and more space. The optical branch is starting to feel threatening, especially the contact lens market. This is the first time that ophthalmologists have entered an area with clearly mercantile objectives – an area in which the focus is not on diseases, but rather on variations within the normal range of people’s vision errors. We have started a process that digresses from the ethical rules that have guided us since the Hippocratic oath was instituted. We doctors are here to help without injuring. Are we moving in the right direction? Seen from our patients’ viewpoint, refractive error is a troublesome handicap. They come to us for help. In today’s world,
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most of our patients are very well informed, especially by the Internet. Our homepages are a form of advertisement, but are they objective? Are we duping our patients into believing that we have the right solutions to their problems? There are huge economic interests at stake here, especially for the eye surgeon. These are normal variations that we are treating, and not diseases! It should be the duty of every eye doctor to weight the risks in relation to the refractive surgery we are going, without casting sideways glances at the financial gain. Of course, our patients can see the many advantages of a successful treatment. The operation is also financially rewarding for the patient in the long run; the operation’s cost is written off in a few years. But, have we given them the proper information about the risks? Several of the methods described above do indeed carry significant risks. Obtaining vision that is precisely as good as the patient has with eyeglasses is actually rare in cases of mild myopia. Night vision and contrast sensitivity are impaired, and aberrations increase, even though visual acuity is good.25 In cases in which there is a greater refraction error, visual acuity usually improves by one or two lines on the eye chart, but the errors mentioned above remain. There are some individuals whose jobs or other situations prevent them from wearing glasses or contact lenses, and it is of course suitable to operate on these individuals – but what about the others? The patient’s motivation is the most important factor in carrying out a treatment, but his motivation should not be influenced by either erroneous information or withheld information. If he is satisfied with his contact lenses or glasses, he would do well to avoid refractive surgery as it is being handled today! The future may bring us even better and safer methods that may change these conditions, but until then, the ophthalmologist’s approach to refractive surgery should be characterized by a certain amount of restrictiveness.
Conclusions Refractive surgery has come to stay. The methods will improve and so also the lasers and other equipment. However we have to be observant of side effects, as infections especially in intraocular surgery, unexpected regression and keratoectasia. So far we have a relatively long experience from PRK which of course can be applied even for LASEK/ ELSA. The long term results from LASIK (>10 years) we do not know but the risk for keratoectasia should not be overlooked as more and more cases are reported. No method can be totally safe when compared to spectacles. The doctor should be aware of the fact that the eyes treated are normal and no risks should be taken that can result in loss of vision. On the other
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hand all our patients will benefit if they can get rid of spectacles and contact lenses in their work and daily life but they have to be totally informed and well motivated prior to any kind of treatment. References 1. Sato T, Akiyama K, Shibata H: A new surgical approach to myopia. Am J Ophthalmol 36:541, 1953 2. Fyodorov SN, Durnev VV: Operation of dosaged dissection of corneal circular ligament in cases of myopia of mild degree. Ann Ophthalmol 11:1185, 1979 3. Waring GO, Lynn MJ, McDonell PJ: The PERK study group. Results of the prospective evaluation of radial keratotomy (PERK) Study ten years after surgery. Arch Ophthalmol 112:1298, 1994 4. Trokel SL, Srinivasan R, Braren R: Excimer laser surgery of the cornea. Am J Ophthalmol 96:710, 1983 5. Munnerlyn C, Kroons SJ, Marshall J: Photorefractive keratectomy: a technique for laser refractive surgery. J Cataract Refractive Surg 14:46, 1988 6. Seiler T, Bende T, Wollensak J, Trokel S: Excimer laser keratectomy for correction of astigmatism. Am J Ophthalmol 105:117, 1988 7. Trokel SL, Munnerlyn C: Excimer laser delivery systems. Laser and Light in Ophthalmology 2:157, 1989 8. McDonald MB, Kaufmann HE, Frantz JM, Shofner S, Salmeron B, Klyce D: Excimer laser ablation in a human eye. Arch Ophthalmol 107:641, 1989 9. Epstein D, Fagerholm P, Hamberg-Nyström H, Tengroth B: Twenty-four months follow-up of excimer laser photorefractive keratectomy for myopia. Ophthalmology 101:1558, 1994 10. Barraquer JL: Keratomileusis for the correction of myopia. Arch Soc Am Oftal Optom 5:27, 1964 11. Pallikaris IG, Papazanaki NE, Stathi EZ, Frenschok O, Georgiadis A: Laser in situ keratomileusis. Lasers Surg Med 10:462, 1990 12. Burrato L, Ferrari M, Rama P: Excimer laser intrastromal keratomileusis. Am J Ophthalmol 113:291, 1992 13. Boerhaave H: Praelectiones publicae de morbus oculorium, Göttingen. A. Vandenhoeck 1746 14. Janin CB: Principles and Practice of Refractive Surgery. In: Elander RR (ed), p 3. WB Saunders Co 1997 15. Barraquer J: Anterior chamber plastic lenses; results of and conclusions from five years experience. Trans Ophthalmol Soc UK 79:393, 1959 16. Fechner PU, Haigis W, Wichmann W: Posterior chamber myopia lenses in phacic eyes. J Cataract Refract Surg 22:178, 1996 17. Seiler T, Kaemmerer M, Mierdel P, Krinke H-E: Ocular optical aberrations after photorefractive keratectomy for myopia and myopic astigmatism. Arch Ophthalmol 118:1721, 2000 18. Fleming JF, Wan LW, Schanzlin DJ: The theory of corneal curvature change with the intrastromal corneal ring. CLAO 15:146, 1989 19. Nose W, Neves RA, Burris TE, Schanzlin DJ, Belfort J: The intrastromal corneal ring: Twelve months sighted myopic eyes. J Refractive Surg 266:229-230, 1996 20. Vassiliadis A: Personal com. 21. Camellin M: Subepithelial corneal ablation. Ocular Surgery News International Edition 3:14, 1999 22. Lohman CP, Winkler van Mohrenfels C, Gabler B, Herrmann W, Müller M: Excimer Laser Subepitheliale Ablation
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Different methods of refractive surgery (ELSA) bzw. Laser epitheliale keratomileusis (LASEK) – ein neuartiges refraktiv-chirurgisches Verfahren zur Myopiekorrektur. Operationstechnik und erste Ergebnisse an 24 Augen und nach 3 Monaten. Klin Monatsbl Augenheilk 219:26-32, 2002
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157 23. Von Helmholtz H, Gullstrand A: Handbuch der physiologischen Optik, p 132 (1867) and p 353 (1909). Leopold Voss 1867 and 1909 24. Mrochen M, Bueeler M, Seiler T: Corneal Laser Surgery for Refractive Correction. 2002 (in print)
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Corneal laser surgery for refractive corrections Michael Mrochen, Michael Bueeler and Theo Seiler University of Zurich, Department of Ophthalmology, Zurich, Switzerland
Keywords: photorefractive keratectomy, LASIK, optical aberrations, success rate, complications
Abstract
Optical aberrations of the human eye
Corneal laser surgery with the modern excimer laser is known to be the most frequently applied laser procedure in medicine. World-wide, more than 2.5 million procedures were performed in 2001. Surgical techniques such as photorefractive keratectomy (PRK) and laser-assisted in-situ keratomileusis (LASIK) are used to correct optical imperfections of the eye, such as short (myopia) and long (hyperopia) sightedness, as well as astigmatism. However, the use of corneal laser surgery is also associated with disadvantages and complications. In particular, the increase of higher-order optical imaging errors (wavefront errors) after laser treatment leads to visual discomfort in poor illumination conditions, e.g., driving a car at night. This chapter provides an overview of the basic principles of corneal laser surgery, side-effects, complications, and future developments.
The term ‘monochromatic optical aberrations’ was coined in the early 19th century and, at that time, included all optical errors of the eye, except for spherical myopic or hyperopic refractive errors.2 During that time, Hermann von Helmholtz described the optics of the eye to be imperfect, and questioned the quality of the retinal image projected by cornea and lens.3 At the end of the 19th century, Tscherning investigated the non-refractive errors of the eye in more detail.4 However, Von Helmholtz described these higher-order errors as being of minor importance,3 in contrast to Tscherning4 and Gullstrand.5 Donders6 defined and specified the measurement and correction of ocular astigmatism, introducing the sphero-cylindrical error and, thus, reducing the ‘monochromatic optical aberrations’ to higher-order optical errors, such as spherical aberrations and coma (Fig. 1). During the second half of the last century, several techniques for measuring monochromatic aberrations of the human eye were developed. One of the first was the crossed cylinder aberroscope introduced by Howland and Howland for subjective measurements,7 and the improved version for objective measurements introduced by Walsh et al.8 and Atchison et al.9 Artal, Santamaria, and Bescos10,11 developed a device to measure the point-spread function of the human eye, and introduced methods to calculate the wavefront aberrations from such data. He et al.12 used a psychophysical procedure to measure wavefront aberrations. In 1994, Liang et al.13,14 presented the first measurement of ocular aberrations using a HartmannShack sensor. The operation principle of a HartmannShack wavefront sensor is demonstrated in Figure
Introduction More than 15 years after the discovery that corneal tissue can be micromachined in a nonthermal fashion with remarkable precision, the application of the ArF-excimer laser on the cornea for myopic, hyperopic, and astigmatic corrections has gradually been accepted by the ophthalmic community.1 During the past years, the safety of photorefractive keratectomy (PRK) and laser-assisted in-situ keratomileusis (LASIK) has improved significantly, mainly due to the increasing experience of surgeons and the development of new and more reliable technologies. However, the use of modern ArF-excimer lasers enables surgeons not only to correct sphero-cylindrical errors, but also to correct optical imaging errors (wavefront aberrations) in order to improve the natural optical quality of the human eye.
Address for correspondence: Michael Mrochen, PhD, University of Zurich, Department of Ophthalmology, Frauenklinik Strasse 24, CH-8091 Zurich, Switzerland. e-mail:
[email protected] Lasers in Ophthalmology – Basic, Diagnostic and Surgical Aspects, pp. 159–169 edited by F. Fankhauser and S. Kwasniewska © 2003 Kugler Publications, The Hague, The Netherlands
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Coma
Fig. 1. Scheme of ray tracing in case of different wavefront aberrations. The representations of wavefront aberrations are enlarged. Typical wavefront aberrations have the units in micrometers and the focal length is normally given in millimeters.
Microlens array
Wavefront
Spot diagram at CCD - camera
Fig. 2. Principle of a Hartmann-Shack sensor. In practice, a laser beam is focused at a point on the retina. The emerging beam from this point source is imaged on to a microlens array, which forms a point pattern that is captured by video camera. The pattern obtained is compared with that of an aberration-free pattern, and the wavefront aberrations are computed from the displacement of the points on the unaberrated pattern.
2; the processing of an aberrated wave is depicted on the left-hand side. The incident wave results in a distorted grid of spots in the focal plane of the microlens array. On the right-hand side of Figure 2, a camera image of a distorted wave is shown. This distorted wavefront causes lateral displacement of the spots on the CCD camera. This displacement is equal to the first derivation of the wavefront. Thus,
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the shape of the incident wavefront can be reconstructed on the basis of appropriate curve-fitting algorithms from the spot pattern. Mierdel et al.15 demonstrated the clinical use of an automated aberrometer based on the principles of Tscherning’s aberrometry,4 as depicted in Figure 3. Basically, this ray-tracing method uses the mathematical analysis of a retinal spot pattern that
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Fig. 3. Principle of the Tscherning aberroscope. A set of light rays is projected onto the cornea and refracted on the retina by the optical structures of the eye. The resulting spot pattern is distorted according to the eye’s optical aberrations. This spot pattern is captured by a video camera, using the principles of indirect ophthalmoscopy. Again, the pattern obtained is compared to that of an aberration-free pattern, and the wavefront aberrations are computed from the displacement of the points on the unaberrated pattern.
patient
-
frame grabber
Fig. 4. Scheme of video-based corneal topography. A set of concentric rings is imaged onto the corneal front surface, and the resulting image of the distorted rings is captured by video camera, analyzed, and then the corneal surface is mathematically reconstructed.
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Fig. 5. Periodic table of Zernike basis functions. Subscription n indicates radial order, which gives the row number in the table. Subscription f indicates meridional frequency, which gives the column number in the table.
is grabbed by a video camera. From the deviation of the spot positions to their ideal positions, the first derivative of the wavefront can be calculated and the measured wavefront easily reconstructed. Wavefront sensing provides detailed information on the image quality at the retina of an individual eye. In contrast, corneal topography provides shape information on the anterior front surface. Consequently, only wavefront sensing serves to obtain relevant data in order to rate the imaging quality of an individual eye. However, corneal topography is used even more frequently in ophthalmology in order to determine the optical imaging quality of the anterior surface, which is known to account for approximately 70% of the total refraction of the eye’s optic. Most of the commercial topography systems are based on the placido disc technique (Fig. 4). Here, a set of concentric rings are imaged onto the corneal front surface, and the resulting image of the ring distortions is captured by video camera, analyzed, and the corneal surface mathematically reconstructed. As usual, the data provided by topography measurements are represented in diopters maps. In technical optics, the quality of an optical system is defined by the aberrations of the wavefront from its ideal plane or spherical shape.16 Anwar Gullstrand5 established a complicated mathematical formula to describe such wavefront aberrations, which never found acceptance in technical optics or
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clinical ophthalmology. Zernike17, Thibos et al.18 introduced a more practical mathematical formula which is still used in technical optics, using special polynomials to describe the wavefront aberrations that are associated with special optical errors such as tilt, defocus, astigmatism of different orders, spherical aberrations, coma, and n-fold of the wavefront (Fig. 5). A quantitative measure of the optical imaging quality is the rms-wavefront error (root-mean-square) of the wavefront deviation. An optical system is considered good if all Zernike coefficients are close to zero and, as a consequence, the rms-wavefront error is smaller than 1/14 of the wavelength (Marechal criterion).19,20 However, the ‘average human eye’ only has minimal wavefront errors, indicating that, in principle, the construction of the human eye provides excellent optics, exceeding the Marechal-criterion only by a factor of two.21 However, such minimal aberrations were only measured in a small percentage (5-10%) of the normal population. This discrepancy is mainly due to the high variability of the wavefront aberrations in each individual eye. Consequently, the ‘average human eye’ consists of a good optical quality, while the optical quality of an individual eye is known to be of poor imaging quality. It should be mentioned that various research groups presented data leaning towards an aberration balancing between corneal aberrations and the optical elements within
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the eye that reduce the aberration from the cornea by a certain degree.22,23 In particular, the spherical aberration of the cornea is nearly fully compensated by the intraocular structures of the eye. Wavefront analysis was never a clinical issue however, mainly because there was no therapeutic approach to correct wavefront errors by optical means such as spectacles or contact lenses. This has changed since the introduction of excimer laser treatment of the cornea24 and, more specifically, customized ablations such as wavefront-guided treatments.25,26 Photoablation of the cornea In 1983, Trokel, Srinivasan and Braren27 realized that intense excimer laser light can be used not only for etching plastics, but also for corneal tissue. Such excimer lasers are gas lasers in which the lasing medium inside the resonator consists of a gas mixture that is excited by special pretreatment. Depending on the gas composition used, an excimer laser is able to emit laser light at various wavelengths, although only the 193-nm wavelength has gained clinical attention in ophthalmology. This far ultraviolet light is obtained by means of a mixture of argon (Ar) and fluorine (F) gas inside the laser tube. The pulse duration of such excimer lasers used for refractive surgery is in the order of 20 nsec. Numerous publications have reported the physics of the laser-tissue interaction for 193-nm tissue removal. However, there is still confusion in the literature as to whether this ‘photoablative decomposition’ or simply ‘photoablation’ (laser-tissue interaction process) is photochemical or photothermal in nature, although the best model probably uses a combination of these processes. Basically, the intense 193-nm radiation is mainly absorbed by the collagen macromolecules of the cornea at the start of the laser pulse. Here, the high photon energy of the 193-nm wavelength is capable of disrupting the chemical bonds of the molecules such as C-C (photochemical process). However, the later part of the laser pulse is more strongly absorbed by the tissue than the early part. The major part of the incoming energy is then converted into heat by inter-systemic energy transfer. The increase in temperature results in a breaking of the hydrogen bonds of the tissue water and, as a consequence, the water of the cornea becomes a strong absorber at a wavelength of 193 nm (photothermal process). The material undergoes a phase transition into gas, which is heated to several hundred degrees Celsius during the photoablation process. This hot gas, including the protein fragments, bursts out due to increased pressure. Gas chromatography and mass spectroscopy of the expelled products have revealed molecules typical of thermal changes in the protein fraction. In contrast, the remainder only shows minimal signs of thermal processing, since the photoablation process occurs within a few nanoseconds. Therefore, the photo-
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Fig. 6. Ablation depth per ArF-excimer laser pulse for corneal tissue as a function of the radiant exposure. The intersection of the logarithmic fit (R² = 0.94; p < 0,001) yields an ablation threshold of 50 mJ/cm². The radiant exposure of commercially available medical excimer lasers ranges from 120-250 mJ/ cm².
ablation process was termed ‘cold laser ablation’. The major advantage of ArF-excimer lasers for photorefractive surgery is the high precision obtained from corneal photoablation. A common method of studying photoablation is to irradiate a sample with series of laser pulses, and then to measure the resulting etch crater depth. A collection of reported corneal etch data for 193-nm laser radiation is shown in Figure 6. The ablation threshold, the minimum radiant exposure for tissue removal, has been measured at approximately 50 mJ/cm². Above this threshold, the etch depth per pulse, known as ablation rate, increases in a logarithmic fashion with the radiant exposure. Since currently-used clinical excimer lasers work with a radiant exposure of between 120 and 250 mJ/cm², the resulting ablation rate ranges from 0.2-0.5 µm per pulse (Fig. 6). The use of intensive ultraviolet laser light in photorefractive surgery is accompanied by minimal thermal, mechanical, and actinic damage to the remaining corneal tissue. As mentioned earlier, the ablation products that are expelled with ultrasonic speed are in the physical state of a hot gas, and may have temperatures of more than 500° K.28 This hot gas condenses and creates heat, as well as a thin layer at the ‘cold’ edges of the remaining tissue. However, this ‘pseudo membrane’ disappears within a few days after surgery, due to wound healing. Under standard surgical conditions, the maximum averaged temperature increase at the edge of the irradiated tissue ranges from 5-10° K, which is not considered to be significant.29 In summary, thermal side-effects during excimer laser surgery are minimal and should not induce any kind of inflammation of the cornea. During laser refractive surgery, mechanical damage to the deeper layers of the cornea may originate from acoustic stress waves produced during the ablation process. Kermani and Lubatschowski found pressure waves with an amplitude of 80 bar at a distance of 3 mm behind the cornea when an excimer laser beam, with a diameter of 4 mm and a radiant expo-
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sure of 200 mJ/cm², was applied to the corneal surface.30 These waves travel with sonic velocity through the eye and, in contrast to the shock waves generated by Nd:YAG laser photodisruption, the amplitude falls slowly with distance. However, stress waves of less than 100 bar probably do not induce damage in the cellular ocular structures, except for the corneal endothelium, but might contribute to the development of postoperative subretinal hemhorrages, which have occasionally been reported after laser refractive surgery. Ultraviolet light is known to have the potential to generate actinic damage. Light with wavelengths of 300 nm or less is considered to be cytotoxic and mutagenic because, in this wavelength range, DNA shows strong absorption. However, ArF-excimer laser light penetrates the tissue by less than 1 µm, and so this radiation is unlikely to reach the nucleus of human cells (cytoplasmatic shielding). Cytobiological experiments showed that the initial 193-nm light carries little risk of causing mutagenic changes in mammalian cells.31 However, during photoablation of the cornea, a faint bluish light is observed, the so-called secondary radiation or fluorescence, which includes wavelengths longer than 193 nm and has components in the dangerous 250-300 nm range. It is this secondary radiation that accounts for the mutagenic cellular damage that has been detected by very sensitive assays. The radiant exposure, however, has been determined to be less than 5 µJ/cm², well below the estimated mutagenic threshold of 10 µJ/cm².32,33 Therefore, the mutagenic potential of 193nm laser light is not considered to be significant, in accordance with the clinical absence of neoplasms after photorefractive surgery in millions of cases. Tissue can be removed from a large area in two ways. Either the cornea can be irradiated with a number of laser pulses, each of which with a varied energy distribution, or the eye can be irradiated with a series of laser pulses of uniform irradiance but varying geometry. In case of a Gaussian beam profile, the center of the laser beam contains the highest concentration of energy, which decreases towards the beam periphery. This beam will remove more tissue centrally than peripherally with each pulse, which flattens the cornea for myopic correction. If the irradiance in the periphery was made larger, more tissue would be removed from the edges than from the center, which would steepen the corneal hyperopic correction. In principle, any contour can be transferred to the cornea by controlling the energy distribution with the laser beam. The same effect can be achieved by exposing the corneal surface to a series of circular laser beams of uniform energy density but increasing diameter up to 7 mm. The result of this is that, for example, the center of the cornea receives more laser pulses, and has more tissue removed. Use of a small diameter spot or narrow slit beam scanned across the ablation area allows wide area ablations by means of a laser not generating the high energies required by lasers that use a stationary large
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beam. Therefore, the clinical device can be smaller and less expensive. Recently, small ArF-excimer lasers have been used to perform such ‘scanningspot’ photorefractive surgery. To complete the treatment within a reasonably short period of time, the small diameter spot must be rapidly scanned across the cornea, and the repetition rate must be higher than with large diameter spot excimer lasers. However, small inevitable eye movements that have almost no impact on large diameter refractive surgery must be considered during scanning-spot corneal laser surgery for the precise positioning of each single laser spot, in order to omit increased surface roughness, which may result in a greater stromal wound healing response and haze. Reasonably fast optical eye-tracking systems have been instituted to overcome this problem. Currently, excimer laser systems with large beam diameters of more than 6 mm and which scan with a circle beam diameter of 0.5-2 mm, as well as with a narrow slit beam, are in clinical use. Some systems combine scanning elements with large area beams in an attempt to develop maximum flexibility in the computer-controlled scanning algorithms. Photorefractive keratectomy PRK and photorefractive astigmatic keratectomy have become clinical standard techniques for correcting myopia and astigmatism. During PRK for the correction of myopia, direct flattening is achieved by the removal of a convex-concave lenticule of tissue from the outer surface of the central cornea. The central depth (a0) of the keratectomy is determined by the intended change of refraction, but is even more dependent upon the diameter of the ablation zone.34 The following approximation for a0 (in microns) enables the rapid estimation of this central ablation depth: a0 =
1 ∆D · d 2 3
(1)
where ∆D represents the refractive change in diopters, and d the diameter of the ablation zone in millimeters. For example, a myopic correction of –6 D, with a typical diameter of 6.0 mm of the ablation zone, results in a central ablation depth of 72 µm. On the whole, data on the efficacy of refractive correction after PRK generally report two overall measures: the percentage of eyes that achieve a postoperative refraction within 1 D of emmetropia, and the percentage of eyes that achieve 20/40 or better uncorrected visual acuity. Clinical studies determined refractive success rates of between 80 and 95% for corrections of up to –6 D of myopia, and the range of patients achieving 20/40 or better distance acuity without correction ranges between 80 and 100%.1 The overall incidence of vision-threatening compli-
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Corneal laser surgery for refractive corrections cations such as a loss of best-corrected visual acuity and decreased contrast sensitivity was established to be in the order of 1-2% for corrections of moderate myopia with < –6 diopters. Higher myopic correction leads to a significantly higher regression and to a high risk of scarring during wound healing. Thus, PRK is considered an effective and safe surgical technique for myopic corrections up to –6 D. It is important to relate these success rates to a defined time after surgery, because of the wound healing. Epithelial as well as stromal wound healing has been documented to occur over a period of months after PRK and, therefore, we choose a follow-up time of 12 months. Also, the success rates are helpful for assessing the efficacy and predictability of a procedure, but are not absolute measures of ‘refractive success’. Patients with preoperative refractive errors close to emmetropia may continue to complain of the need for spectacle or contact-lens correction, at least for part of the day. In contrast, patients with a residual myopia of, for example, –1.5 D, may consider the procedure successful if the eye was highly myopic prior to surgery. In addition, glare and halos around sources of light, or loss of contrast sensitivity, may not be detected under conditions typically used for measuring postoperative acuity. Because pupil diameter and variation in pupil diameter affect acuity, measuring visual acuities with a single ambient light will, therefore, not necessarily reflect the acuity experienced by the patient.
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Cutting the flap
Laser treatment
Repositioning of the flap
Fig. 7. Scheme of the LASIK procedure.
Laser in-situ keratomileusis The older literature on excimer laser PRK for myopia describes ablation starting at the corneal surface and extending into the deeper structures of the cornea, such as Bowman’s layer and the stroma. A modification of this technique involves the initial creation of a lamellar flap (average thickness, 120160 µm) of anterior corneal stroma, followed by refractive ablation of the exposed stromal bed. This flap is then repositioned on to the exposed stroma, and good adhesion is usually obtained without the need for sutures (Fig. 7). Known as LASIK, this procedure was particularly investigated in eyes needing high myopic corrections of more than –6 D, for which the precision and stability of the refractive outcome after PRK have been somewhat disappointing. However, today the percentage of eyes achieving a postoperative refraction within 1 D of emmetropia is in the range of 90-98%, and the percentage of eyes achieving 20/40 or better uncorrected visual acuity is in the order of 90-100% in cases of moderate myopia. Other advantages of this application include rapid visual recovery, because no central epithelial defect is created, and a relative decrease in corneal haze compared to surface ablations of similar magnitude. A potential disadvantage of this technique is the large amount of tissue removed during surgery and
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Fig. 8. Currently approved refractive corrections using PRK or LASIK.
laser treatment, commonly leading to an effective central thickness of a ‘normal’ cornea (520 µm) of 260 µm. When applying biomechanical data, the maximum myopic corrections (6-mm ablation zone; 160-µm flap thickness) would be limited to –3 D in
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a thin (500 µm) and –10 D in a thick (600 µm) cornea. Thus, the maximum myopic correction with LASIK is somewhere in the order of –10 D. Besides the risk of creating a residual corneal thickness of less than 250-270 µm, the reported intra- and postoperative complications with LASIK are in the order of 1-2%, comparable with the complication rate that occurs after PRK.35 The acceptable ranges for correcting refractive errors of the eye by means of corneal laser surgery are shown in Figure 8. Optical aberrations after corneal laser surgery Both corneal (topographic irregularities) and total wavefront aberrations are found to be significantly increased after corneal laser surgery. In general, the higher the preoperative refraction, the higher the increase in optical aberrations. Such an increase of optical aberrations impairs the visual performance of the eyes treated. To put this in more detail, scotopic visual measures such as low-contrast visual acuity and glare visual acuity suffer most from refractive corrections. This loss of visual performance is more dominant in dim light (larger pupil) because of the strong dependence of the optical aberrations on pupil size. Seiler et al.36 reported an increased factor of 17.65 (pupil diameter 7 mm) measurable in 15 eyes three months after conventional PRK. Martinez et al.37 have shown that PRK changes the relative contribution of coma-like and spherical-like corneal aberrations by analyzing videokeratographs obtained pre- and 24 months postoperatively, and for a 7-mm virtual pupil, the total wavefront aberration increased 11-fold from the preoperative situation. Furthermore, while pupillary dilation from 3-7 mm in the preoperative eye only caused a nine-fold increase in total wavefront aberrations, the same dilatation caused a 100-fold increase one month after surgery, and an approximately 70-fold increase thereafter. Similar results were reported earlier by Oliver et al.38 Marcos et al.23 studied the optical response to LASIK surgery for myopia from total and corneal aberration measurements. Because LASIK surgery induces changes in the anterior corneal surface, most changes in the total aberration pattern can be attributed to changes in anterior corneal aberrations. However, because of individual interactions of the aberrations in the ocular components, a combination of corneal and total aberration measurements is critical to understanding individual outcomes. Unfortunately, little is known about optical aberrations when comparing PRK and LASIK, and, therefore, there may be different increased factors for PRK and LASIK due to corneal wound healing.39 The increase in higher-order optical aberrations could have a variety of reasons: pupil centration, preoperative internal optics, biomechanical response, and changes in the anterior corneal surface due to epithelial and stromal wound healing. In particular,
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misalignment of the ablation on the cornea, whether this is caused by an initial placement error or by the intraoperative eye movements of the patient, was shown to play an important role in this context.40-43 Alignment of the procedure relative to the eye is a task with six degrees of freedom. Apart from lateral shifts (with a horizontal and a vertical component) and axial movements, rotations with potential components in three axes have to be taken into account. One of the major difficulties in centering any ophthalmic procedure is the fact that the alignment cannot be completely controlled by the operator himself. He relies on the collaboration of the patient, who is asked to fixate on a target during the measurement and the surgical procedure. Alignment errors can occur systematically or randomly. In the first case, an initial displacement of the treatment relative to the reference coordinate system is maintained as a constant offset error and causes a drastic increase, especially in coma-like aberrations. Precise alignment techniques are required in order to avoid this type of centration error. Random decentrations are due to eye movements such as drifts and tremors, and might be practically avoided by active intraoperative eye tracking. Such eye-tracking systems detect the eye’s motion and move the scanning mirrors that are used for positioning of the laser beam, with respect to the detected eye movement. Wavefront-guided corneal laser surgery The aim of customized ablations such as wavefrontguided treatments is, first of all, to avoid an unpleasant increase in higher-order aberrations that might cause the visual performance to deteriorate. Nevertheless, the possibility of improving the optical performance by means of a surgical procedure might further improve the visual outcome to its neuronal limits. A key concept in wavefront-guided refractive procedures is the transfer of the wavefront aberration measurements into an ablation profile to correct the aberrations.15,44 In more detail, the wavefront measured by means of a wavefront sensor is translated into a set of laser spot positions and directly transferred to the laser without taking the subjectively determined refraction data into consideration. Briefly, the plane wavefront of parallel (laser) light entering the eye is distorted after propagating through the optics of the eye at the retinal image plane, according to geometrical irregularities and inhomogeneities in the refractive index. Since the wavefront aberration W(x,y) is defined by optical paths, its dimension is length in meters or in units of the wavelength (Fig. 9). To first order, the wavefront aberrations introduced by the eye are considered to be independent of the wavelength of the light within the visible spectrum. As main refractive surface and site of ablation, we accept the anterior corneal surface with a refrac-
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Fig. 9. Relationship between wavefront, rays, and phase in a point light source example.
Treatment zone
Ablation profile
Fig. 10. Scheme of the transfer of an ablation profile into a set of laser pulses, as used in modern scanning-spot excimer lasers for corneal laser surgery.
tive index of n = 1.337 (tear film) and, in principle, the ablation profile is determined by:
max(W(x,y)) – W(x,y) a(x,y) =
n–1
(2)
This ablation profile a(x,y) has to be approximated by a series of spot ablations, including appropriate overlaps (Fig. 10). The ablation pattern calculated from the wavefront deviation map typically has an optical zone with a diameter of between 6.0 and 7.0 mm, surrounded by a transition zone of up to 2.0 mm. Wavefront-guided treatments are currently under investigation by various groups. However, the first clinical data presented already demonstrate that wavefront-guided treatments are a promising technique
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with the potential for correcting refractive errors, improving visual acuity, and increasing the quality of vision, especially under mesopic conditions. Mrochen et al.45 have reported a prospective study in which 35 eyes of 28 patients were enrolled, with an average preoperative spherical refraction of –4.8 ± 2.3 D and a cylinder of –1.1 ± 0.9 D. Pre- and postoperative wavefront analysis was performed with a Tscherning-type aberrometer. A scanning spot laser with a spot size of 1 mm and a laser repetition rate of 200 Hz was used. The eye tracking system had a response time of less than 6 msec. The treatment area diameter ranged from 6-7 mm, surrounded by a transition zone of 1 mm. At three months, 68% of the eyes were within ± 0.5 D and 93.5% within ± 1.0 D of emmetropia. Unaided visual acuity (UVA) was 20/20 or better in 93.5% of the cases. None of
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the eyes lost more than one line of low contrast, glare, and best-spectacle corrected visual acuity (BCVA). Supernormal vision (20/10 or better in BCVA) was achieved in 16%. The correction of higher-order aberrations (spherical aberration, coma) was insufficient, with an increased factor of the overall rmswavefront error of 1.44 ± 0.74 at three months after surgery. Coma could be better corrected than spherical aberration. Based on these data, wavefront-guided LASIK offers the potential for correcting refractive errors, improving visual acuity, and increasing the quality of vision, especially under mesopic conditions. In order to achieve better correction of the aberrations, further studies are necessary and should include selective overcorrrections of different Zernike components. In addition, further prospectively-controlled clinical studies are needed to clarify the major benefits of wavefront-guided LASIK.
12.
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15.
16. 17.
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Conclusions There has been rapid evolution in the technology for treating patients with corneal laser surgery. Newly refined diagnostic technology, such as topography and wavefront sensing, and more sophisticated spot laser delivery systems with eye tracking, provide the refractive surgeon with much greater flexibility for tackling the often-challenging optical abnormalities of the human eye. The next decade in corneal laser surgery promises to provide huge gains in visual function by increasing retinal image quality. References
19.
20. 21.
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1. Seiler T, McDonnell PJ: Excimer laser photorefractive keratectomy. Surv Ophthalmol 40:89-118, 1995 2. Volkmann AW: Sehen. In: Wagner R (ed) Handwörterbuch der Physiologie III, pp 289-293. Braunschweig: Vieweg und Sohn 1846 3. Helmholtz H: Handbuch der physiologischen Optik, pp 137147. Leipzig: Leopold Voss 1867 4. Tscherning M: Die monochromatischen Aberrationen des menschlichen Auges. Z Psychol Physiol Sinne 6:456-471, 1894 5. Gullstrand A: Die Dioptrik des Auges, V. Die monochromatischen Aberrationen des Auges. In: Von Helmholtz H (ed) Handbuch der physiologischen Optik, 3rd edn, pp 353-367. Leipzig: Leopold Voss 1909 6. Donders FC: Astigmatismus und cylindrische Glaeser. Verlag von Hermann Peters 1862 7. Howland HC, Howland B: A subjective method for the measurement of monochromatic aberrations of the human eye. J Opt Soc Am 67:1508-1518, 1977 8. Walsh G, Charman WN, Howland HC: Objective technique for determination of monochromatic aberrations of the human eye. J Opt Soc Am 1(9):987-992, 1984 9. Atchison DA, Collins MJ, Wildsoet CF, Christensen J, Waterworth MD: Measurement of monochromatic ocular aberrations of human eyes as a function of accommodation by the Howland aberroscope technique. Vision Res 35(3): 313-323, 1995 10. Artal P, Santamaria J, Bescos J: Retrieval of wave aberra-
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tion of human eyes from actual point-spread-function data. J Opt Soc Am 5(8):1201-1206, 1988 Santamaria J, Artal P, Bescos J: Determination of the pointspread function of human eyes using a hybrid optical-digital method. J Opt Soc Am 4(6):1109-1114, 1987 He JC, Marcos S, Webb RH, Burns SA: Measurement of the wave-front aberration of the eye using a fast psychophysical procedure. J Opt Soc Am 15(9):2449-2456, 1998 Liang J, Williams DR, Miller DT: Supernormal vision and high-resolution retinal imaging through adaptive optics. J Opt Soc Am 14:2884-2892, 1997 Liang J, Williams DR: Aberrations and retinal image quality of the normal human eye. J Opt Soc Am 14(11):28732883, 1997 Mierdel P, Wiegand W, Krinke HE, Kaemmerer M, Seiler T: Measuring device for determining monochromatic aberration of the human eye. Ophthalmologe 6:441-445, 1997 Born M, Wolf E: Principles of Optics. Oxford: Pergamon Press 1987 Zernike F: Beugungstheorie des Schneidenverfahrens und seiner verbesserten Form der Phasenkontrastmethode. Physica I 2:689-704, 1934 Thibos LN, Applegate RA, Schwiegerling JT, Webb R: Standards for reporting the optical aberrations of eyes. In: McRae S, Applegate RA, Krueger R (eds) Customized Corneal Ablation: The Quest for Super Vision, pp 348-361. Thorofare, NJ: Slack Inc 2001 Marechal A: Etude des effets combines de la diffraction et des aberrations geometriques sur l’image d’un point lumineux. Rev Opt: 257-277, 1947 Mahajan VN: Aberration theory made simple, Part 2. SPIE Optical Engineering Press 1991 Mrochen M, Kaemmerer M, Mierdel P, Krinke HE, Seiler T: Is the human eye a perfect optic? SPIE Proc Ophthalmic Technol 11(4245):30-36, 2001 Artal P, Berrio E, Guirao A, Piers P: Contribution of the corneal and internal surfaces to the changes of ocular aberrations with age. J Opt Soc Am 19:137-143, 2002 Marcos S, Babero S, Llorente L, Merayo-Lloves J: Optical response to LASIK surgery for myopia from total and corneal aberration measurements. Invest Ophthalmol Vis Sci 42: 3349-3356, 2001 Seiler T, Genth U, Holschbach A, Derse MSO: Aspheric photorefractive keratectomy with excimer laser. Refract Corneal Surg 9:166-172, 1993 Gibralter R, Trokel SL: Correction of irregular astigmatism with the excimer laser. Ophthalmology 101(7):1310-1314, 1994 Mrochen M, Kaemmerer M, Seiler T: Wavefront-guided laser in situ keratomileusis: early results in three eyes. J Refract Surg 16:116-121, 2000 Trokel SL, Srinivasan R, Braren B: Excimer laser surgery of the cornea. Am J Ophthalmol 96:710-715, 1983 Valderame GL, Fredin LG, Berry MJ: Temperature distribution in laser-irradiated tissue. SPIE Proc 1427:200-213, 1991 Bende T, Seiler T, Wollensak J: Side effects in excimer corneal surgery: corneal thermal gradients. Graefe’s Arch Clin Exp Ophthalmol 226:277-280, 1988 Kermani O, Lubatschowski H: Struktur und Dynamic photoakustischer Schockwellen bei der 193 nm Excimerlaserphotoablation. Fortschr Ophthalmol 88:748-753, 1991 Kochevar IE: Cytotoxicity and mutagenicity of excimer laser radiation. Laser Surg Med 9:440-445, 1989 Lubatschowski H, Kermani O: 193 nm Excimerlaserphotoablation der Hornhaut: Spektrum und Transmissionsverhalten von Sekundärstrahlung. Ophthalmologe 89:134-138, 1992 Machette LS, Waynant RW, Royston DD et al: Induction
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of lambda prophage near the site of focused UV laser radiation. Photochem Photobiol 49:161-167, 1989 Munnerlyn CR, Koons SJ, Marshall J: Photorefractive keratectomy: a technique for laser refractive surgery. J Cataract Refract Surg 14:46-52, 1988 Knorz MC, Jendritza B, Hugger P, Liermann A: Komplikationen der Laser-in-situ Keratomileusis (LASIK). Ophthalmologe 96:503-508, 1999 Seiler T, Kaemmerer M, Mierdel P, Krinke HE: Ocular optical aberrations after photorefractive keratectomy for myopia and myopic astigmatism. Arch Ophthalmol 118:1721, 2000 Martinez CE, Applegate RA, Klyce SD, McDonald MB, Median JP, Howland HC: Effect of pupillary dilation on corneal optical aberrations after photorefractive keratectomy. Arch Ophthalmol 116:1053-1062, 1998 Oliver KM, Hemenger RP, Corbett MC et al: Corneal optical aberrations induced by photorefractive keratectomy. J Refract Surg 13:246-254, 1997 Oshika T, Klyce SD, Applegate RA, Howland HC, Danasoury MAE: Comparison of corneal wavefront aberrations after photorefractive keratectomy and laser in situ kerato-
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169 mileusis. Am J Ophthalmol 127(1):1-7, 1999 40. Verdon W, Bullimore M, Maloney RK: Visual performance after photorefractive keratectomy. Arch Ophthalmol 144: 1465-1472, 1996 41. Mrochen M, Kaemmerer M, Mierdel P, Seiler T: Increased higher-order optical aberrations after laser refractive surgery; a problem of subclinical decentration. J Cataract Refract Surg 27:362-369, 2001 42. Guirao A, Williams DR, Cox IG: Effect of rotation and translation on the expected benefit of an ideal method to correct the eye’s higher-order aberrations. J Opt Soc Am 18(5):1003-1015, 2001 43. Bueeler M, Mrochen M, Seiler T: Required accuracy of lateral centration in aberration sensing and wavefront guided treatments. J Cataract Refract Surg 2002 (in press) 44. Klein SA: Optimal corneal ablation for eyes with arbitrary Hartmann-Shack aberrations. J Opt Soc Am A 15(9):25802588, 1998 45. Mrochen M, Kaemmerer M, Seiler T: Clinical results of wavefront-guided LASIK at 3 months after surgery. J Cataract Refract Surg 27:201-207, 2001
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Selective laser trabeculoplasty Mark A. Latina1 and David H. Gosiengfiao2 1 Department of Ophthalmology, Tufts University; 2Massachusetts Eye and Ear Infirmary; Boston, MA, USA
Keywords: Q-switched KTP laser, selective trabeculoplasty, argon laser trabeculoplasty, mechanism of action, clinical efficiency
Argon laser trabeculoplasty The application of lasers on the trabecular meshwork (TM) was pioneered by Worthen and Wickham in 1973.1,2 They described the use of a continuous wave argon laser to photocoagulate the TM, resulting in a significant reduction in intraocular pressure (IOP). The technique was called argon laser trabeculotomy and the results were temporary, lasting only several months. In 1979, Wise and Witter3 described the use of a low energy argon laser not to penetrate the TM, but rather to create a series of superficial scars circumferentially with the end result of decreased IOP. Their work paved the way for the widespread use of lasers in the treatment of open-angle glaucoma. This technique became known as argon laser trabeculoplasty (ALT). Selective laser trabeculoplasty Despite the proven success of ALT, its use has been limited by several factors. Most important of these is the underlying coagulative damage to the TM induced by the procedure. Such changes in the architecture of the TM may ultimately result in trabecular fusion and/or occlusion of the trabecular spaces, possibly leading to obstruction of aqueous outflow. In 1995, using a procedure they called selective laser trabeculoplasty (SLT), Latina and Park4 demonstrated that coagulative damage to the TM is not necessary for achieving IOP reduction. Initial in vitro studies were conducted to evaluate the possibility of selectively targeting the pigmented TM cells without damaging the architecture of the TM.5
This was achieved using a Q-switched, frequen-cydoubled Nd:YAG laser at energy fluences of 30-1000 mJ/cm². A subsequent study using owl monkeys (sp. Aotus trivirgatus) confirmed actual selective targeting of pigmented TM cells in vivo, and evaluated the safety and morphological effects of the procedure in a living system (unpublished data presented at ARVO; Latina MA, Sibayan S: S408, 1996). Light and electron microscopy revealed only disruption of pigmented TM cells with an intact structural architecture and no damage to the trabecular collagen beams. Endothelial membrane formation on the TM, which is usually found in ALT-treated eyes, was not seen. These results were confirmed by Kramer and Noecker6 Using scanning and transmission electron microscopy to compare acute morphological changes in the TM of freshly enucleated human eye bank eyes, they demonstrated crater formation, coagulative damage, fibrin deposition, and disruption of trabecular beams and endothelial cells after ALT, while eyes treated with SLT showed none of these findings. SLT preserved the architecture of the TM. Mechanism of selective laser trabeculoplasty The in vitro and in vivo findings after SLT are observed because the pulse duration of the Qswitched, frequency-doubled (532 nm) Nd:YAG laser is shorter than the thermal relaxation time of melanin.4 Thermal relaxation time defines the absolute time required by a chromophore to convert electromagnetic energy into thermal energy. Melanin has a thermal relaxation time in the microsecond range, while the pulse duration of the Nd:YAG
Address for correspondence: Mark A. Latina, MD, 20 Pond Meadow Drive, Suite 203, Reading, MA 01867, USA
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laser ranges from 3-10 nsec. This means that the pulse duration of this laser is too short for the melanin to convert the laser energy into heat. This spares the surrounding non-pigmented tissues from any coagulative damage. The demonstrable clinical efficacy of SLT suggests that it works on the cellular level, either through migration and phagocytosis of TM debris by the macrophages,7 or by stimulation of the formation of healthy trabecular tissue which may enhance the outflow properties of the TM.8,9 Alvarado10 observed a five- to eight-fold increase in the number of monocytes and macrophages present in the TM of monkey eyes treated with SLT, compared with untreated controls. He theorized that injury to the pigmented TM cells after SLT results in the release of factors and chemoattractants which recruit monocytes. These are activated and transformed into macrophages upon interacting with the injured tissues, and engulf and clear the pigment granules from the TM before exiting the eye to return to the circulation via Schlemm’s canal.11 Clinical efficacy of selective laser trabeculoplasty A pilot study evaluating the IOP lowering effect of SLT in 53 patients whose IOPs could not be controlled by medication, or who had failed traditional ALT, demonstrated the safety and efficacy of SLT.5 The patients were followed for 26 weeks and a mean IOP reduction of 18.7% (4.6 mmHg) with minimal adverse effects was noted. Furthermore, 66% of patients who had previously failed ALT showed an average reduction of IOP of 5.9 mmHg, without any significant adverse effects. In a longer term, randomized, prospective clinical study comparing the effects of SLT versus ALT on the IOP of patients with open-angle glaucoma on medication, Damji et al.7 noted a mean IOP reduction of 6.5 mmHg at 12 months in those patients treated with SLT over 180°. Eyes treated with ALT achieved an IOP reduction of 6.03 mmHg, which was not statistically different from the SLTtreated eyes. Indications for selective laser trabeculoplasty Almost all forms of open-angle glaucoma are candidates for treatment with SLT. However, caution should be used with conditions not amenable to ALT, such as juvenile open-angle glaucoma and secondary inflammatory glaucomas. Selective laser trabeculoplasty as initial therapy The Glaucoma Laser Trial (GLT),12-15 demonstrated that initial treatment with ALT was shown to have
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a greater pressure lowering effect than medication (timolol maleate 0.5%) at two years of follow-up. The GLT study indicated that initial treatment with ALT in patients with primary open-angle glaucoma was at least as effective as intervention with timolol.15 In a similar fashion, SLT has been demonstrated to be safe and efficacious as a first line therapy for open-angle glaucoma. In a study of 26 previously untreated pseudoexfoliative and primary open-angle glaucoma patients, Latina and Smith showed a 30month cumulative probability of successfully remaining off medication of 70% without significant adverse effects after SLT (unpublished data presented at the AAO. Latina MA, Smith J: 2001). Preoperative assessment The preoperative evaluation should include the assessment of corneal clarity and gonioscopy. Corneal clarity is necessary in the delivery of energy. Opacifications in the cornea from edema, scarring, or dystrophy must be noted and appropriate interventions performed. Gonioscopy performed preoperatively will facilitate the laser procedure. Narrow angles may require laser gonioplasty or iridotomy to deepen the angle. Peripheral anterior synechiae should be noted. Extensive synechiae may preclude an effective procedure. Operative technique The operative technique is similar to ALT. Even though SLT is not associated with large IOP spikes, apraclonidine or brimonidine should be given to minimize post-laser IOP elevation. A topical anesthetic such as proparacaine, given immediately prior to the procedure, will also be helpful. A Goldmann three-mirror lens or a Ritch lens may be used, but a Latina lens, optically tuned for the 532 nm wavelength, is preferred for SLT. Goniosolution is placed on the contact lens and the lens is placed on the eye. The aiming beams should be focused on the TM through the center of the mirror of the contact lens in order to ensure a round beam and to maximize energy delivery. The 400-µm aiming beam is large enough to cover the antero-posterior breadth of the TM. SLT has a fixed spot size of 400 µm – as large as the aiming beam. Treatment is begun at 0.8 mJ and increased or decreased in 0.1 mJ increments until the threshold energy is reached. This can be identified by the formation of a cavitation bubble. Treatment is consummated just below the threshold energy. Approximately 50 laser shots are delivered to cover 180° of TM. SLT treatment should be confluent, but non-overlapping.
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Selective laser trabeculoplasty Postoperative period An increased anterior chamber reaction can be expected during the first hour after treatment.5,7 This quickly resolves and usually disappears by the fifth day. A wide variety of anti-inflammatories has been used successfully as post-treatment medication, including indomethacin 0.1% t.i.d. for ten days,16 dexamethasone-neomycin q.i.d. for seven days,17 and prednisolone acetate 1% q.i.d. for seven days.7 Without pre-treatment with apraclonidine, Latina et al. reported IOP spikes of 8 mmHg or greater in 9% of treated eyes, and IOP spikes of 5 mmHg or greater in only 24% of treated eyes.5 Pretreament with apraclonidine 1% resulted in a mean drop of 1.4 mmHg at one hour.7 An IOP spike occurred in only three of 18 treated eyes (maximum, 4 mmHg) in the same study. Within two hours, up to 40% IOP decrease has been observed.16 This was maintained at 24 hours. Gracner reported a 22.6% or 5.12 mmHg mean reduction in IOP at one day, which still remained at six months.17 However, there were cases in which a less precipitous drop in IOP occurred. Hence, clinical decisions regarding repeat treatment should be deferred for at least six weeks. Repeat treatments If IOP reduction is inadequate after six weeks, treatment of the untreated 180° may be contemplated. Because of its non-destructive nature, multiple treatments with SLT are theoretically possible. Repeat treatment with SLT has been found to be safe and effective (unpublished data presented at the AAO. Latina MA, Smith J: 2001). Conclusions In summary, SLT is safe and effective in lowering IOP. Moreover, clinical studies have demonstrated that SLT treatment following failed ALT has a success rate similar to that seen in patients who have not had prior laser treatment. The treatment may be repeated because of the lack of coagulation damage to the TM. Furthermore, it can be considered as a primary treatment option, especially in patients who are medication-intolerant or in those who will not comply with their glaucoma medication, without interfering with the success of future surgery. Due to its non-destructive properties and low
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173 complication rate, SLT has the potential to become an ideal first line treatment for open-angle glaucoma. References 1. Worthen DM, Wickham MG: Laser trabeculotomy in monkeys. Invest Ophthalmol Vis Sci 12:707-711, 1973 2. Worthen DM, Wickham MG: Argon laser trabeculotomy. Trans Am Acad Ophthalmol Otolaryngol 78:371-375, 1974 3. Wise JB, Witter SL: Argon laser therapy for open-angle glaucoma: a pilot study. Arch Ophthalmol 197:319-322, 1979. 4. Latina M, Park C: Selective targeting of trabecular meshwork cells: in vitro studies of pulse and continuous laser interactions. Exp Eye Res 60:359-372, 1995 5. Latina MA, Sibayan SA, Shin DH, Noecker RJ, Marcellino G: Q-switched 532-nm Nd:YAG laser trabeculoplasty (selective laser trabeculoplasty). Ophthalmology 105:20822090, 1998 6. Kramer TR, Noecker RJ: Comparison of the acute morphologic changes after selective laser trabeculoplasty and argon laser trabeculoplasty in human eye bank eyes. Ophthalmology 108:773-779, 2001 7. Damji KF, Shah KC, Rock WJ et al: Selective laser trabeculoplasty vs argon laser trabeculoplasty: a prospective randomised clinical trial. Br J Ophthalmol 83(6):718722, 1999 8. Dueker DK, Norberg M, Johnson DH et al: Stimulation of cell division by argon and Nd:YAG laser trabeculoplasty in cynomolgous monkeys. Invest Ophthalmol Vis Sci 31:115124, 1990 9. Bijlsma SS, Samples JR, Acott TS, Van Buskirk EM: Trabecular cell division after argon laser trabeculoplasty. Arch Ophthalmol 106:544-547, 1988 10. Alvarado JA: Mechanical and biochemical comparison of ALT and SLT. Ocul Surg News March 7-10, 2000 11. Alvarado JA, Murphy CG: Outflow obstruction in pigmentary and primary open angle glaucoma. Arch Ophthalmol 110: 1769-1778, 1992 12. Glaucoma Laser Trial Research Group: Acute effects of argon laser trabeculoplasty on intraocular pressure. Arch Ophthalmol 107:1135-1142, 1989 13. Glaucoma Laser Trial Research Group: The Glaucoma Laser Trial II: Results of argon laser trabeculoplasty versus topical medicines. Ophthalmology 97:1403-1413, 1990 14. Glaucoma Laser Trial Research Group: The Glaucoma Laser Trial VI: Treatment group differences in visual field changes. Am J Ophthalmol 120:10-22, 1995 15. Glaucoma Laser Trial Research Group: The Glaucoma Laser Trial VII: The glaucoma laser trial (GLT) and glaucoma laser trial follow-up study results. Am J Ophthalmol 120:718731, 1995 16. Lanzetta P, Menchini U, Virgili G: Immediate intraocular pressure response to selective laser trabeculoplasty. Br J Ophthalmology 83:29-32, 1999 17. Gracner T: Intraocular pressure response to selective laser trabeculoplasty in the treatment of primary open-angle glaucoma. Ophthalmologica 215:267-270, 2001
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Photocoagulation, transpupillary thermotherapy and photodynamic therapy for choroidal neovascularization Richard S.B. Newsom1, Adam H. Rogers2 and Elias Reichel2 Southampton Eye Hospital, Southampton, UK; 2New England Eye Center, Tufts University School of Medicine, Boston, MA, USA
1
Keywords: photocoagulation, thermotherapy, photodynamic therapy, senile, neovascular, exudative maculopathy
Introduction
Table 1.
Potential treatments for retinal disease are becoming more diverse as the pathogenic mechanisms underlying retinal disease are better understood and laser-tissue interactions better refined. At present, lasers are used in photomechanical, photothermal, photochemical, and photodynamic modalities specific for disease entities. The range of laser treatments is likely to increase with the possibility of adjunctive therapies to enhance treatments. Within this chapter, we aim to discuss developments in laser and photodynamic therapy for age-related macular degeneration. Photocoagulation
Native macula chromophores Melanin RPE and choroid Xanthophill inner and outer plexiform layers Photo pigments photoreceptor outer segments Lipofucin RPE Hemoglobin choroidal and retinal vessels Exogenous chromophores in AMD Lipofuscin, melanolipofuscin RPE Metabolic debris retina/RPE Blood subretinal, RPE Cellular and fibrous subretinal, sub-PE
400-1000 nm 420-500 nm 420-780 nm
450 and 550 nm
420-510 nm
not characterized
Background Photocoagulation relies on the conversion of light energy to heat by retinal chromophores (Tables 1 and 2).1,2 The total energy released depends on light wavelength (λ), irradiance (power per unit area), and retinal chromophore concentrations.3 Threshold retinal photocoagulation occurs when the retinal temperature is raised by 20°C for ten seconds, or by 39°C for 0.1 seconds.4 Clinically, lasers are used at suprathreshold levels, raising retinal temperatures between by 40-60°C.5 The energy needed to achieve threshold coagulation increases in proportion to pulse length. During longer pulses, proportionally more heat is conducted from irradiated tissues.6,7 The highest temperatures generated occur at the center of the laser spot, while temperatures at the burn periphery are lower, due to light scatter, heat conduction and eye movements.8 The effects of tem-
Table 2. Lasers
Wavelength
Argon Doubled YAG Yellow Red Krypton Diode
488 532 577 630 647 810
+ 514 nm nm nm nm nm nm
perature gradients along a laser spot are minimized with short 0.1-0.5-second exposures.9 Retinal pigment epithelium (RPE) melanin is the primary chromophore during photocoagulation. Retinal hemoglobin and xanthophyll, absorbing light between 400-550 nm, are key chromophores for argon green (514 nm) and frequency-doubled YAG
Address for correspondence: Elias Reichel, MD, New England Eye Center, Tufts University School of Medicine, 750 Washington Street, Box 450, Boston, MA 02111, USA. e-mail:
[email protected] Lasers in Ophthalmology – Basic, Diagnostic and Surgical Aspects, pp. 175–182 edited by F. Fankhauser and S. Kwasniewska © 2003 Kugler Publications, The Hague, The Netherlands
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Table 3.
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Micropulse
TTT
PDT
Wavelength
514-810 nm
532-810 nm
810 nm
689 nm
Pulse duration
0.2-0.05
0.001
1 minute
83 seconds
Laser-tissue interaction
photothermal coagulation
photothermal photobiological
photothermal photobiological
photochemical photobiological
Retinal irradiance
80 W/cm2
7.5 W/cm2
0.6 W/cm2
Maximum temperature rise
42°
10°
2-4°
42°
(532 nm) lasers. RPE and choroidal melanin absorb longer wavelengths of light between 600 and 1000 nm.10 Thus, the inner retina is heated more with shorter wavelengths and the outer retina/choroid with longer wavelengths.11 Histological studies12 show argon blue-green lasers (488 and 514 nm) cause full thickness burns, and argon green lasers (514 nm)13 and frequency-doubled YAG lasers form cone shaped burns mainly effecting photoreceptors and RPE.14 Krypton (647 nm) and diode lasers (810 nm) cause outer retina coagulation, since most of their energy passes unabsorbed through macular xanthophyll and hemoglobin.15,16 However, at suprathreshold levels, all the aforementioned lasers can cause full thickness retinal burns. When treating choroidal neovascularization (CNV), light energy is converted to heat energy, which denatures proteins within the surrounding tissue and causes intravascular capillary coagulation. Green.17 found histological evidence of CNV obliteration following photocoagulation, but also evidence of recurrence in nine (75%) lesions. A scar comprised of hyperplastic retinal pigment epithelium was noted. Inner retinal layers were preserved following krypton photocoagulation, however, blue-green argon caused full-thickness destruction of the retina.17 RPE cells at the edge of the lasered site react by spreading, migrating,18,19 and releasing inhibitory growth factors.20 CNV resolution may also depend on altering growth factor expression.21 Studies of RPE following photocoagulation show up-regulation of TGF-β, VEGF, IL-8, and ETS-1.22,23 However, several weeks following treatment of CNV, VEGF down-regulation occurs, due to increased retinal oxygenation from the choroidal circulation.24 The damaging collateral effects of retinal photocoagulation have driven laser research to target specific retinal layers. Lowering laser irradiance with micropulsed treatments or subthreshold levels avoiding retinal coagulation and photoreceptor loss are gaining acceptance (Table 3).5,8,25,26
be targeted using laser pulses shorter in duration than the thermal relaxation time of a given tissue (adiabatic heating).25 However, very short pulses may generate micro-explosions and other thermomechanical effects offsetting the beneficial effects of shorter pulse duration. Thermomechanical effects of shortpulsed lasers can be prevented using lower power micropulses. An animal model of argon laser micropulsing was developed,25 and is now used for diode laser applications.27-30 Initial studies reported positive results for treatment of diabetic macular edema, showing improvement in 29% and stabilization of acuity in 69%.26 Other authors reported that micropulsed lasers were less painful and clinically effective in treating macular edema.27 Friberg and Karatza28 observed clinical resolution of macular edema from branch retinal vein occlusion in 92% of eyes, with visual acuity stabilizing in 77%. Subthreshold treatment by placing the laser in continuous wave (cw) mode has also been shown to reduce energy transmitted to the retina, thus avoiding photocoagulation.5 Judging the endpoint of subthreshold treatments is difficult, due to the lack of a visible retinal reaction. Some investigators have suggested using suprathreshold treatment to adjacent tissue as a guide to titrating laser power, followed by a 50% power reduction when treating the intended choroidal lesion.27 However, retinal chromophore composition may vary over a short distances, so that laser uptake may differ between the two sites, leading to variable laser reactions. This form of photocoagulation has been criticized, as it tends to create variable response in the retinal pigment epithelium.29 Subthreshold laser treatment of retinal disease may mirror the trend to use lower power lasers in other medical applications such as low power laser irradiation (LPLI). LPLI in the visible and infrared regions causes bio-stimulation of cellular processes which clinically accelerates wound healing and tissue repair.30-34 Similar tissue reactions may occur following sub-threshold retinal laser treatments.
Subthreshold photocoagulation
Current indications for photocoagulation
Attempts at limiting collateral retinal damage from suprathreshold laser photocoagulation have led to a focus on subthreshold techniques. Initial animal studies demonstrated that specific retinal layers could
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Prophylactic laser photocoagulation of soft drusen is a controversial treatment intended to stimulate
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Photocoagulation, transpupillary thermotherapy and photodynamic therapy drusen re-absorption and theoretically diminish the risk of CNV. In 1973, Gass first reported his observation that laser applied to drusen. 35 This initial finding was corroborated by other reports demonstrating a beneficial effect of laser to drusen. Little et al.36 reported on 27 patients with bilateral drusen. Laser photocoagulation was applied to drusen in one eye with the second eye acting as a control. All 27 eyes treated with laser demonstrated resolution of the drusen with an associated 1.2-line improvement in visual acuity. Seven percent of treated eyes progressed to CNV, compared with 15% in the control group. While the results of these individual reports appear to be impressive, an associated increased incidence in laser-induced CNV appears to be a complication of this treatment. Guymer et al.37 observed CNV following photocoagulation in patients with drusen. In these patients, histology demonstrated choroidal endothelial processes breaching the elastic lamina, possibly representing early changes prior to choroidal neovascularization. In 1998, the Choroidal Neovascularization Prevention Trial Research Group found that CNV occurred in 5% of treated eyes compared to 2% in control eyes.38 With this degree of controversy surrounding photocoagulation treatment to drusen, Olk et al.39 explored the use of subthreshold micropulsed diode laser to drusen. Sixty-five percent of patients had drusen resorption with an improvement in vision. No increased incidence in the formation of CNV was reported.39 Although this latest modality appears to be promising, studies of drusen prophylaxis will need to be re-evaluated in the light of recent (AREDS) study, which showed vitamins C and E, beta-carotene and zinc could reduce progression of visual loss and neovascular events in late age-related macular degeneration (AMD) by an odds ratio of 0.72.40 Current indications for photocoagulation of choroidal neovascularization Extrafoveal CNV The macular photocoagulation study (MPS) demonstrated a beneficial effect of treating extrafoveal choroidal neovascularization with argon laser photocoagulation in eyes with AMD.41 Lesions in this group occurred a minimum of 200 µm from the geometric center of the foveal avascular zone (FAZ). Forty-five percent of the treated group, compared with 63% of the untreated group, suffered severe visual loss (a loss of six or more lines measured on an ETDRS chart) at three years. This trend was observed to continue at five years, with 46% of treated eyes and 64% of observed eyes showing severe visual loss (SVL). At five years, treated eyes had lost 5.2 lines and untreated eyes 7.1 lines of visual acuity. Recurrent CNV proved to be a major problem with laser photocoagulation, occurring in 54%
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of laser treated eyes at five years. Seventy-five percent of recurrent CNV occurred within the first year following photocoagulation, with 80% of these eyes suffering SVL.41 Photocoagulation of CNV secondary to non-AMD etiologies fared better. In lesions secondary to presumed ocular histoplasmosis (POHS), 10% of the treated group compared with 43% of the control group lost six or more lines of vision. Patients with idiopathic CNV also fared better, since 23% of the treated group compared to 48% of the control group lost six or more lines of visual acuity at five years, with most of the vision loss occurring within the first six months.41 Juxtafoveal CNV Treatment results were less clear for classic juxtafoveal CNV occurring between 1 and 199 µm from the geometric center of the FAZ. Severe vision loss in the treated group compared to the control group was 55% versus 65%.42 In this trial, 54% of patients suffered recurrence at one year and 78% at five years. The median visual acuity was 20/200 in the treated group compared to 20/250 in the control group. The mean number of lines lost in the treated group was 5.5, compared to 6.5 in the control group. Patients with juxtafoveal CNV secondary to POHS fared better than eyes with AMD, since they exhibited fewer recurrences and better preservation of vision. SVL occurred in 12% of the treated group compared to 28% of the control group. The mean change in acuity was a loss of 0.1 lines in the treatment compared to 2.1 lines in the control group. The eyes in the idiopathic CNV group exhibited similar results, with 21% of the treated group compared to 34% of the control group experiencing SVL at five years. For extrafoveal and juxtafoveal classic lesions, laser photocoagulation remains the standard of treatment. Although the TAP Study43 focused on subfoveal disease, a small subgroup of patients with juxtafoveal classic lesions was included in the clinical trial. It is interesting to note that these patients tended to show a greater visual benefit than those patients who had true subfoveal lesions, suggesting a role for photodynamic therapy in treating juxtafoveal CNV. However, a randomized clinical trial is necessary to prove this point. Subfoveal CNV Photocoagulation for classic, subfoveal choroidal neovascularization has demonstrated modest prevention of SVL compared to the natural history of classic CNV.44 Three months following enrollment, 20% of the treated eyes, compared to 11% of the control eyes, had lost six or more lines of visual acuity. Fortyeight months following randomization, the efficacy of treatment was realized with 23% of the treated group losing six or more lines of acuity compared
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to 45% of the control group. A mean of 3.5 lines was lost in the treated group compared to 5.0 lines in the control group. The limiting feature of foveal ablation for subfoveal CNV is an immediate and marked decline in visual acuity. With this technique, there is generally a four-line decline in visual acuity immediately following treatment. Despite a statistically significant difference in visual acuity between treated and untreated eyes, conventional photocoagulation of subfoveal CNV has not gained wide acceptance. Techniques aimed at treating subfoveal CNV, while sparing the surrounding neurosensory retina, are gaining in popularity. Photodynamic therapy (PDT) and transpupillary thermotherapy (TTT) offer, for the first time, treatments that rely on gradual involution of CNV while sparing the neurosensory retina. Transpupillary thermotherapy Transpupillary thermotherapy (TTT) treats CNV by producing a mild rise in retinal temperature that has been calculated to be between 4 and 10°C using biophysical models developed by Mainster.8 The choroid acts as a heat sink by dissipating excessive heat produced, thus preventing excessive temperature elevations and retinal coagulation.8 An elevation of temperature to 41°C leads to the inhibition of DNA, RNA, protein synthesis, and respiratory enzymes. Above a temperature of 43°C, extensive free radical release and protein denaturation occurs.8 Heat shock proteins (Hsps) are theoretically produced during TTT. Hsps may protect cells against the deleterious effects of hyperthermia, radiation, ischemia, hypoxia, cytokines, oxygen free radicals, and metabolic poisons.48-50 However, the primary role of Hsps is as a molecular chaperon aiding in the folding of unstable newly translated proteins.51-53 In heat stressed cells, they bind thermally denatured proteins and play an important role in generating thermotolerance.53 They also prevent apoptosis (programmed cell death) inhibiting the action of the capsase cascade.49 HSP 70, a specific Hsps, is present in all retinal layers, except the outer segments.54 The primary site of synthesis of Hsps following hyperthermia is the photoreceptor layer,55 since they protect the retina against phototoxicity. Following diode laser TTT, HSP 70 is expressed in the RPE and choroidal vessels45 controlling apoptotic events within the RPE and choroidal vessels. Several reports have demonstrated positive results of TTT for the treatment of occult subfoveal CNV. Reichel et al.46 reported on 16 eyes of 15 patients with symptomatic visual loss from occult CNV secondary to AMD. Three eyes (19%) showed a two or more line improvement in visual acuity. Visual acuity remained stable (no change or one-line improvement) in nine treated eyes (56%). The remaining four eyes (25%) showed a decline of one or more
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lines of visual acuity. Fifteen eyes (94%) demonstrated decreased exudation on fluorescein angiography and optical coherence tomography.46 Newsom et al.47 demonstrated similar results in eyes with occult CNV. Treatment of classic CNV was also evaluated in this study, with 80% of eyes demonstrating stable vision within two lines of their pretreatment visual acuity. Most investigators have found that between 70% and 80% of patients with occult CNV demonstrate stabilization, with similar results for classic CNV. While the follow-up in these studies is relatively short compared to publications evaluating argon laser photocoagulation, TTT appears to decrease the rate of recurrent CNV. The scope of the usefulness of TTT continues to broaden. Investigators have shown that TTT is effective among different racial groups. TTT has also been used to treat CNV associated with myopia, angioid streaks, and idiopathic etiologies, and in individuals with idiopathic polypoidal choroidal vasculopathy. Reported complications from TTT are rare. Corneal burns and cataract do not appear to be a significant complication. TTT may cause transient visual loss in 2-5% of patients, and has been shown to temporarily reduce choroidal blood flow similar to PDT.56 There have been reports of negative outcomes in patients with serous PED treated with TTT, 57 though this complication has been observed in patients who have been treated with conventional photocoagulation and photodynamic therapy. In some reported series, 10% of occult CNV converted to classic lesions following TTT. These patients can be effectively treated with PDT.58 In natural history studies, 30-50% of occult lesions demonstrate the development of classic lesions. While conversion to classic lesions following TTT may merely represent the natural history of occult CNV, TTT may accelerate this process. The ‘TTT4CNV’ study, a multicenter, randomized, blinded trial comparing TTT with sham therapy for occult subfoveal lesions, will provide more information regarding occult conversion to classic and the effectiveness of TTT compared with the natural history of disease. Practical considerations Thermal modelling has provided additional information with regard to TTT. For long exposures and large spot sizes, the energy required is proportional to the diameter of the spot size. The calculation of the spot size at the retina is critical in attaining therapeutic retinal irradiances.8 The authors have found that a 1.0 × magnification lens is useful for lesions of up to 3000 µm in diameter and a 2 × magnification lens for lesions of up to 6000 µm. Power settings are increased proportionally to the laser spot diameter. However, in lesions greater than 4000 µm in diameter, the calculated power may have to be reduced by 20-40%. Maintenance of beam circularity is important for preventing astigmatism and ensuring equal retinal irradiance.59 Avoidance of exces-
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Photocoagulation, transpupillary thermotherapy and photodynamic therapy sive ocular pressure by the contact lens, during TTT is crucial for successful treatment. Increased pressure on the globe may diminish choroidal blood flow, reducing the capacity of the choroid to dissipate generated heat that may result in retinal coagulation. The endpoint of treatment with TTT is to visualize no retinal whitening. Subtle retinal changes can be observed by viewing the retina with a narrow slit beam centered on the red circular aiming beam. Variations in retinal pigmentation must be considered when treating with TTT. Darkly pigmented fundi produce higher temperatures for a given irradiance than paler fundi, which can lead to over treatment. Similarly, pigment clumps in the RPE following previous laser photocoagulation may cause focal hyperthermia and retinal coagulation, potentially giving rise to recurrent CNV in that area.5 The presence of subretinal blood also increases diode laser uptake, leading to overtreatment. Shallow serous retinal elevation may also require lower power settings, since they may be at a higher risk for retinal damage. If whitening is observed, then treatment should be stopped and the patient closely monitored postoperatively. The time taken to reach a stable temperature is approximately 0.2 seconds, so if the TTT treatment is interrupted for any reason,60 only the remaining time is needed to complete the treatment. Re-treatments of TTT are necessary in 2050% of occult membranes.46-47 Following TTT, CNV membranes slowly close and re-treatments are not usually performed before six weeks for classic membranes and three months for occult membranes.46-47 Different properties of classic CNV compared to occult CNV may explain the need for earlier and more frequent re-treatments using TTT. TTT is an emerging treatment with great potential. Clinical experience has shown there is variability in reaction to TTT possibly due to variations in blood flow and chromophore concentrations. Developing techniques to monitor individual responses to this subthreshold treatment will be important in developing this technology. Photodynamic therapy Photodynamic therapy has revolutionized the treatment of classic, subfoveal CNV. Verteporfin, a benzoporphyrin derivative monoacid ring A,61-63 is an intravenously injected drug that is infused over a tenminute period, followed by a five-minute pause. Verteporfin accumulates in areas of neovascularization and normal blood vessels by binding to low-density lipoproteins (LDL) receptors that are expressed on endothelial and tumor cells.62,64,65 Fifteen minutes after the start of intravenous infusion, the verteporfin is activated by a low power laser (λ 689 nm). When treating CNV, the spot size used is 1000 µm larger than the greatest linear dimension of the lesion. Involution of the CNV occurs through the
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production of free radicals (type 1 reaction) and the formation of singlet oxygen (type II reaction) after light activation of the photosensitizing agent.63 Tissue damage is characterized by membrane peroxidation,66 nuclear damage, protein denaturation, organelle and apoptotic mechanism damage.66 Lesions treated with PDT demonstrate histological evidence of vascular occlusion, endothelial cell damage, platelet aggregation, histamine, tumor necrosis factor-a (TNF-a), and cytokine release.67-72 Preclinical trials have shown that liposomal preparations and intravenous preparations were effective in closing experimental CNV.65 The Treatment of Age-related Macular Degeneration with Photodynamic Therapy (TAP) Study Group reported positive outcomes for patients with predominately classic CNV (classic component of CNV equal to or greater than 50% of the lesion) treated with PDT using verteporfin. Moderate visual loss (MVL), consisting of a loss of three or more lines Snellen acuity or 15 or more ETDRS letters, was reduced from 62% in treated eyes and 47% in control eyes 24 months after randomization.43 Severe visual loss was reduced from 30% in the control group to 18% in the treated group. However, patients with prior laser treatment, older age (greater than 75 years), and lesions with fibrosis showed no benefit from PDT. Results for occult CNV without evidence of classic CNV were reported in the Verteporfin in Photodynamic Therapy (VIP) Study Group. The results were not as dramatic as those reported by the TAP study group. Fifty-five percent of the treated group compared to 68% of the control group had MVL at two years.73 Only one subgroup of patients, those having either lesions greater than four disc areas in size or relatively poor vision (less than 65 ETDRS letters), showed benefit at two years. A sudden decrease in vision following verteporfin treatment occurred in 4.4% of eyes, due to subretinal bleeding or choriocapillaris occlusion. Fifty percent of these affected eyes regained some vision. These study findings have been corroborated by the authors’ clinical experience with verteporfin. The VIP study group similarly reported positive results for the treatment of subfoveal CNV secondary to myopia. Twelve months following randomization, 14% of the verteporfin treated eyes compared to 33% of the placebo-treated eyes had lost more than 15 letters (p < 0.01).74 Two years after randomization, no statistical benefit between verteporfintreated eyes and control eyes could be identified. One drawback of PDT is that many patients needed multiple treatments (five on average in both the TAP and VIP studies).43,73 Other potentially life-threatening reactions have been reported, as well as local intravenous site complications.75 Ocular complications include pigment epithelial rips following the treatment of classic CNV.76 It is important to note that the prevention of severe vision loss in treating classic CNV is no different from when compared to conventional laser photocoagulation after two years’ follow-up.43,44
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Similarities exist between the effects of TTT and PDT. Both induce involution of CNV while avoiding photocoagulation of the neurosensory retina, thereby stabilizing visual acuity. Transient interruption of choroidal blood flow is observed angiographically.70 TTT and PDT both release free radials within the choroidal blood vessels and injure the vascular endothelium, resulting in closure of the CNV. However, one difference is that PDT closes membranes within 24 hours of treatment.70 These usually re-open and require re-treatment. Membranes treated with TTT appear to close in more gradually, which appears to protect against revascularization of the treated CNV. Therefore, TTT-treated lesions appear to require fewer retreatments compared to PDT. Post-treatment angiographic changes are subtle following both treatments, and may be difficult to interpret. Membranes may take several weeks to regress and continue to stain even after stabilization. Subretinal fibrosis is identified clinically, by redfree photos, and on angiography when the treated lesion develops a scalloped edge. Our experience is that serial OCT images are useful in monitoring the progression of the CNV when angiographic features become quiescent.77 This enables the ophthalmologist to monitor the activity of the CNV through the presence or absence of subretinal and intraretinal fluid collection. The treatment of serous pigment epithelial detachments (PED) has not been established, however, PDT or TTT may be of use when CNV can be demonstrated on ICG. There is a high risk of rips in the retinal pigment epithelium with either procedure. Feeder vessel photocoagulation before or after PDT/ TTT may also have possible therapeutic advantages in that laser energy can be targeted directly to a few abnormal vessels.78 Again, however, the problems of damage to Bruch’s membrane and suprathreshold photocoagulation mean that there can be a relatively high rate of recurrence in CNV treated in this manner. Conclusions Macular photocoagulation for subfoveal and juxtafoveal CNV is associated with high rates of visual loss and recurrence of CNV. New therapies such as TTT and PDT treat CNV with less injury to the surrounding neurosensory retina. PDT has demonstrated good efficacy for classic CNV associated with AMD and presumed ocular histoplasmosis, and moderate efficacy for occult CNV. Initial reports for TTT have been supportive for the treatment of both occult and classic CNV. Further developments in subthreshold laser treatments may also be useful for the prophylactic treatment of early AMD. However, without a clearly observable endpoint, new techniques for evaluating laser tissue interactions may be necessary before this treatment becomes widespread. Huge strides in our understanding of laser
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tissue interactions have allowed laser treatments to be refined and tailored to individual conditions, however, adjunctive therapies may hold the final answer for challenging retinal conditions such as AMD. References 1. Mainster MA: Wavelength selection in macular photocoagulation: tissue optics, thermal effects, and laser systems. Ophthalmology 93(7):952-958, 1986 2. Liao JC, Roider J, Jay DG: Chromophore-assisted laser inactivation of proteins is mediated by the photogeneration of free radicals. Proc Nat Acad Sci US 91(7):2659-2663, 1994 3. Mainster MA, White TJ, Allen RG: Spectral dependence of retinal damage produced by intense light sources. J Opt Soc Am 60(6):848-855, 1970 4. Cain CP, Welch AJ: Measured and predicted laser-induced temperature rises in the rabbit fundus. Invest Ophthalmol 13:60-70, 1974 5. Mainster MA: Decreasing retinal photocoagulation damage: principles and techniques. Sem Ophthalmol 14(4):200-209, 1999 6. Roider J, El Hifnawi ES, Birngruber R: Bubble formation as primary interaction mechanism in retinal laser exposure with 200-ns laser pulses. Lasers Surg Med 22(4):240-248, 1998 7. Mainster MA, Sliney DH, Belcher CD III, Buzney SM: Laser photodisruptors: damage mechanisms, instrument design and safety. Ophthalmology 90(8):973-991, 1983 8. Mainster MA, Reichel E: Transpupillary thermotherapy for age-related macular degeneration: long-pulse photocoagulation, apoptosis, and heat shock proteins. Ophthalmic Surg Lasers 31(5):359-373, 2000 9. Roider J, Lindemann C, El-Hifnawi el S, Laqua H, Birngruber R: Therapeutic range of repetitive nanosecond laser exposures in selective RPE photocoagulation. Graefe’s Arch Clin Exp Ophthalmol 236(3):213-219, 1998 10. Trempe CL, Mainster MA, Pomerantzeff O, Avila MP, Jalkh AE, Weiter JJ et al: Macular photocoagulation: optimal wavelength selection. Ophthalmology 89(7):721-728, 1982 11. Mainster MA, White TJ, Tips JH, Wilson PW: Retinaltemperature increases produced by intense light sources. J Opt Soc Am 60(2):264-270, 1970 12. Smiddy WE, Fine SL, Quigley HA, Hohman RM, Addicks EA: Comparison of krypton and argon laser photocoagulation: results of stimulated clinical treatment of primate retina. Arch Ophthalmol 102(7):1086-1092, 1984 13. Smiddy WE, Fine SL, Green WR, Glaser BM: Clinicopathologic correlation of krypton red, argon blue-green, and argon green laser photocoagulation in the human fundus. Retina 4(1):15-21, 1984 14. Mosier MA, Champion J, Liaw LH, Berns MW: Retinal effects of the frequency-doubled (532 nm) YAG laser: histopathological comparison with argon laser. Lasers Surg Med 5(4):377-404, 1985 15. Mainster MA, Ho PC, Mainster KJ: Argon and krypton laser photocoagulators. Ophthalmology Suppl:48-54, 1983 16. Brancato R, Pratesi R, Leoni G, Trabucchi G, Vanni U: Histopathology of diode and argon laser lesions in rabbit retina: a comparative study. Invest Ophthalmol Vis Sci 30(7):1504-1510, 1989 17. Green WR: Clinicopathologic studies of treated choroidal neovascular membranes: a review and report of two cases. Retina 11(3):328-356, 1991
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Photocoagulation, transpupillary thermotherapy and photodynamic therapy 18. Nicolaissen B Jr: Argon laser lesions in the human RPE in vitro. Acta Ophthalmol (Kbh) 66(3):277-285, 1988 19. Del Priore LV, Glaser BM, Quigley HA, Green WR: Response of pig retinal pigment epithelium to laser photocoagulation in organ culture. Arch Ophthalmol 107(1):119-122, 1989 20. Yoshimura N, Matsumoto M, Shimizu H, Mandai M, Hata Y, Ishibashi T: Photocoagulated human retinal pigment epithelial cells produce an inhibitor of vascular endothelial cell proliferation. Invest Ophthalmol Vis Sci 36(8):16861691, 1995 21. Lip PL, Belgore F, Blann AD, Hope-Ross MW, Gibson JM, Lip GY: Plasma VEGF and soluble VEGF receptor FLT1 in proliferative retinopathy: relationship to endothelial dysfunction and laser treatment. Invest Ophthalmol Vis Sci 41(8):2115-2119, 2000 22. Ogata N, Ando A, Uyama M, Matsumura M: Expression of cytokines and transcription factors in photocoagulated human retinal pigment epithelial cells. Graefe’s Arch Clin Exp Ophthalmol 239(2):87-95, 2001 23. Ishida K, Yoshimura N, Yoshida M, Honda Y: Up regulation of transforming growth factor-beta after panretinal photocoagulation. Invest Ophthalmol Vis Sci 39(5):801-807, 1998 24. Stefansson E, Machemer R, De Juan E Jr, McCuen BW II, Peterson J: Retinal oxygenation and laser treatment in patients with diabetic retinopathy. Am J Ophthalmol 113(1): 36-38, 1992 25. Roider J, Michaud NA, Flotte TJ, Birngruber R: Response of the retinal pigment epithelium to selective photocoagulation. Arch Ophthalmol 110(12):1786-1792, 1992 26. Moorman CM, Hamilton AM: Clinical applications of the MicroPulse diode laser. Eye 13(2):145-150, 1999 27. Stanga PE, Reck AC, Hamilton AM: Micropulse laser in the treatment of diabetic macular edema. Sem Ophthalmol 14(4):210-213, 1999 28. Friberg TR, Karatza EC: The treatment of macular disease using a micropulsed and continuous wave 810-nm diode laser. Ophthalmology 104(12):2030-2038, 1997 29. Pollack JS, Kim JE, Pulido JS, Burke JM: Tissue effects of subclinical diode laser treatment of the retina. Arch Ophthalmol 116(12):1633-1639, 1998 30. Berger TW: Thermal modeling of micropulsed diode laser retinal photocoagulation. Lasers Surg Med 20:409-415, 1997 31. Almeida-Lopes L, Rigau J, Amaro Zangaro R, GuidugliNeto J, Marques Jaeger MM: Comparison of the low level laser therapy effects on cultured human gingival fibroblasts proliferation using different irradiance and same fluence. Lasers Surg Med 29(2):179-184, 2001 32. Lagan KM, Clements BA, McDonough S, Baxter GD: Low intensity laser therapy (830 nm) in the management of minor post surgical wounds: a controlled clinical study. Lasers Surg Med 28(1):27-32, 2001 33. Guzzardella GA, Tigani D, Torricelli P, Fini M, Martini L, Morrone G et al: Low-power diode laser stimulation of surgical osteochondral defects: results after 24 weeks. Artif Cells Blood Substit Immobil Biotechnol 29(3):235-244, 2001 34. Stadler I, Lanzafame RJ, Evans R, Narayan V, Dailey B, Buehner N et al: 830-nm irradiation increases the wound tensile strength in a diabetic murine model. Lasers Surg Med 28(3):220-226, 2001 35. Gass JD: Drusen and disciform macular detachment and degeneration. Arch Ophthalmol 90(3):206-217, 1973 36. Little HL, Showman JM, Brown BW: A pilot randomized controlled study on the effect of laser photocoagulation of confluent soft macular drusen. Ophthalmology 104(4):623631, 1997 37. Guymer RH, Hageman GS, Bird AC: Influence of laser
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56. Cullia TA, Harris A, Kagemann L, Maturi R, McNulty L, Pratt LM, Xiao M, Crisswell MH, Weinberger D: Transpupillary thermotherapy for subfoveal occult choroidal neovascularization: effect on ocular perfusion. Invest Ophthalmol Vis Sci 42:3337-3340, 2001 57. Thompson JT: Retinal pigment epithelial tear after transpupillary thermotherapy for choroidal neovascularization. Am J Ophthalmol 131(5):662-664, 2001 58. Park CH, Duker JS, Mainster MA, Puliafito CA, Reichel, E: Transpupillary Thermotherapy (TTT) of occult choroidal neovascularization. A retrospective, noncomparative case series of 57 eyes. Seminars in Ophth 16(2):66-69, 2001 59. Whitacre MM, Timberlake GT, Stein RA, Stanley AM, Van Vleck S, Dominick KE: Light distribution of ocular endophotocoagulator probes and its surgical implications. Lasers Surg Med 15(1):62-73, 1994 60. Sliney DH, Marshall J: Tissue specific damage to the retinal pigment epithelium: mechanisms and therapeutic implications. Lasers Light Ophthalmol 5:17-28, 1992 61. Scott LJ, Goa KL: Verteporfin. Drugs Aging 16(2):139-146, discussion 147-148, 2000 62. Turnbull RG, Chen JC, Labow RS, Margaron P, Hsiang YN: Benzoporphyrin derivative monoacid ring A (Verteporfin) alone has no inhibitory effect on intimal hyperplasia: in vitro and in vivo results. J Invest Surg 13(3):153-159, 2000 63. Adili F, Statius Van Eps RG, Flotte TJ, LaMuraglia GM: Photodynamic therapy with local photosensitizer delivery inhibits experimental intimal hyperplasia. Lasers Surg Med 23(5):263-273, 1998 64. Haimovici R, Kramer M, Miller JW, Hasan T, Flotte TJ, Schomacker KT et al: Localization of lipoprotein-delivered benzoporphyrin derivative in the rabbit eye. Curr Eye Res 16(2):83-90, 1997 65. Schmidt-Erfurth U, Bauman W, Gragoudas E, Flotte TJ, Michaud NA, Birngruber R et al: Photodynamic therapy of experimental choroidal melanoma using lipoprotein-delivered benzoporphyrin. Ophthalmology 101(1):89-99, 1994 66. Levy JG: Photosensitizers in photodynamic therapy. Sem Oncol 21(6 Suppl 15):4-10, 1994 67. Belzacq AS, Jacotot E, Vieira HL, Mistro D, Granville DJ, Xie Z et al: Apoptosis induction by the photosensitizer verteporfin: identification of mitochondrial adenine nucleotide translocator as a critical target. Cancer Res 61(4):12601264, 2001 68. Sickenberg M: Early detection, diagnosis and management
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Photodynamic therapy: basic principles and mechanisms Hubert Van den Bergh and Jean-Pierre Ballini EPFL, Lausanne, Switzerland
Keywords: photodynamic therapy, photochemistry, photophysics, photobiology, selectivity of action
Abstract
Introduction
In this article, the authors describe the principles and mechanisms of photodynamic therapy (PDT). After an introduction on the PDT of cancer, the photophysics of photosensitizer molecules is described, followed by the elementary photochemical mechanisms and some of the PDT-induced biomolecular cascades. The authors then focus on the effects of PDT on the blood vessels and the treatment of choroidal neovascularization (CNV) of the retina associated with the exudative form of age-related macular degeneration (AMD). The selectivity of the PDT of CNV in AMD is discussed, and some possible improvements are proposed.
The history of photodynamic therapy (PDT) has recently been reviewed,1,2 and excellent overviews on the subject of PDT exist.3-5 PDT for cancer is presented by the schematic description shown in Figure 1, in which the human body is simplified as a three-compartment system. Figure 1 shows the PDT of early-stage superficial non-metastasized cancer in, for example, a hollow organ (lung or esophagus). After (1) injection of a photosensitizer (PS), (2) a certain amount of selective uptake and/or removal leads to the local preferential concentration of the PS in the tumor compared to the surrounding normal tissue. At the
Fig. 1. The principle of partial selectivity, which is fundamental to photodynamic therapy.
Address for correspondence: Dr. H. van den Bergh, Ecole Polytechnique Fédérale de Lausanne, ENAC/ISTE/LPAS, CH 1015 Lausanne, Switzerland. e-mail:
[email protected] Lasers in Ophthalmology – Basic, Diagnostic and Surgical Aspects, pp. 183–195 edited by F. Fankhauser and S. Kwasniewska © 2003 Kugler Publications, The Hague, The Netherlands
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Fig. 2. ‘Selective’ necrosis of an early squamous cell cancer in the bronchi.
same time, a high concentration of PS may be found in the liver, kidneys, or other organs, by which PS is removed from the body. We call this situation partial selectivity. It is important that PS is applied at concentrations at which, by itself, it is not toxic to any part of the body. PS is then activated (3) by applying light to the surface of the hollow organ. Generally, this is performed in such a way that it shields the superficial normal tissue surrounding the tumor. Light delivery systems for this purpose are described in detail elsewhere.6 The photosensitizer activated by light (PS*) then destroys the tissue in which it is located. (4) The volume of destroyed (mostly necrotic) tissue is largely determined by the area where the light is applied, by the limited penetration of light of a given wavelength into the tissue, and by the selectivity of PS concentration in the tumor compared to the surrounding normal tissues. Thus, the deeper lying tissue below the cancer should not be destroyed and, most importantly, the high concentration of PS in other organs, such as the liver, spleen, or kidneys, is of no consequence, since they are not reached by light. Exceptions to this are, of course, the skin and the eyes, which must be protected by the patient staying out of sunlight
(or strong artificial light) until most PS has been removed from these organs by natural pharmacokinetic processes. Finally, in time, the damaged surface will heal. It should also be underlined that, in themselves, the low light intensities used in PDT are not harmful to the body either. Very promising results have been obtained using PDT on early stage cancer.7,8 These demonstrate the efficacy and simplicity of this minimally-invasive and repeatable procedure. Furthermore, side-effects, such as skin photosensitivity, stenosis, and perforation of the hollow organs can essentially be completely avoided by using the newest PSs, light applicators, and PDT procedures. In particular, most new PSs are designed to be rapidly removed from the body. Figure 2 shows the ‘selective’ necrosis of an early squamous cell cancer in the bronchi about ten days after PDT. Tables 1 and 2 demonstrate the low rates of recurrence in this procedure after several years of follow-up using two different PSs for early squamous cell carcinomas (SCC) in the upper aerodigestive tract (UAT), the esophagus, and the bronchi. The data shown in Tables 1 and 2 show that, while ‘carcinoma in situ’ can essentially be treated with
Table 1. Results of PDT with HPD (hematoporphyrin derivative) and PhotofrinII in 51 patients with SCC Location
Patients with in situ SCC
Patients with micro-invasive SCC
Patients with no recurrence
Rate of recurrence %
UAT Esophagus Bronchi Total
3/3 7/8 7/8 17/19
4/7 9/13 8/12 21/32
7/10 16/21 15/20 38/51
30 24 25 25
(Reproduced from Radu et al.9 by courtesy of the publisher) Table 2. Results of PDT with mTHPC (meso-tetrahydroxy-phenyl-chlorin) in 50 patients with SCC Location
Patients with in situ SCC
Patients with micro-invasive SCC
Patients with no recurrence
Rate of recurrence %
UAT Esophagus Bronchi Total
3/4 9/10 11/12 23/26
2/2 12/18 2/4 16/24
5/6 21/28 13/16 39/50
17 25 19 22
(Reproduced from Radu et al.9 by courtesy of the publisher)
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Photodynamic therapy: basic principles and mechanisms 100% efficacy, recurrence increases as the staging of the cancer advances. This is probably in part due to an increase in distant metastatic disease, which cannot be treated easily due to the limited penetration of light into the tissue. In PDT, the maximum depths of necrosis that have been demonstrated range from several millimeters to about 1 cm, depending on the PS and the excitation wavelength; beyond that, PDT can still be quite effective using interstitial illumination. Photophysics Absorption of light by PS or other molecules is governed by the so-called Beer-Lambert law log
$% I0 I
= ε (cl)γ
(1)
where I0 is the light intensity at a given wavelength (λ) incident on a sample of thickness l in units of cm, which contains photosensitizer molecules at
rmolesi concentration cu u. ε is the decadic molar ex3 q cm t tinction coefficient, which gives the probability that light is absorbed by the PS at wavelength λ. The cm 2 units of ε in this example would be . I is the mole light intensity transmitted through the sample. γ which is often unity, is a factor that corrects for the nonlinearity of the system, which can be due, for instance, to the width of the exciting light being much broader than the spectral features of the PS. When a photon is absorbed by a PS it gains energy according to E = hν =
hc λ
where E is the photon energy (erg), h is Planck’s constant (erg.sec), ν is the frequency of the light (sec-1), c here is the velocity of light (cm sec-1) and λ is the wavelength in cm.
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If we measure the light absorption as a function of the wavelength for a given PS, the resulting curve of ε as a function of wavelength is called an absorption spectrum. This is shown in Figure 3 or a ‘typical’ porphyrin (protoporphyrin IX in this case), together with its chemical structure. The absorption of a photon by a PS means that the PS energy is increased by the energy of the photon E = hν. In other words, the spectrum in Figure 3 describes the relative probability of photon-induced energy level changes in the PS. The absorption in Figure 3 near 400 nm is called the Soret band, and the other four bands between 500 and 630 nm are called the Q bands. The probability of absorption of a photon by a PS depends upon several properties of the ground state and the state excited by the photon: i.e., on the electronic spin multiplicity, the overlap of the orbitals in space, the symmetry of the wave functions, and the magnitude of change of momentum. The electrons in a molecule are generally found in pairs in spatial volumes called orbitals 1 with either spin (s) equal to + (↑) or spin equal 2 1 to – (↓), and the multiplicity of an electronic con2 figuration is defined as 2S + 1 where S = Σ s. In a PS with an even total number of electrons, we have singlet states where 2S + 1 = 2 U$2 % + $– 2%Y + 1 = 1. 1
1
These states may be indicated by (↑↓). We also have triplet states where 2S + 1 = 2
U$ 12 % + $ 12%Y + 1 = 3
which are labelled by (↑↑). Thus, the ground state and the lower electronic states in a PS may typically have the configuration shown in Figure 4. Photon absorption from the ground singlet state S0 to the first excited singlet state S1 is a process with no change in multiplicity. This is called a spinallowed transition. Following this photon-induced transition to S1 different energy redistribution pathways may occur, as shown schematically in Figure 5, a so-called Jablonski diagram. The different electronic state are indicated as S0, S1, S2 or T1, and all have sublevels of energy rep-
Fig. 3. The absorption spectrum and structure of protoporphyrin IX, an intermediate in the synthesis of haem, or iron-protoporphyrin IX.
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Fig. 4. The multiplicity (2S + 1) of ground and excited states showing the configuration of some of the outer electrons of a PS.
Fig. 5. Simplified Jablonski diagram showing the energy levels and the energy flow pathways between these levels. Some vibrational sublevels are shown for each of the electronic levels, rotational levels are omitted for the sake of simplicity. The three major pathways for the S1 state are: (1) fluorescence, (2) internal conversion (heat), and (3) ISC or intersystem crossing.
resenting some excess rotational and vibrational energy. Following absorption of a photon on an extremely short time scale, about 10-15 seconds, the PS is in a vibrationally excited state of S1. This state then undergoes fast (~10-12 seconds) vibrational relaxation (VR) to give the S1 de–excited singlet state. From this ‘ground’ S1 state five different processes can basically occur, depending on the PS: 1. Fluorescence, generally on a time scale of 10-8 to 10-9 seconds. Emission of a fluorescent photon is to one of several rovibrational states of the S0 ground electronic state. Averaged over many emitted photons, this gives a broad-banded emission, which is somewhat red shifted from the absorption wavelength. This is the so-called Stokes shift. 2. Internal conversion (IC). This is a non-radiative change between states of the same multiplicity. In Figure 5, the IC process between S1 and vibrationally excited S0 is followed by the loss of vibrational excess energy colliding with the molecules surrounding the PS. Thus, the photon energy is rapidly degraded into heat.
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3. Intersystem crossing (ISC). This is also a nonradiative change between electronic states, however, of a different multiplicity. The ISC process indicated in Figure 5 is a non-radiative transfer between S1 and T1. Due to the energy difference (so-called singlet-triplet splitting), T1 is produced in a vibrationally excited state which once again undergoes rapid VR to give the metastable triplet vibrational ‘ground’ state T1. This state can then either undergo a chemical reaction, transfer its energy to another molecule, or internally convert to S0, or radiate in a non-allowed spin-forbidden transition to one of several rovibrational states of the S0 ground electronic state. The latter, since it is a spin-forbidden process, is quite slow and is called phosphorescence. 4. S1 can react chemically in many different ways: for instance, it can rearrange itself, it can add itself to another molecule, or dissociate itself. 5. Energy exchange with a neighboring molecule, which is then excited to its singlet, excited state. Different PS are selected or synthesized for their
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Fig. 6. Schematic representation of some of the photochemical steps thought to be involved in the sensitization of a biological substrate molecule M by a photosensitizer activated to its triplet state (3PS).
different properties, depending on the application. For instance, fluorescein and indocyanine green (ICG) are used for fluorescence angiography in ophthalmology. They are chosen in particular for having a high fluorescence quantum yield. Most PSs used in PDT are chosen for their high yield of long-lived triplet states. The oxygen present in the tissue can diffuse to the triplet (T1) state and pick up its energy by colliding with it. The excited oxygen can then diffuse away from the PS and oxidize certain types of the surrounding molecules (or get quenched). In a given PS, the preferential pathway can be modified by chemical changes being made. For instance, substituting hydrogen atoms by heavy halogen atoms or oxygen by sulphur in the PS macrocycle gives higher triplet yields. This is the case, for instance, when going from fluorescein (a predominantly fluorescing molecule) to tetrachlorotetraiodofluorescein or Bengal rose, which is a good photosensitizer with a high triplet yield. Adding a central metal atom or, for instance, changing it from diamagnetic to paramagnetic can also change the properties of the PS significantly. Thus, protoporphyrin IX is a good PS which fluoresces rather well, while upon insertion of Fe2+ into the center of the macrocycle, both these phenomena are strongly suppressed, indicating a change to internal conversion as the predominant pathway. Photochemistry In the previous section, the photophysics of photosensitizers (PS) showed that short-lived singlet states (S1) and longer-lived metastable triplet states (T1) are produced following photon absorption. Direct reactions of substrate molecules ‘M’ with the short-
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lived S1 state of the PS are probably not important, and even if S 1 collides with oxygen, which may induce intersystem crossing (ISC) to give the triplet state T1, singlet oxygen would probably not be produced as this process is highly endothermic. This is because the singlet-triplet splitting in porphyrins is frequently of the order of 8 kcal/mole compared to the 1O2 excitation energy of nearly 23 kcal/mole. The chemical reactions that are induced by the triplet state photosensitizer (T1) are summarized in Figure 6 (see also Jori and Spikes10). The chemical reactions shown in Figure 6 follow three main pathways, which start from the common metastable triplet state of the photosensitizer (3PS): (1) energy transfer, either to oxygen to give O2(1∆) which then either reacts with PS (bleaching it in a ‘cage’ reaction) or oxidizes a biomolecular target close to the PS. This is generally called a Type II reaction. Alternatively, the 3PS can transfer its energy to M. The second pathway: (2) hydrogen atom exchange between the PS and M leads to radical intermediates which may combine with oxygen to form peroxide radicals as subsequent intermediates leading to oxidized M. Finally, (3) electron transfer between 3PS and M leads to radical ions which together with O2 can lead to other oxidizing reactive intermediates, such as the superoxide radical anion (O2•-), H2O2 and the very reactive hydroxyl (OH•) radical. Processes (2) and (3) together are called Type I reactions. It should be noted that Type II reactions require oxygen in the first step, while Type I reactions only involve oxygen further downstream in the reaction mechanism. From the three pathways indicated in Figure 6, the energy transfer from 3PS to oxygen is generally envisaged to be one of the more important steps. This is shown in more detail in Figure 7 (see also, Turro11).
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1 ∆
2 3 Fig. 7. The triplet-triplet annihilation mechanism for energy transfer from triplet PS to 3O2 may be described as (1) electron exchange, (2) electron transfer, and (3) diradical formation.
Fig. 8. Two typical reactions of 1O2.
Finally, it should be noted that the energy transfer to M (3PS + 0M 0PS + 3M) is too endothermic for most biomolecules, β-carotene being a typical exception. In a biological environment, singlet oxygen (1O2) either reacts rapidly or is quenched so that reactions take place at distances of only a few tens of nanometers from the location of the PS, so that PS localization determines where PDT damage takes place. Some typical reactions of 1O2 are shown in Figure 8. Biomolecular pathway changes In vivo, PDT effects are interpreted in terms of three mechanisms: cellular, vascular, and immunological effects. The relative contribution of each of these to the overall PDT efficacy in destroying a tumor
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depends on the tissue, the sensitizer, and the conditions applied. The latter include the time between drug and light application and the drug delivery system, both of which determine drug localization at the time of light application. The quantities of drug and light also play a role. In this overview, we are particularly concerned with the vascular PDT effect,1,2,5,12-16 which may be observed as a vasoconstriction directly upon light exposure or several hours later. Endothelial cells are supposedly the main target and, depending upon the drug, either mainly the mitochondria (this is the case for BPD-MA) or predominantly the lysosomes or the plasma membrane may be targeted, among others. Lysosomal localizing (water soluble) PSs tend to cause cell death by necrosis, or can lead to apoptosis via the release of cathepsins and caspase 3. On the other hand, mitochondrial localizing PSs mainly lead to cell death by apoptosis. In these organelles, PDT-induced lipid
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Fig. 9. PDT-induced biochemical cascades leading to apoptosis. Initiation depends on the localization of the photosensitizer.
peroxidation may lead, among other things, to organelle membrane disruption, change in membrane potential, or damage to membrane proteins. Apart from apoptosis and necrosis as a death response to PDT, partially damaged cells will show an activated rescue response in the form of release of heat shock proteins, glucose-regulated proteins and heme oxygenase (see, Xue et al.,17,18 Gomer et al.19-21 and Morgan et al.22). PDT also influences surface receptors and induces cytokines, which probably influence the immune response. For an example of some of the known details of the biochemical apoptotic pathways induced by PDT, the reader is referred to Figure 9 and to Oleinick et al.23 and Roth and Reed.24 Damage by PDT may occur either mainly in the cell membrane, the mitochondria, or the lysosomes. We take the case of a PS (BPD-MA) mainly localized in the mitochondria at the time of light application, in some cases causing the loss of the mitochondrial membrane potential, possibly due to the opening of a large conductance channel called the mitochondrial permeability transition pore complex (PTPC). This can contribute to the release of cytochrome-c into the cytosol, where it forms a multi-protein complex called an apoptosome with apoptosis-activating factor-1 (APAF-1) and ATP. The latter recruits and activates pro-caspase 9 via the caspase dimer. The name caspase comes from a cysteine protease acting on aspartic acid, i.e., a protein that cleaves another protein at a specific site. The apoptosome now releases caspase 9 which, in turn, induces caspase 3 release. The latter is a key protein in apoptosis induction and causes the cleavage of multiple proteins, among which PARP (poly-ADP-ribose polymerase) and DNA-PK, which normally acts to repair DNA damage. Caspase 3 also attacks ICAD, thus releasing CAD which itself attacks DNA. Caspase 3 can also
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activate pro-caspase 6, thereby releasing caspase 6 which cleaves the lamins of the cell nucleus, thus inducing nuclear breakdown. Many other substrates are attacked by caspase 3, among which SHREBs, Gelsolin, caspase 7, caspase 9, MDM2, GAS2, FODRIN, FAK, etc., all of which are involved in apoptosis. In a parallel process, the release of cytochrome-c can also indirectly influence the activity of all these caspases by the release of Smac/Diablo,2428 which interferes with the release of IAPs (inhibitor of apoptosis proteins). The latter controls the release of all these caspases. For water-soluble PSs that might localize in lysosomes, PDT leads to the release of cathepsins which, among other things, help to release caspase 3. For the sake of simplicity, this has not been added to Figure 9. It cannot be completely ruled out that, in the case of a vascular PDT effect following damage to endothelial cells at very short intervals after PS injection, cell surface receptors play a role. Activation of cell death receptors such as TRAIL (TNF-related apoptosis-inducing ligand), FAS, and TNFR1 (tumor necrosis factor receptor 1) via specific ligands will cause adaptor proteins such as FADD (FAS associated death domain) to bind to their cytosolic component. This complex then recruits and activates procaspase 8, causing caspase 8 release. The latter acts to give caspase 3. This picture is further complicated by the fact that proteins of the Bcl-2 family regulate apoptosis and contain a number of both pro-apoptotic members such as Bid and Bax (the latter may be able to form a transmembrane pore, leading to cytochrome-c release) as well as many anti-apoptotic members, among which Bcl-2 itself, as well as Bcl-xL which are located in the outer mitochondrial membrane and promote cell survival. Caspase 8 cleaves Bid in the cytosol, generating truncated Bid (t-Bid), which then relocates to the
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mitochondrial membrane, thus helping to release cytochrome-c. It is hoped that a better understanding of these pathways will lead to more effective PDT via apoptosis, and hence possibly lower drug and light doses compared to necrosis. It may also be speculated that the selectivity of apoptosis induced by PDT might be easier to manipulate than that of necrotic cell death, so that improved selective destruction of certain tissues can be envisaged. More information on caspases can be found in Salvesen and Dixit,29 Cohen,30 Qin et al.,31 and Thornberry and Lazebnik.32 Vascular effects of photodynamic therapy Photodynamic therapy with various photosensitizers under a specific set of applied conditions can have a predominant vascular component. This has led to the use of PDT for treating several non-cancerous lesions such as atherosclerotic plaque, choroidal neovascularization (CNV) of the retina, re-stenosis after balloon angioplasty, and port-wine stains. Microvascular damage and blood flow stasis during and after PDT have consequently been the focus of a number of recent studies.1,2,12-15,33-35 Although many details still remain unknown, a certain picture illustrated by Figure 10, can be extracted from the available information. Early damage, after the i.v. injection of PS and illumination, is observed in endothelial cells and the sub-endothelium. In particular, PDT-induced changes to endothelial cells can be found on the luminal surface, in cytoplasmic microtubules, on cytoskeletal proteins, and in mitochondria. Endothelial cells undergoing PDT then retract and lose their tight junctions with adjacent cells, thus exposing the basement membrane as shown in Figure 10b. The latter leads to the activation of platelets and polymorphonuclear leukocytes which, in turn, causes the increased release of eicosanoids, and finally the aggregation of platelets on the exposed basement membrane of the vessel wall. A further observation following PDT is the adhesion of polymorphonuclear leukocytes to the vessel
wall. A consequence of these changes is the release of biochemicals, including the eicosanoids thromboxane and leukotriene B4 and C4. It should be noted that the rounding of the endothelial cells combined with the biochemical changes first causes an increase in vessel permeability and leakage, before finally leading to blood flow stasis, as illustrated by Figure 10c. The PDT-induced release of biochemicals causes disturbance of the normal equilibrium between aggregating and disaggregating behavior, as well as between vasoconstriction and vasodilatation processes. This imbalance results in increased smooth muscle cell activity and effective vasoconstriction, combined with platelet aggregation to form a plug, which is stabilized by fibrins. In a parallel mechanism, PDT damage to membrane lipids can cause the release of arachidonic acid and a consequent biochemical cascade involving cyclooxygenase, prostaglandin endoperoxides, and finally also the production of the vasoconstrictor, thromboxane. The overall result of all this is blood flow stasis, which may to some extent have been aided by the increased interstitial pressure following enhanced leakage. In order to investigate the effects of PDT on the blood flow in more detail, PDT-induced leakage and vessel constriction were studied in a CAM model (chicken embryo’s chorioallantoic membrane) in our laboratory. Some results obtained with this model using various PSs are shown below. In Figure 11, typical experimental results compare photosensitizer leakage from CNV in a human eye (on the left) with leakage from small CAM vessels (on the right). The leakage of PS is important in the PDT of CNV associated with age-related macular degeneration (AMD), since too much leakage of the PS before light application may result in photodynamic damage to parts of the retina close to the leaky neovessels that we want to preserve (such as, for instance, the photoreceptors). The two images on the left-hand side of Figure 11 show the strong localized leakage of fluorescein at short and at longer times after injection. This leakage is typical of aggressive ‘classical’ CNV, which in this case was later treated by PDT with
Fig. 10. A simplified mechanism of PDT leading to blood flow stasis.
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Fig. 11. Photosensitizer leakage: a comparison between a human eye on the left and a CAM model on the right. The intervals after injection at which the fluorescence angiograms were taken increases from top to bottom.
Fig. 12. PS efficacy in blood flow stasis. A comparison between the human eye one week after PDT, and the CAM model vasculature at different times after PDT. The drug used was Visudyne®.
Visudyne®. On the right-hand side, in the CAM model, the fluorescence pharmacokinetics is shown for a water-soluble dye (Rhodamine 101) after i.v. injection,36 demonstrating the arrival of the fluorescing substance in different parts of the vasculature at short intervals after injection (top), and the leakage of the substance at longer intervals after injection (bottom). The basic idea behind this experiment is to establish a ‘leakage-scale-relationship’ between the human eye and the CAM model. The efficacy of a PS in blood flow stasis is shown in Figure 12, once again comparing this property between the human eye (on the left) and the CAM vessels (on the right). The goal also being to screen new PSs for PDT of CNV in AMD and to establish a ‘PDT-blood-flow-stasis’ relationship between the clinical observations and the results in the preclinical model. Thus, on the left-hand side of Figure 12,
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the leakage of fluorescein is shown in one patient, one week after PDT treatment with Visudyne®. The dark hypofluorescent spot that can be seen in the macular region shows, one week after PDT, both the disappearance of the CNV leakage that existed prior to PDT, and the decreased leakage of the partially closed choriocapillaries in this region. At the right top of Figure 12, the time frame of PDT with Visudyne® on CAM 36 is shown by Visudyne® angiograms prior to PDT (40 sec), while light application (2 min) shows the treated area. At the right bottom of Figure 12, the blood flow stasis is shown 24 hours after PDT by a Rhodamine 101 angiogram. A hypofluorescent spot, quite similar to that shown in the clinical contest, indicative of blood flow stasis in the irradiated region, does not show immediately after PDT, but becomes clearly visible 24 hours after treatment.
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Selectivity in photodynamic therapy ‘Selectivity’ in the PDT of tumors will depend on the drug, drug carrier, application of light, and tissue properties. Selectivity may be defined in many ways. Here, we present some possible reasons for ‘drug concentration-selectivity’ in PDT: Selective uptake Leaky neovessels in tumors can cause more of the drug to be delivered. This leakiness may be temporarily enhanced by PDT itself. Tumors may also have increased the activity of low density lipoprotein (LDL) receptors and/or other receptors, such as albumin. Hydrophobic sensitizers may be bound in the lipid core of lipoproteins, while hydrophilic PSs can be transported by albumin. Selective retention Tumor-associated macrophages may accumulate both aggregated lipophilic PSs and lipoproteins overloaded with PSs. pH-lowering advanced cancers can cause a decrease in the solubility of certain porphyrins. Reduced lymphatic drainage in neoplastic tissue may cause a build-up of certain molecules. Furthermore, a PS may be targeted to some receptor, which is selectively expressed in a tumor or in tumor neovasculature. This may be done by attaching a PS covalently to a monoclonal antibody or antibody fragment, which is targeted to a specific antigen. Alternatively, the PS could be attached to a small peptide targeted to an integrin such as αvβ3, which has been shown to be expressed selectively on some tumor neovasculature. The case of selectivity in the PDT of CNV associated with exudative AMD is discussed in detail by Birngruber in another chapter in this volume. Here, this topic is discussed briefly from a slightly different point of view. Figure 12 showed a fluorescein angiogram of the macular region of a human eye after PDT. The irradiated zone shows up as a dark hypofluorescent region. This implies that, not only have the exudative neovessels been closed, but also the choriocapillaries have at least partially been closed. The retinal capillaries are clearly patent. Hence, the treatment is perfectly selective in the sense of sparing the retinal capillaries – which is all-important – and not completely selective in the sense of partially closing the choriocapillaries. In the following section, we discuss the selectivity issues that apparently play a role in the PDT of CNV associated with AMD. Figure 13 shows a schematic and simplified diagram of the central region of the retina, known as the macula. The light is shown coming in from the top. This activates the photoreceptors (there are mostly cones for high resolution and color vision in the macula). The signal from the cones is pre-
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processed by bipolar and horizontal cells, amacrine cells, and ganglion cells (none shown) before being transferred to the brain by axons, which leave the eye in a bundle via the optic nerve. Below the photoreceptors is the retinal pigment epithelium (RPE), a layer of ‘macrophage-like’, highly pigmented cells. There are five types of blood vessels in the diagram. Near the surface (top) of the retina are the retinal arteries and veins, which feed the small and not very dense network of retinal capillaries, which surrounds the macula. In the macular region itself, there are no retinal vessels, as these would tend to blur the high-resolution image. Closure of the retinal capillaries in PDT must absolutely be avoided, since this is extremely damaging to the visual acuity. Below the RPE there is a zone called Bruch’s membrane, and below this are the choroidal capillaries, which are fed by the larger choroidal vessels. Although the etiology of exudative (neovascular) AMD is not completely understood, one reason for the growth of CNV from the choriocapillaries is probably a lack of oxygen and other nutrients at the level of the photoreceptors. The leaking CNV end up by disturbing visual acuity, as has been described in detail in the literature.37 The following question may now be asked: what selectivity is required from PDT? First, as described above, closure of the retinal capillaries must be avoided. This is probably facilitated by the lower partial pressure of oxygen in the retinal circulation compared to the choroidal circulation,38-41 the former being part of the blood-brain barrier associated vasculature. In fact, the human retina receives its oxygen from two separate sources: the retinal circulation is the main supply of O2 for the inner retina between the vitreal surface and the inner nuclear layer, while the choroidal circulation oxygenates the remainder of the retina and the RPE (see Fig. 13). Furthermore, oxygen consumption is not homogeneous across the retina. It is probably highest near the photoreceptor inner segments (IS in Fig. 13). Therefore, the non-uniformity of the oxygen supply and oxygen consumption in the retina must lead to oxygen gradients.40,42-45 On an even smaller scale, these gradients may be further disturbed by the differences in oxygen solubility in different tissue compartments. Thus, O2 is less soluble in the cytosol than in the plasma membrane. Tighter contacts between endothelial cells in the retinal circulation may also render PDT relatively less effective. A second type of selectivity is that PDT should, at least to some extent, be restricted to the vasculature. This means that PSs should not be allowed to leak significantly out of the CNV on the timescale of the treatment as it is important to avoid damage to the photoreceptors and the neural retina, as well as major damage to the RPE. The RPE cells probably pick up leaked PSs rapidly. And even though some damage to the RPE (which might lead to replacement of ‘aged’ RPE cells) might be of interest, major damage to the RPE should probably
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Fig. 13. A simplified schematic diagram of the central part of the retina. The fovea (not indicated) can be found in the middle of the macula, and is the zone for the high visual acuity of the human eye. Note that: (1) no major leakage of PS from the vessels is desired prior to PDT, and (2) CNV must be ‘closed’ to minimalize damage to the normal choriocapillaries and retinal capillaries. Oxygen concentration is lowest near the inner segments (IS) of the photoreceptors, higher near the vitreous, and highest at the choriocapillaries. (Adapted from Van den Bergh1 by courtesy of the publisher.)
Fig. 14. BPD-MA angiograms recorded with a modified Topcon fundus camera, showing the localization of the PS in the CNV area. Plot of the analyzed BPD-MA fluorescence versus time, up to 30 minutes, in main retinal (line) and choroidal (dotted line) vessels, respectively, for comparison with the concentration of PS measured in the plasma.
be avoided. A third type of selectivity implies that we want to close the CNV without destroying the choriocapillaries or larger choroidal vessels. The larger choroidal vessels may be protected by a significantly different amount of collagen in the vessel walls compared to the smaller diameter vessels. It has been pointed out by Hasan et al.2 that neovasculature in some diseases tends to have enhanced activity of receptors for LDL and albumin. Thus, lipophilic substances such as BPD-MA attached to LDL may be preferentially taken up in the CNV. Some proof of enhanced BPD-MA in the neovascular region can be seen in Figure 14, which shows the fluorescence pharmacokinetics of BPD-MA in the eye fundus at intervals up to 30 minutes.
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Analysis of these images46 indicates a possible degree of selectivity of BPD-MA for the CNV. Nevertheless, PDT also leads to at least temporary closure of a significant part of the choriocapillaries, as can be seen from the hypofluorescent spot in the fluorescence angiograms obtained at early and late time intervals with both ICG and fluorescein (Fig. 12). The hypofluorescence observed with both compounds indicates a significant closure of the choriocapillaries one week after PDT. Thus, the desired selectivity between closing CNV and not closing choriocapillaries is incomplete. In order to attain this selectivity, different approaches are taken, two of which involve the increased expression of αvβ3 integrins at the surface of neovascular endothelial cells.
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Birchler et al.47 from Zurich have tried to link a PS to an antibody fragment that attaches to fibronectin, which itself attaches to αvβ3. Other groups, including ourselves, are presently trying to attach PS to RGD-peptide-related structures which attach directly to αvβ3. Only the future will show whether improved PDT selectivity can be attained with such approaches, and whether this selectivity will significantly influence the treatment efficacy. Non-closure of the choriocapillaries should increase tissue oxygenation after PDT, and decrease PDT-related inflammation, both of which should reduce the recanalization/regrowth of CNV that hampers present-day treatments and makes repeat treatments necessary. Finally, of course, it should be mentioned that, with PDT of CNV in AMD, we treat the effects of the disease but not its cause(s). Conclusions In this overview, we have described some of the basic photophysics, photochemistry, and photobiology associated with PDT. A simplified mechanism of the biological cascades leading to apoptosis has also been described, as well as details of PDT in vessels leading to blood flow stasis. Finally, selectivity issues in the PDT of exudative AMD are discussed. Acknowledgment We would like to thank N. Lange for the preparation of Figures 11 and 12.
References 1. Van den Bergh H: Photodynamic therapy of age related macular degeneration: history and principles. Semin Ophthalmol 25: 2002 (accepted for publication) 2. Hasan T et al: Photodynamic therapy of cancer. In: Holland JF et al (eds) Cancer Medicine, pp 489-502. Hamilton, BC: Decker Inc 2000 3. Milgrom LR: The Colours of Life: An Introduction to the Chemistry of Porphyrins and Related Compounds. New York, NY: Oxford University Press Inc 1997 4. Bonnett R: Chemical Aspects of Photodynamic Therapy. New York, NY: Gordon and Breach Sci Publ 2000 5. Henderson BW, Dougherty TJ: Photodynamic Therapy: Basic Principles and Clinical Applications. New York, NY: M Dekker 1992 6. Van den Bergh H: On the evolution of some endoscopic light delivery systems for photodynamic therapy. Endoscopy 30:392-407, 1998 7. Monnier P et al: Photodetection and photodynamic therapy of ‘early’ squamous cell carcinomas of the pharynx, oesophagus and tracheo-bronchial tree. Lasers Med Sci 5:149-169, 1990 8. Monnier P et al: Further appraisal of PDI and PDT of early squamous cell carcinomas of the pharynx, oesophagus and bronchi. In: Spinelli P et al (eds) Photodynamic therapy and biomedical lasers, pp 7-14. Amsterdam: Elsevier Sci Publ 1992
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9. Radu A et al: Photodynamic therapy for 101 early cancers of the upper aerodigestive tract, the esophagus, and the bronchi: a single-institution experience. Diagnost Therapeut Endoscop 5:145-154, 1999 10. Jori G, Spikes D: Topics in photomedicine. In: SmithKendric C (ed) Topics in Photomedicine, pp 183-318. Cambridge: Perseus Publ 1984 11. Turro NJ: Modern Molecular Photochemistry. University Science Books, Sausalito, CA: 1991 12. Fingar VH et al: The role of microvascular damage in photodynamic therapy: the effect of treatment on vessel constriction, permeability, and leukocyte adhesion. Cancer Res 52:4914-4921, 1992 13. Fingar VH: Vascular effects of photodynamic therapy. J Clin Laser Med Surg 14:323-328, 1996. 14. Fingar VH et al: Analysis of acute vascular damage after photodynamic therapy using benzoporphyrin derivative (BPD). Br J Cancer 79:1702-1708, 1999. 15. Krammer B: Vascular effects of photodynamic therapy. Anticancer Res 21:4271-4277, 2001 16. Star WM et al: Destructive effect of photoradiation on the microcirculation of a rat mammary tumor growing in ‘sandwich’ observation chambers. Prog Clin Biol Res 170:637645, 1984 17. Xue LY et al: Elevation of GRP-78 and loss of HSP-70 following photodynamic treatment of V79 cells: sensitization by nigericin. Photochem Photobiol 62:135-143, 1995 18. Xue LY et al: Rapid tyrosine phosphorylation of HS1 in the response of mouse lymphoma L5178Y-R cells to photodynamic treatment sensitized by the phthalocyanine Pc 4. Photochem Photobiol 66:105-113, 1997 19. Gomer CJ et al: Photodynamic therapy-mediated oxidative stress can induce expression of heat shock proteins. Cancer Res 56:2355-2360, 1996 20. Gomer CJ et al: Glucose regulated protein induction and cellular resistance to oxidative stress mediated by porphyrin photosensitization. Cancer Res 51:6574-6579, 1991 21. Gomer CJ et al: Increased transcription and translation of heme oxygenase in Chinese hamster fibroblasts following photodynamic stress or Photofrin II incubation. Photochem Photobiol 53:275-279, 1991 22. Morgan J et al: GRP78 induction by calcium ionophore potentiates photodynamic therapy using the mitochondrial targeting dye victoria blue BO. Photochem Photobiol 67:155164, 1998 23. Oleinick NL et al: The role of apoptosis in response to photodynamic therapy: what, where, why, and how. Photochem Photobiol Sci 1:1-21, 2002 24. Roth W, Reed JC: Apoptosis and cancer: when BAX is TRAILing away. Nat Med 8:216-218, 2002 25. Reed JC: Apoptosis-based therapies. Nature Rev Drug Disc 1:111-121, 2002 26. Nicholson DW: From bench to clinic with apoptosis-based therapeutic agents. Nature 407:810-816, 2000 27. LeBlanc H et al: Tumor-cell resistance to death receptorinduced apoptosis through mutational inactivation of the proapoptotic Bcl-2 homolog Bax. Nat Med 8:274-281, 2002 28. Deng Y et al: TRAIL-induced apoptosis requires Bax-dependent mitochondrial release of Smac/DIABLO. Genes Dev 16:33-45, 2002 29. Salvesen GS, Dixit VM: Caspases: intracellular signaling by proteolysis. Cell 91:443-446, 1997 30. Cohen GM: Caspases: the executioners of apoptosis. Biochem J 326:1-16, 1997 31. Qin H et al: Structural basis of procaspase-9 recruitment by the apoptotic protease-activating factor 1. Nature 399:549557, 1999 32. Thornberry NA, Lazebnik Y: Caspases: enemies within. Science 281:1312-1316, 1998
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Photodynamic therapy: basic principles and mechanisms 33. Star WM et al: Destruction of rat mammary tumor and normal tissue microcirculation by hematoporphyrin derivative photoradiation observed in vivo in sandwich observation chambers. Cancer Res 46:2532-2540, 1986 34. Henderson BW et al: Effects of photodynamic treatment of platelets or endothelial cells in vitro on platelet aggregation. Photochem Photobiol 56:513-521, 1992 35. Henderson BW, Dougherty TJ: How does photodynamic therapy work? Photochem Photobiol 55:145-157, 1992 36. Lange N et al: A new drug-screening procedure for photosensitizing agents used in photodynamic therapy for CNV. Invest Ophthalmol Vis Sci 42:38-46, 2001 37. Schmidt-Erfurth U et al: Photodynamic therapy of subfoveal choroidal neovascularization: clinical and angiographic examples. Graefe’s Arch Clin Exp Ophthalmol 236:365-374, 1998 38. Yu DY, Cringle SJ: Oxygen distribution and consumption within the retina in vascularised and avascular retinas and in animal models of retinal disease. Prog Retin Eye Res 20:175-208, 2001 39. Tornquist P, Alm A: Retinal and choroidal contribution to retinal metabolism in vivo: a study in pigs. Acta Physiol Scand 106:351-357, 1979 40. Tsacopoulos M et al: Studies on retinal oxygenation. In:
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195 Grote J et al (eds) Oxygen Transport to Tissue II, pp 413416. New York, NY: Plenum Publ Corp 1976 Pournaras CJ: Retinal oxygen distribution: its role in the physiopathology of vasoproliferative microangiopathies. Retina 15:332-347, 1995 Drujan BD, Svaetichin G: Characterization of different classes of isolated retinal cells. Vision Res 12:1777-1784, 1972 Linsenmeier RA: Effects of light and darkness on oxygen distribution and consumption in the cat retina. J Gen Physiol 88:521-542, 1986 Dollery CT et al: Oxygen supply to the retina from the retinal and choroidal circulations at normal and increased arterial oxygen tensions. Invest Ophthalmol 8:588-594, 1969 Alder VA et al: The retinal oxygen profile in cats. Invest Ophthalmol Vis Sci 24:30-36, 1983 Sickenberg M et al: A computer-based method to quantify the classic pattern of choroidal neovascularization in order to monitor photodynamic therapy. Graefe’s Arch Clin Exp Ophthalmol 237:353-360, 1999 Birchler M et al: Selective targeting and photocoagulation of ocular angiogenesis mediated by a phage-derived human antibody fragment. Nat Biotechnol 17:984-988, 1999
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The concept and experimental validation of photodynamic therapy in neovascular structures in the eye Reginald Birngruber Medical Laser Center Lübeck, Lübeck, Germany
Keywords: photodynamic therapy, age related macula degeneration, vascular occlusion, Verteporfin, choroidal neovascular membrane, photothrombosis
Introduction One of the essential advantages of the principle of photodynamic therapy is the possibility of using photosensitizer molecules which preferentially get attached to or incorporated into the desired target structure, in order to achieve optimal selectivity of the photodynamic action after light activation.1 Several targeting mechanisms and sensitizer distribution modalities are presently being evaluated for a variety of different PDT applications. In order to treat neovascular structures, a specific targeting process leading to the preferential uptake of the photosensitizer in new vessels has to be developed. An ultimate selectivity seems to be particularly necessary if the neovascularization is situated directly adjacent to functionally important structures, such as in the clinically important pathology of macular choroidal neovasculature (CNV). This chapter describes the concept of selective targeting proliferating vascular endothelial cells and reviews the experimental studies that eventually led to a new treatment modality for CNV-related macular diseases, such as age-related macular degeneration, pathological myopia, and other chorioretinal diseases. Ocular PDT The principle of PDT was first applied in ophthalmology in the 1980s by Sery,2 Gomer et al.,3 and Murphree et al.4 in order to investigate the potential for treating intraocular tumors. In all those investigations, hematoporphyrin derivate (HPD), the first commercially-available photosensitizer, was used. However, the selectively and efficiency of tumor
destruction, as well as the debridement capacity of the eye, were not high enough to lead to a successful new treatment modality. In the late 1980s, a number of papers were published that evaluated the possibilities of treating ocular neovascularizations photodynamically.5-9 Different photoactive substances, such as dehematoporphyrin ether, phthalocyanin and rose bengal were used with HPD. All these investigations showed thrombosis due to photodynamically-induced intraluminal action in normal retinal vessels,7 iris neovasculature,5 and experimentally-induced CNV in monkeys.6,8,9 These investigations demonstrated different degrees of efficacy and selectivity ranging from ‘PDT-augmented laser coagulation’6 to ‘closure of CNV with retinal preservation’,8 depending on the type of photosensitizers treatment protocol and dosimetry used in the experiments. Vascular selectivity of photodynamic therapy The rational approach for optimizing vascular PDT is to systematically search for the right kind of sensitizer with maximum affinity to vascular structures, investigate the optimal drug-light interval with the highest concentration gradient of the photosensitizer in the neovascular structures, and evaluate the appropriate drug and light dosimetry, avoiding systemic side-effects by applying the minimum drug dose necessary, and eliminating undesired unspecific thermal collateral damage by using low light irradiances for the photodynamic drug excitation. Therefore, the optimal compound would be a highly phototoxic sensitizer which selectively binds to vascular structures. In general, it is known that hydrophillic molecules have the tendency to be preferentially taken
Address for correspondence: Professor R. Birngruber, PhD, MD, Medical Laser Center Lübeck, Peter Monnik Weg 4, D-23562 Lübeck, Germany. e-mail:
[email protected] Lasers in Ophthalmology – Basic, Diagnostic and Surgical Aspects, pp. 197–204 edited by F. Fankhauser and S. Kwasniewska © 2003 Kugler Publications, The Hague, The Netherlands
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up by the vasculature. Moreover, numerous experimental studies have shown that damage to the endothelial cell layer of the inner vessel wall – be this mechanical10 or thermal due to diathermia11 or argonlaser irradiation12 – leads to intravascular thrombosis due to thrombocyte agglutination with associated platelet and erythrocyte clumping.13,14 Therefore, the concept of selective vascular photodynamic therapy was to maximize photosensitizer uptake in the vascular endothelial cells. This specific targeting of endothelial cells can be mastered by complexing a potent photosensitizer with a carrier molecule with a high affinity to endothelial cells. Serum lipoproteins are promising candidates for such carrier systems, with their receptor-mediated internalization into endothelial cells.15 The therapeutic principle of neovascular selectivity The clinical goal of neovascular photodynamic therapy is to treat neovascular diseases – most importantly CNV membranes and other chorioretinal vascular neoplasias such as choroidal and retinal hemangiomas – by a mechanism of action ultimately selective to the neovascular vessel wall. Consequently, the concept of selective vascular targeting could be even more specifically adapted to neovasculature if the carrier system were to be preferentially taken up by neovascular endothelial cells. Studies on tumor selectivity in photodynamic therapy have indicated that the uptake of lipophilic sensitizers in tumorous neovasculature is mediated by low density lipoprotein (LDL).16 Based on this argumentation, we hypothesized for the first time on the possibility of selective photothrombosis of the ocular neovasculature, using a sensitizer LDL complex for the selective targeting of proliferating neovascular endothelial cells.17,18,43 The desired selectivity is thought to be due to the increased LDL receptor expression in proliferating cells.44 In addition, improved efficacy might be achieved by the fact that the sensitizer LDL complex is internalized by endocytosis, causing enzymatically enhanced endothelial cell destruction after photoexcitation.19 The concept of selective neovascular photothrombosis is outlined in Figures 1a-g: a CNV membrane originating from the choriocapillaris proliferates through Bruch’s membrane into the
subretinal and retinal spaces (Fig. 1b). The new vessels (Fig. 1c), with their heavily leaking vessel walls, cause exudation and hemorrhages into the outer retinal space, eventually leading to the formation of a fibrovascular scar associated with a substantial loss of neural tissue (Fig. 1d). By applying a neovascular selective photosensitizer, the proliferating endothelial cells of the neovascular structures will be targeted (Fig. 1e), and after light application, the activated sensitizer will produce enough toxic singlet oxygen to selectively damaging the neovascular structure (Fig. 1f). The reactive intravascular thrombosis will then occlude the new vascular tissue without damaging the neural retina and retinal pigment epithelium, which should be able to re-establish the anatomical outer retinal structure (Fig. 1g). The assumed specific binding of the photosensitizer to LDL in the intravascular system is illustrated schematically in Figure 1h. Experimental studies The first experimental validation of the concept of the preferential uptake of the photosensitizer LDL complex in the ocular neovasculature, and consequently of selective thrombotic occlusion of the neovascular structures with minimal collateral photodynamic damage to the surrounding tissues, was performed in experimental models of the corneal neovasculature18 and tumor-induced choroidal neovascularization17 in rabbits. Benzoporphyrin derivative monoacide (BPD-MA) – currently available commercially in liposomal formulation under the name Verteporfin™ – was used either in liposomal formulation or was complexed with LDL before intravenous injection. Allison et al.20 and Richter et al.19 have shown that BPD in liposomal formulation injected into the blood circulation forms BPD-LDL complexes similar to those obtained extracorporeally. In corneal neovasculature, we were able to demonstrate that liposomal BPD and directly complexed BPD-LDL showed comparable vascular selectivity, and that subsequent photoactivation induced the same vascular effects.18 Figure 2 shows the chemical structure and action spectrum of BPD-MA. The uptake of liposomal BPD in the corneal vasculature, measured using laser-induced fluorescence (LIF) at different time intervals after drug injection, clearly demonstrates the preferential
→ Fig. 1a-h. Schematic illustration of the development of choroidal neovascularization (CNV) and the principle of its photodynamic therapy. a: Normal structure of the RPE with the photoreceptors on top and the capillaries of the choroid below. b: The thickened Bruch’s membrane separates the photoreceptor layer and the choroid, reducing the normal metabolism between the outer retina and the choroid. c: New pathological vessels emerge from the choriocapillaris, penetrate Bruch’s membrane, and grow into the subretinal space. d: The leaking neovasculature causes subretinal fluid deposition, retinal edema, and hemorrhage, leading to different degrees of RPE detachment, fibrovascular scar formation, and loss of visual function. e: The LDL complexed photosensitizer BPD-MA (Verteporfin™) is receptor-mediated accumulated in the neovascular structures. f: The light-activated photosensitizer produces highly reactive singlet oxygen, causing damage to the neovascular structure followed by thrombotic occlusion of the neovasculature. g: Resorption of the extravascular fluid and partial restoration of the retinal structure causes stabilization of visual function. h: The photosensitizer BPD-MA binds to serum LDL after it has been applied intravenously.
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a. Fig. 2. The chemical structure of the two regio-isomers of BPD-MA and the action spectrum of the photosensitizer. Photoactivation is performed with diode laser light around the peak wavelength of 689 nm.
b.
Fig. 3. Uptake and retention of LDL-complexed BPD-MA in the corneal neovasculature of rabbit eyes measured over 48 hours by laser-induced fluorescence.
binding to neovascular structures. The normalized relative BPD fluorescence measured at a wavelength of 696 nm was excited at 337 nm, using a pulsed nitrogen laser.18 Figure 3 shows the kinetics of uptake and retention with maximum sensitizer accumulation at about one hour post-injection. The retention of the sensitizer shows exponential decay with a halflife time of about six hours, which is much longer than its life time in the blood serum. This rapid accumulation and retention suggest a short drug-light interval for the effective photodynamic activation of BPD. The pronounced maximum of the action spectrum at an infrared wavelength of 694 nm enables good light transmission through the optical media of the eye, even with moderately cataractous lenses, and provides enough light penetration to reach chorioretinal neovascularizations below the retinal pigment epithelium (RPE) and/or hemorrhagic exudation. The photothrombotic effect in the neovasculature after light activation of BPD-MA is demonstrated in Figures 4a-d. In this example, the liposomal LDL sensitizer with a BPD dose of 2 mg/ kg body weight was administered intravenously into Dutch belted rabbits. One hour post-injection, laser irradiation at a wavelength of 694 nm and an irra-
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d. Figs. 4a-d. Photothrombotic effects in the corneal neovasculature of a rabbit eye. a,b: Color photographs of a corneal neovascular structure induced by mechanical irritation with intrastromal silk sutures before (a) and several minutes after (b) PDT, with light doses of 10 J/cm2 (arrow) and 25 J/cm2 (double arrow), respectively. c,d: Fluorescein angiography of the same eye before and after PDT. The neovascularizations are completely occluded in both treated areas.
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Fig. 5a,b. Light microscopy of the corneal neovasculature, before (a) and one hour after (b) PDT. After applying a light dose of 100 J/cm2, total occlusion of the vessels could be achieved by intravascular thrombus formation. The corneal stroma appears undamaged.
Fig. 6. Light micrographs of a Green’s melanoma planted into the choroid of a rabbit one hour post-PDT. The endothelial cells of the tumor vessels show hyperchromatic cell nuclei as a typical sign of cell damage. The lumina of the vessels are completely thrombosed.
Fig. 7. Electron micrograph of a vessel from an implanted iris tumor in a rabbit eye two hours post-PDT. The primary damage to the vascular endothelial cells can be clearly seen: cytoplasmatic extrusion into the vessel lumen (arrows) indicates total cell destruction.
diance of 100 mW/cm2 was performed using light doses of 10 J/cm2 and 25 J/cm2, respectively. Photographic documentation (Figs. 4a,b) and fluorescein angiography (Figs. 4c,d) demonstrate the structure and function of the neovasculature before (Figs. 4a,c) and several minutes after (Figs. 4b,d) photoactivation. Histological preparation shows the morphology of photodynamically-induced vessel occlusion. Light microscopy of semi-thin histological
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preparations one hour after PDT clearly shows complete thrombotic occlusion of the corneal neovasculature (Fig. 5b) in contrast to the completely perfused neovasculature prior to PDT (Fig. 5a). Further investigations into neovascular structures in experimental tumor models showed similar results after PDT with BPD-LDL complexes.17 One hour post-treatment, complete thrombosis of the tumor vasculature in Green’s melanomas implanted into the iris and choroid of rabbits was observed ophthalmoscopically and fluorescein-angiographically. No signs of recanalization of the occluded vessels were observed during the follow-up of up to three months. Light and electron microscopic evaluation of the treated vessel walls elucidated the mechanisms of thrombosis formation: as can be seen in Figure 6, one hour post-PDT, the endothelial cells are damaged and have substantial hypercromatic cell nuclei. The dilated vessel lumina are thrombosed. No collateral damage can be seen light microscopically in the adjacent tissue structures. Figure 7 shows the endothelial cell damage in greater detail in a higher resolution electron micrograph of an iris tumor vessel two hours post-PDT. The endothelial cell membranes have been disrupted and cytoplasm has extruded into the vascular lumen. All these experimental findings strongly support our proposed concept of the photodynamically-generated selective occlusion of neovascular structures: the primary photo-oxidative damage to the vascular endothelial cells induces thrombus formation by intravascular platelet activation and thrombocyte aggregation. The photodynamically-generated, so-called photothrombosis of neovasculatures demonstrated here, with its spatial confinement to the inner structures of the vessel walls and their lumina, stimulated further animal studies,21,22 mainly in order to establish optimal drug and light dosimetry. It also led to the first clinical evaluation of the photodynamic treatment of CNV membranes in age-related macular degeneration,23-28 and other chorioretinal pathologies.29-31 The results of extended multicenter clinical trials of
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the photodynamic therapy of choroidal neovascular membranes using Verteporfin™, and the current status of clinical experience, will be reviewed in the following chapter.32 Photosensitizers The experimental proof of the principle of the selective photodynamic therapy of neovascular structures, as well as the first establishment of this therapeutic modality for the treatment of CNV, was carried out using BPD-MA as photosensitizer complexed with LDL or a liposomally formulated (Verteporfin™). The huge success of this therapeutic approach and the encouraging experimental and clinical results,59,23-31,43 led to a number of further investigations using different photosensitizer compounds for the treatment of ocular neovascularities. More than ten new photosensitizers were experimentally tested for neovascular treatment in the eye: SnET2,33 Lutex,34 ATX S10,35 NPe6,36 ICG,37 Lambda 27,38 Hypocrellin A,39 MV 6401,40 and Tookad.41 Since NPe6 and ICG are hydrophillic compounds and Lambda 27, Hypocrellin A, and MV 6401 are lipophilic dyes, only ATX S10 has amphyphillic properties with a measurable affinity to lipid membranes and, at the same time, the advantage of water solubility.42 Practically all these sensitizers are designed to have a strong light absorption band in the near-infrared region (between 660 and 780 nm), in order to provide rather undisturbed ocular light transmission, even in eyes with mild cataract, and to facilitate enough light penetration through the hemorrhages and pigmented layers masking the neovasculature. The vessel closure experiments in rats, rabbits, and monkeys always followed the same principle: intravenous drug application was followed by light exposure of the vascular structure at the fundus of the animal eyes. Fluorescein angiography and histopathological examinations were performed in order to investigate the vessel closure and collateral damage in the treated areas. In most cases, variation of the drug and light dose and the drug-light interval resulted in clinically useful treatment parameters. A scientifically-based comparison between the different sensitizers is difficult at this point because of the very different stages of the various investigations. In general, the initial photodynamic effect with all sensitizers is damage to the vascular endothelium, which is then followed by intravascular thrombosis. However, the efficacy of this vessel closure mechanism and the side-effects in the retinal and choroidal vasculature, as well as the collateral damage in the neural retina and the retinal pigment epithelium, vary substantially for each sensitizer. Two of the newly-investigated sensitizers were also used in clinical trials to treat CNV in age-related macular degeneration: tin ethyletiopurpurin (SnET2; Purlitin™) and lutetium texaphyrin (Lutex). Generally speaking, in both these sensitizers, the pre-
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clinical experiments to validate the principle of intravascular thrombogenesis and to evaluate realistic treatment parameters, as well as the treatment protocols for the phase I/II and III clinical trials, were similar to the strategy originally used with the photosensitizer Verteporfin™. It is worth mentioning that almost all experimental studies, including those with new sensitizers not yet clinically evaluated, produced comparable results in terms of drug and light dosimetry, therapeutic mechanism, and time course for closure of the neovascular choroidal structures. Drug doses of 1-2 mg/kg body weight were applied systemically, and light doses of 20-50 J/cm2 were necessary to achieve vessel closure in the choroidal structures of rabbits and monkeys. Collateral damage in the RPE and neurosensory retina could be identified histologically, but this was small, thus indicating a substantial selectivity to vascular structures. These experimental findings were essentially consistent with those obtained with the liposomally formulated BPD-MA (Verteporfin™), a somewhat unexpected result due to the different biochemical properties of the sensitizers used. Verteporfin™ is a highly lipophilic compound in a liposomal formulation, Purlitin™ is also lipophilic, but is formulated in an emulsion to become water-soluble, and Lutex is hydrophillic in itself. These various properties should result in different drug kinetics in the ocular structures because of different uptake and retention in the tight retinal and in the fenestrated choroidal vascularization. However, the experimental results seemed to show a rather uniform damage profile in the vasculature, even with very different retention times in the skin, ranging from weeks (Purlitin™) to several hours (Verteporfin™). The clinical trials with Purlitin™ (phase III) and Lutex (phase I/II), were terminated before study reports had been published. Treatment efficacy as well as systemic and ocular side-effects were obviously not satisfactory, at least in comparison with the approved sensitizer, Verteporfin™. In contrast to the experimental findings, the clinical results indicate that efficacy and biodistribution of the sensitizers play an important role in the required selectivity of the ocular neovasculature, where new vascular structures have to be treated in the immediate vicinity of the anatomical resting retinal and large choroidal vasculature, and where functionally important tissues, such as the neurosensory retina, as well as extremely active phagocytic structures, such as the RPE, must remain undamaged. Further large scale research and studies will be required to establish new sensitizers for the treatment of neovascular structures in ocular tissue. Conclusions The concept of selective vascular PDT as the treatment principle for neovascular diseases in the eye was proven in animal experiments. Corneal neovas-
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Concept and experimental validation of photodynamic therapy cularizations as well as neovascular structures in chorioidal tumors were selectively thrombosed using the photosensitizer benzoporphyrine derivative monoacid ring A (BPD-MA) complexed with low density lipoproteins (LDL). The LDL-formulation was achieved experimentally either by covalently binding BPD-MA with LDL before intravenous injection or by liposomal delivery of BPD-MA. The selectively targeted neovascular structures showed damage of the intravascular endothelial layer right after photodynamic light activation. This damage was then followed by intravascular thrombosis. The therapeutic value of the achieved photothrombosis was successfully evaluated in clinical trials. The current status of the potential value and the limitations of this new treatment modality is described in a separate chapter of this book.32 After the first successful clinical trials numerous other photosensitizers have been investigated for vascular PDT in the eye both experimentally as well as clinically. Despite of experimentally encouraging results no photosensitizer other than BPD (Visudyne) has passed prospective clinical trials to date. References 1. Henderson BW, Dougherty TJ: How does photodynamic therapy work? Photochem Photobiol 55:145-157, 1992 2. Sery TW: Photodynamic killing of retinoblastoma cells with hematoporphyrin derivative and light. Cancer Res 39:96100, 1979 3. Gomer CJ, Dorion DR, White L, Jester JV, Dunn S, Szirth BC, Razum NJ, Murphree AL: Hematoporphyrin derivative photoradiation induced damage to normal and tumor tissue of the pigmented rabbit eye. Curr Eye Res 3:229-237, 1984 4. Murphree AL, Cote M, Gomer CJ: The evolution of photodynamic therapy techniques in the treatment of intraocular tumors. Photochem Photobiol 46:919-923, 1987 5. Packer AJ, Tse DT, Yuon-Qing G, Hayreh SS: Hematoporphyrin photoradiation therapy for iris neovascularization. Arch Ophthalmol 102:1193-1197, 1987 6. Thomas EL, Langhofer M: Closure of experimental subretinal neovascular vessels with hematoporphyrin-ether augmented argon green laser photocoagulation. Photochem Photobiol 46:881-886, 1987 7. Nanda SK, Hatchell DL, Tiedeman JS, Dutton JJ, Hatchell M, McAdoo T: A new method for vascular occlusion. Arch Ophthalmol 105:1121-1124, 1987 8. Kilman GH, Stern D, Gregory WA: Angiography and photodynamic therapy of experimental choroidal neovascularization using phthalocyanin dye. Invest Ophthalmol Vis Sci 30:S371, 1989 9. Miller H, Miller B: Photodynamic therapy of subretinal neovascularization in the monkey eye. Arch Ophthalmol 111:855-860, 1993 10. Honour AJ, Mitchell JRA: Platelet clumping in injured vessels. Br J Exp Pathol 45:75-87, 1964 11. Callahan AB, Lutz BR, Fulton GP, Dengelmann I: Smooth muscle and thrombus thresholds to unipolar stimulation of small blood vessels. Angiology 2:35-39, 1969 12. Boergen KP, Birngruber R, Hillenkamp F: Laser-induced endovascular thrombosis as a possibility of selective vessel closure. Ophthalmic Res 13:139-150, 1981
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13. Baumgartner HR: Platelet interaction with vascular structures. Thromb Diath Haemorrh 51:151-176, 1972 14. Ashford TB, Freiman DG: The role of the endothelium in the initial phases of thromboses. Am J Pathol 50:257-273, 1967 15. Jori G, Reddi E: The role of lipoproteins in the delivery of tumor-targeting photosensitizers. Int J Biochem 25:13691375, 1993 16. Roberts WG, Hasan T: Role of neovasculature and vascular permeability on the tumor retention of photodynamic agents. Cancer Res 52:924-930, 1992 17. Schmidt-Erfurth U, Baumann W, Gragoudas E, Flotte TJ, Michaud NA, Birngruber R, Hasan T: Photodynamic therapy of experimental choroidal melanoma using lipoproteindelivered benzoporphyrin. Ophthalmology 101(1):89-99, 1994 18. Schmidt-Erfurth U, Hasan T, Schomacker K, Flotte TJ, Birngruber R: In vivo uptake of liposomal benzoporphyrin derivative and photothrombosis in experimental corneal neovascularization. Lasers Surg Med 17:178-188, 1995 19. Richter AM, Waterfield E, Jain AK, Canaan AJ, Allison BA, Levy JG: Liposomal delivery of a photosensitizer, benzoporphyrin derivative monoacid ring A (BPD), to tumor tissue in a mouse tumor model. Photochem Photobiol 57:1000-1006, 1993 20. Allison BA, Pritchard PH, Levy JG: Evidence for low-density lipoprotein receptor-mediated uptake of benzoporphyrin derivative. (Published erratum appears in Br J Cancer 71(1):214, 1995) Br J Cancer 69:833-839, 1994 21. Miller JW, Walsh AW, Kramer M, Hasan T, Michaud N, Flotte TJ, Haimovici R, Gragoudas ES: Photodynamic therapy of experimental choroidal neovascularization using lipoprotein-delivered benzoporphyrin. Arch Ophthalmol 113:810-818, 1995 22. Kramer M, Miller J, Michaud N, Moulton RS, Hasan T, Flotte TJ, Gragoudas ES: Liposomal BPD verteporfin photodynamic therapy: selective treatment of choroidal neovascularization in monkeys. Ophthalmology 103:427-438, 1996 23. Schmidt-Erfurth U, Miller W, Sickenberg M, Bunse A, Laqua H, Gragoudas E, Zofragos L, Birngruber R, Van den Bergh H, Strong A, Manjuris U, Fsadi M, Lane AM, Piguet B, Bressler NM: Photodynamic therapy of subfoveal choroidal neovascularization: clinical and angiographic examples. Graefe’s Arch Clin Exp Ophthalmol 236:365-374, 1998 24. Miller JW, Schmidt-Erfurth U, Sickenberg M, Pournaras CJ, Laqua H, Barbazetto I, Zografos L, Piguet B, Donati G, Lane AN, Birngruber R, Van den Bergh H, Strong A, Manjuris U: Photodynamic therapy with Verteporfin for choroidal neovascularization caused by age-related macular degeneration: results of a single treatment in a phase 1 and 2 study. Arch Ophthalmol 117:1161-1173, 1999 25. Schmidt-Erfurth U, Miller JW, Sickenberg M, Laqua H, Barbazetto I, Gragoudas ES, Zografos L, Piguet B, Pournaras CJ, Donati G, Lane AN, Birngruber R, Van den Bergh H, Strong A, Manjuris U, Gray T, Fsadni M, Bressler NM: Photodynamic Therapy with Verteporfin for choroidal neovascularization caused by age-related macular degeneration: results of retreatment in a phase 1 and 2 study. Arch Ophthalmol 117:1177-1187, 1999 26. Photodymamic therapy of subfoveal choroidal neovascularization in age-related macular degeneration with Verteportin: one-year results of 2 randomized clinical trials: treatment of age-related macular degeneration with photodynamic therapy (TAP) study group: TAP report 1. Arch Ophthalmol 117:1329-1345, 1999 27. Photodynamic therapy of subfoveal choroidal neovascularization in age-related macular degeneration with Verteporfin:
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tizer, ATX-S10. Lasers Surg Med 24:209-222, 1999 36. Kazi AA, Peyman A, Unal M, Knoobehi B, Yoneya S, Mori Keisuke, Moshfeghi D, Moshfeghi AA: Threshold power levels for Npe6 photodynamic therapy. Ophthalmic Surg Lasers 31:136-142, 2000 37. Costa RA, Farah ME, Morales PH, Freymuller E, Smith R, Cardillo JA: Choriocapillaris photodynamic therapy using indocyanine green. Invest Ophthalmol Vis Sci 42:S438, 2001 38. Kazi AA, Peyman GA, Genaidy M, Farahat HG, Morgan A, Bisland S: Evaluation of threshold parameters for choroidal vessel closure with the new photosensitizer Lambda 27. Invest Ophthalmol Vis Sci 42:S437, 2001 39. Grinstead LR Jr, Khoobehi B, Peyman GA, Passos E: Experimental photodynamic effects of hypocrellin A on the choriocapillaris. ARVO Abstracts. Invest Ophthalmol Vis Sci 42:S437, 2001 40. Pratt LM, Criswell MH, Ciulla TA, Danis RP, Hill TE, Snyder WJ, Small W: Choriocapillaris closure in normal primates using photosensitizer MV6401 at different postinjection activation times. ARVO Abstracts. Invest Ophthalmol Vis Sci 42:S437, 2001 41. Framme C: Personal communication 42. Mori M, Kuroda T, Obana A, Sakata I, Hirano T, Nakajima S, Hikida M, Kumagai T: In vitro plasma protein binding and cellular uptake of ATX-S10(Na), a hydrophillic chlorine photosensitizer. Jpn J Cancer Res 91:845-852, 2000 43. Schmidt-Erfurth U, Hasan T, Gragoudas E, Michaud NA, Flotte TJ, Birngruber R: Vascular targeting in photodynamic occlusion of subretinal vessels. Ophthalmology 101(12): 1953-1961, 1994 44. Vlodavsky I, Fielding PE, Fielding CJ, Gospodarowicz D: Role of contact inhibition in the regulation of receptormediated uptake of low-density lipoprotein in cultured vascular endothelial cells. Proc Natl Acad Sci USA 75:356360, 1978
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Photodynamic therapy: clinical status Ursula Schmidt-Erfurth and Stephan Michels University Eye Hospital Luebeck, D-23538 Luebeck, Germany
Keywords: photodynamic therapy, age-related macular degeneration, indications, vascular effects
Introduction Within a few years, photodynamic therapy (PDT) with Verteporfin has become a standard treatment in exudative age-related macular degeneration (AMD). Just four years after the first promising results on the experimental PDT treatment of ocular neovascularization were published by SchmidtErfurth et al.,1,2 the first phase III multicenter placebo-controlled trials were conducted. This reflects the necessity for a new treatment modality in exudative age-related macular degeneration, one of the leading causes of legal blindness in northern America and Europe.3,4 The only proven treatment was nonselective thermal laser photocoagulation, which was only applicable in a limited number of patients with exudative AMD. In particular, the initial decrease of visual acuity in subfoveal lesions and the high rate of recurrences were considered detrimental.5 The aim of any new treatment for exudative AMD is the closure or removal of subretinal choroidal neovascularization (CNV) without significant damage to the surrounding tissues, such as photoreceptors or retinal pigment epithelium (RPE). Current treatment modalities other than PDT are radiation, transpupillary thermotherapy, and submacular surgery. The concept of radiation therapy is to effect fastproliferating neovascular structures. In tumor therapy, radiation is well established. However, a recent multicenter placebo-controlled trial of radiation in exudative AMD showed no treatment benefit compared to sham treatment.6 Transpupillary thermotherapy using an 810-nm laser within a range of 360-1000 mW is a relatively inexpensive treatment. By raising the choroidal temperature by about 10°C, pathological vessels are thought to be more susceptible to damage and clo-
sure than mature vessels. The first promising results were published by Reichel et al.7 However, the temperature margins are very close and over- or undertreatment can easily occur. The current TTT4 AMD trial will provide further information. Submacular surgery techniques have significantly improved, which has made macular translocation feasible. Significant visual improvement has been shown in selected cases,8 however, complications such as retinal detachment, proliferative vitreoretinopathy (PVR), cystoid macular edema, and cataract are common. In large multicenter, placebo-controlled trials over two years, photodynamic therapy (PDT) with Verteporfin has shown the stabilization of vision in patients with subfoveal choroidal neovascularization due to AMD and pathological myopia.9,10 Major advantages are the lack of significant adverse events and the possibility of being able to provide treatment in an outpatient setting. Hence, PDT should be ideal for the treatment of the elderly population affected by exudative AMD. Photodynamic therapy in predominantly classic choroidal neovascularization due to age-related macular degeneration The currently most relevant studies for the treatment of choroidal neovascularization in AMD and pathological myopia are the TAP (‘treatment of age-related macular degeneration with photodynamic therapy’) and the VIP (‘Verteporfin in photodynamic therapy’) trials.9,10 These were multicenter, double-masked, placebo controlled, randomized clinical trials, including a total of 948 patients with exudative subfoveal lesions due to AMD. The primary objective of the TAP and VIP trials
Address for correspondence: U. Schmidt-Erfurth, M.D., University Eye Hospital Lübeck, Ratzeburger Allee 160, D 23538 Lübeck, Germany. e-mail:
[email protected] Lasers in Ophthalmology – Basic, Diagnostic and Surgical Aspects, pp. 205–215 edited by F. Fankhauser and S. Kwasniewska © 2003 Kugler Publications, The Hague, The Netherlands
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was to prove a significant visual acuity benefit in Verteporfin-treated eyes. Vision was defined as stable if no more than three lines were lost on ETDRS (‘early treatment of diabetic retinopathy study’) vision charts. Patient selection for both studies was based on the fluorescein angiographic characteristics of subfoveal CNV. Depending upon their appearance on fluorescein angiography (FLA), lesions are characterized by their occult and classic components, as defined by the criteria of the Macular Photocoagulation Study (MPS) group.5 These characteristics correlate with the natural clinical course of the lesions. Predominantly classic lesions tend to show more aggressive growth and more rapid vision loss than occult lesions, which can be stable for longer periods. Based on the angiographical proportion of the classic component, a classification was created that divided the lesions into occult only (0% classic component), minimally classic (classic component >0% 50% classic component), and classic only (100% classic component). Inclusion criteria for TAP and VIP patients suffering from AMD were similar with regard to lesion localization and size (subfoveal; up to 5400 µm). However, the TAP studies only included eyes with at least minimally classic CNV due to AMD. In contrast, the VIP study was designed for eyes with occult lesions only, showing recent progression of disease. The two-year results of the TAP study showed a significant treatment benefit for the Verteporfintreated eyes, with stabilization of vision in 53% versus 38% in the placebo group. The number of patients who had severe vision loss (at least six lines lost on the ETDRS chart) was also significantly
higher in the in the placebo group (30% versus 18%). An increase of visual acuity (gain of at least one line on the ETDRS chart) was found in 13% of the Verteporfin- versus 7% of the placebo-treated eyes. Subgroup analysis of the above-mentioned components of the lesions with FLA showed significant differences in treatment benefits (Fig. 1). Subfoveal neovascularization composed of classic components only, treated with PDT, showed the highest percentage of vision stabilization (70% versus 29%), followed by predominantly classic lesions (59% versus 31%). No significant treatment benefit was seen in the minimally classic group. The few patients graded by the central reading center (Wilmer Angiographic Reading Center, Johns Hopkins University School of Medicine, Baltimore)9,10 as occultonly lesions were primarily misdiagnosed by the treating study center and were not intended to be included in the study. Further treatment benefits were found at 24 months for all patients with respect to stabilized contrast sensitivity (1.3 versus 5.2 letters lost), smaller lesion size, and less lesions with a classic component (48.8% versus 71.5%). At baseline, all demographic and lesion characteristics were sufficiently balanced. A total of 5.6 PDT treatments on average was applied from the onset of the study to the two-year endpoint. Factors predicting outcome were patient age, baseline visual acuity, lesion size, status of the fellow eye, and extent of invasion of the foveal avascular zone (FAZ). In general, visual results were better in patients of less than 65 years of age, with lower initial visual acuity, smaller lesions and better vision levels in the fellow eye. The chance of improvement in vision was highest in eyes in which
classic 4 DA
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207 Photodynamic therapy in choroidal neovascularization due to pathological myopia Pathological myopia is the second most common disease to present with subfoveal CNV. Due to ruptures in Bruch’s membrane, CNV can develop and frequently induces rapid vision loss. The population affected is characteristically younger and tends to have a better visual outcome than spontaneous AMD patients with CNV. The VIP study included 120 patients with subfoveal CNV due to pathological myopia (distance correction of at least –6 diopters). Stabilization of vision as the primary outcome parameter was defined more strictly than in the AMD population, i.e., a visual loss of less than eight letters (~1.5 lines on the ETDRS chart). However, baseline inclusion criteria were comparable to the AMD population, vision had to be 20/100 or better, and treatment was not limited to any specific age-group. There was a significant visual benefit in the PDTtreated versus the placebo group, with stabilization of vision in 72% versus 44% at 12 and 64% versus 49% at 24 months. An improvement in vision (a gain of at least one line on the ETDRS chart) was found in 32% versus 15% at 12 and 40% versus 13% at 24 months. The distribution of significant visual loss (a loss of eight letters or more) is shown in Figure 3. Moreover, the results regarding lesion size with FLA and contrast sensitivity were favorable. The most striking observation was the high rate of improvement in visual acuity, unlike that seen in the AMD population. At two years, 40% of patients in the PDT group showed an increase in vision of at least one line, compared to 13% in the placebo group; 12% even improved by three lines or more, an event that was never seen in the placebo-treated group. It is noteworthy that, due to the well-defined, classic nature of CNV in myopia, retreatments were indicated, based on angiographical leakage, at a higher rate in PDT patients than in placebo patients. It is probable that retreatments should not be indicated on angiographical data only, but also on the visual outcome and evidence of clinical activity. In our experience, CNV in myopic eyes requires less retreatments and appears to respond much faster to PDT. Photodynamic therapy indications unrelated to age-related macular degeneration and myopia Other diseases with secondary CNV that can be treated with PDT with Verteporfin are: parafoveal telangiectasia, angioid streaks, chorioretinitis, presumed ocular histoplasmosis syndrome (POHS), and idiopathic CNV. So far, no placebo-controlled, randomized trials have been conducted regarding these overall less frequent causes of secondary CNV. However, a number of case control studies support the benefit of PDT in patients with CNV not due to AMD or to pathological myopia.
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Choroidal neovascularization on FLA
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Fig. 2. Guideline for the management of neovascular macular disease in AMD.
VIP: Pathologic Myopia – Severe vision loss
Verteporfin % of patients (loss of at least 8 letters)
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Fig. 3. Proportion of patients presenting vision loss of more than 1,5 lines (≥ 8 letters) during a one year follow-up.
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Fig. 4. Clinical and ultrasonographical images of a macular choroidal hemangioma treated with PDT. A large hemangioma is seen affecting the temporal and central portion of the macula with a vision decline to 20/80 (pre PDT), ultrasound indicates a tumor height of 6.2 mm. Past PDT the lesion demonstrates progressive atrophy clinically together with a reduction in tumor prominence. A second (past PDT II) and third PDT were performed, which induced a complete regression of the angioma to a flat scar with granular RPD atrophy. No prominence can be detected by ultrasound (past PDT III).
In parafoveal telangiectasia, CNV develops on average within 73 months of the initial diagnosis, and the final visual acuity is less than 20/200 in 80% of patients.11 Potter et al. and Müller-Velten et al. reported stabilization and improvement of vision in patients with secondary CNV due to parafoveal telangiectasia, chorioretinitis, and idiopathic CNV.12,13 Secondary CNV in angioid streaks tends to show rapid progression and loss of vision during its natural course. A 70% stabilization of vision (loss of less than three lines) over 12 months has been shown in a study by Virgili et al.14 Response to PDT seems to be quite variable, and shorter retreatment intervals have been recommended.15 A multicenter, uncontrolled study on the treatment of CNV in POHS with Verteporfin showed an average increase of vision of one line after 12 months. Eighty-eight percent of patients had stabilized vision and 28% had an improvement in vision of more than three lines.16 Extended indications for PDT may occur when considering the individual situation of each patient. Results of controlled studies are to be expected within the next few years, but are not available currently. Photodynamic therapy as a novel modality in the treatment of intraocular hemangioma Decompensated choroidal hemangioma may have a devastating course, with 75% of patients having a long-term visual acuity of less than 20/50, and onethird of patients vision of less than 20/160 after ten years.17,18 In an interventional case control study, Schmidt-Erfurth et al.19 showed progressive reduction of tumor height, improvement and stabilization
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of vision in all patients, as well as a continuously improving central visual field. The mean increase in visual acuity was three lines on the patients’ ETDRS charts, with increases of up to eight lines. Choroidal tumors responded reliably to PDT with a predictable decrease in height following each treatment course (Fig. 4). All tumors in the choroidal angioma group showed complete regression, together with resolution of the macular edema and exudative detachments. The parameters used in this study were 100 J/cm2 of light, a drug dose of 6 mg/m2 body surface area, and a retreatment interval of six weeks. No side-effects were observed and no recurrence was seen in any of the patients, even after as long as five years. PDT appears to have become the treatment of choice in choroidal hemangioma involving the macula. Retinal capillary angioma of the optic nerve only responded poorly to PDT. Moreover, lesions showed progressive fibrosis, and massive transudation following PDT, intravitreal hemorrhage, and occlusion of large retinal vessels were observed. Currently, it is not recommended to apply PDT for this condition.20 Photochemical mechanisms of photodynamic therapy PDT with Verteporfin PDT is based on two interacting mechanisms: the accumulation of a light-activated chromophore in the pathological tissue and the induction of a chemical reaction by light activation in the presence of oxygen. Each photosensitizer has its characteristic wavelength for activation. In neovascular AMD, the target lesion is easily accessible via a slit-lamp setting, and may be covered completely and homogenously with the sensitizing light.
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In the clinical setting, the photosensitizer is usually applied systemically and is distributed throughout the body. Neoplastic and neovascular tissues typically show preferential uptake and retention of the photosensitizer. Liposomally-delivered Verteporfin aggregates intravascularly with low-density lipoprotein (LDL) and accumulates in proliferating cells.21,22 LDL receptors are expressed at a ten-fold higher rate in rapidly proliferating cells, which probably enhances the binding and uptake of Verteporfin in the neovascular tissue compared to normal vasculature.23,24 Angiography using Verteporfin in experimental CNV demonstrated the preferential uptake of Verteporfin in CNV lesions. A similar photochemical reaction can be found for all photosensitizers.25 The photosensitizer absorbs light energy at its characteristic wavelength peak. Verteporfin is a second-generation lipophilic photosensitizer derived from porphyrin. In comparison to earlier photosensitizers, Verteporfin is four times more efficient in absorbing light. Consequently, the cytotoxic effect is significantly higher.26 Absorption of light leads to a shift of its singlet electronic state from the lower energetic level S0 to the higher energetic level S1. By intersystem crossing at the level of electrons, S1 gives rise to the exited triplet state T1. A photochemical reaction is initiated by T1 either directly by generating reactive, cytotoxic, free radicals, or indirectly by transferring its energy to ground state oxygen, and consequently to singlet oxygen. The photo-oxidative damage to biological structures is similar for both pathways.27 However, it is thought that the formation of singlet oxygen is the primary mechanism of PDT-induced damage for most photosensitizers currently in use. Oxygen is abundantly present in the choroid of the human eye, which facilitates oxidative processes. The PDT-induced cell and tissue damage is primarily based on cellular, vascular, and immunological mechanisms.28,29 The relative contribution of each mechanism is dependent upon the photosensitizer, the tissue being treated, and the timing.30 However, the characteristic mechanism induced by PDT with Verteporfin in vivo seems to be vascular damage followed by blood-flow stasis, and finally vasoocclusion.29 Endothelial damage with swelling, fragmentation, and exposure of the basement membrane, as well as thrombosis including red blood cells, thromocytes, and fibrin, was seen histologically within treated CNV surgically removed after unsuccessful treatment, as well as in the physiological choroid of human eyes. Damage to the endothelial cells is the first step in the cascade towards blood flow stasis. Breaks in the intraluminal endothelial membranes are seen at an ultrastructural level. This consequently leads to the endothelial cells shrinking away from each other.31 Damage to the endothelial cells with exposure of the basement membrane triggers platelet binding and aggregation. The release of vasoactive substances, such as thromboxane, histamine, and tumor necrosis factor-α (TNF-α), in-
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duces amplification of platelet activity, thrombosis, vascoconstriction, and increased vascular permeability. Blood-flow stasis, tissue hypoxia, and possible shutdown of the vasculature is the consequence.32 Due to the selective nature of the treatment, the RPE, photoreceptors, and retinal vasculature remained intact during PDT using Verteporfin, which is the basis for the maintenance of visual function. Parameters in clinical use The biodistribution of a photosensitizer is crucial for the timing of photoactivation, in order to maximize closure of the CNV and to minimize collateral damage. Optimizing treatment parameters with the aim of minimizing the dose of the photosensitizer, without excessively prolonging laser application, is essential (see also ref. 33 and 34). The currently recommended treatment protocol has been standardized, and has been proven to be safe and effective in TAP and VIP trials. A diode laser with a slit-lamp delivery system is used. Fifty J/cm2 is delivered at an intensity of 600 mW/cm2 for a duration of 83 seconds. Before treatment, the greatest linear diameter of the lesion is measured on a standardized fluorescein angiogram. An additional 1000 µm is added to ensure a circular safety margin of 500 µm. Verteporfin is prepared in a 30-ml glucose solution at a dose of 6 mg per square meter of body surface area, in a darkened room to avoid premature activation of the photosensitizer. The drug is given intravenously over ten minutes, via a cubital access in order to avoid extravasation of verteporfin which can cause ulceration of tissue following light exposure. Fifteen minutes after the start of infusion, laser light is applied via a hand-held contact lens. Follow-up is recommended at three-month intervals. Retreatment should be considered if FLA shows persistent leakage from the lesion. Further studies are being conducted to compare three- versus twomonth intervals, and to evaluate the clinical appearance (visual acuity, metamorphopsia, biomicroscopy) on decision-making for PDT retreatment. Optical coherence tomography may be a useful tool to detect and quantify subretinal fluid due to active CNV reliably (Puliafito C, ARVO abstract 2002). Characteristic angiographical features following photodynamic therapy Angiography is the most important method for evaluating and classifying CNV. Indications for PDT treatment are based on fluorescein angiography and angiographical follow-up is essential to document the closure and regression of CNV. Not as commonly used, but at least as informative, is indocyanine green angiography (ICGA). The advantages of this technique are its longer wavelength in comparison to FA, with better visualization of the choroid by reducing
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A
B
C
D
Fig. 5. Pre PDT a lesion, identified as retinal angiomatous proliferation with a subfoveal CNV is seen by early ICGA frames (A). Five hours after treatment the neovascular net is less well perfused, but not completely occluded (B). However, one day after PDT, no choroidal net is detectable (C). One week after PDT, a neovascular lesion is absent angiographically and the choriocapillary reveals marked non-perfusion within the treatment site (D).
A
B
C
Fig. 6. Late ICG reveals the sequence of PDT-related events in the choroid: A plaque-like hyperfluorescent lesion is seen before PDT (A). Five hours after exposure an additional break-down of vascular barriers in the choroid leads to increased fluid pooling (B), which is followed by choroidal hypofluorescence consistent with occlusion of choriocapillary vessels (C).
masking phenomena. Less extravasation by binding to larger proteins is also useful when documenting the neovascular structures precisely. The typical ICG angiographic course of CNV in AMD responding to PDT presents the following characteristics: Before PDT the vascular structure is well delineated in early frames (Fig. 5A). At five
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hours after treatment the peripheral structures of the choroidal net are rarified, but most of the CNV is still perfused (Fig. 5B). One day after PDT the membrane is mostly completely occluded as seen in early frames of the angiogram (Fig. 5C). Marked hypofluorescence within the treatment spot is the characteristic appearance in early frames at one week after
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Fig. 7. Typical fluorescein angiographic course of a patient treated with PDT. Images: before PDT (A/B) at 6 months (C/D) and at 12 months (E/F)
PDT with no significant activity of the CNV (Fig. 5D). The late phase of an angiogram at five hours (Fig. 6B) reveals typically an increased leakage in comparison to pre PDT images (Fig. 6A) as a consequence of significant vascular damage to the CNV and to the choroid covered by the treatment spot. Angiograms at one week (Fig. 6C) show in late frames a persistent hypofluorescence as sign of significant choroidal hypoperfusion but no leakage from the CNV. At eight to 12 weeks after PDT, partial recanali-
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zation of the CNV net and leakage can be seen. The TAP study reported a mean of 3.4 treatments during the first, and 2.2 during the second, year.8 Fluorescein angiographically, the lesion retracts progressively and looses its activity, documented by the transformation of leakage to staining. During follow-up, CNV typically persists, but is converted into inactive, partially-fibrotic tissue. A typical course with repeated PDT is shown in Fig. 7A-F. For the maintenance of long-term function and structural integrity, overtreatment should be avoided. Histological studies in human eyes showed closure
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Fig. 8. Characteristic features by ICG angiography. The CNV is seen before PDT in an intact choroidal bed (A-B). After 12 months choroidal perfusion is markedly reduced (C-D) and remains blocked with choriocapillary closure (E-F).
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of the choriocapillary layer one week after treatment, which correlates with a hypofluorescent spot particularly seen in ICGA.33 The choriocapillary typically recovers within eight to 12 weeks, but repeated treatments can induce persistent hypoperfusion, and consequently atrophy of the retinal pigment epithelium.34,35 A typical example of angiography is shown in Figures 8a-f. Significant hypofluorescence can be seen in the early and late phases of the angiogram at 12 months, and partial recovery at 24 months of follow-up. Clinical recommendations The current indications for PDT with Verteporfin are quite well-defined, however, experience in angiographical and clinical evaluation is necessary in order to decide which patients will benefit from PDT. Patients with at least predominantly classic CNV in AMD, and those with CNV in pathological myopia, seem to benefit most from PDT. But also, progressive occult lesions with low visual acuity and/or smaller lesion size show promising results after PDT. Juxtafoveal lesions should be treated with PDT, if it is expected that laser photocoagulation will affect the center of the FAZ. Figure 2 provides a guideline for PDT treatment in exudative AMD. Guarded prognosis is to be expected in patients with chorioretinal anastomosis, which can best be seen with ICGA after single PDT treatment failure. There are a number of ocular diseases with secondary CNV other than AMD and pathological myopia. Preliminary case reports indicate that PDT is an effective treatment option, as long as the lesion is composed of a predominantly classic component. However, each patient should be evaluated on an individual base.16 Further investigations will show whether modified treatment intervals will improve outcomes, and whether adjuvant therapy with antiangiogenic compounds or modified steroids (e.g., Anecortave Acetat) will significantly improve treatment outcomes in choroidal neovascular diseases (Slakter J et al., ARVO abstract 2002). Conclusion Photodynamic therapy is a useful novel modality to reduce progression of neovascular age-related macular degeneration and classic choroidal neovascularization due to other pathologies. Future studies are underway to expand applications and optimize vision outcome. References 1. Schmidt-Erfurth U, Birngruber R, Hasan T: Selektiver Verschluß okulärer Neovaskularisationen durch photodynamische Therapie. Ophthalmologe 89:391-394, 1992
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2. Schmidt-Erfurth U, Hasan T, Gragoudas E et al: Vascular targeting in photodynamic occlusion of subretinal vessels. Ophthalmology 101(12):1953-1961, 1994 3. Klein R, Klein BE, Linton KL: Prevalence of age-related maculopathy: the Beaver Dam Eye Study. Ophthalmology 99:933-943, 1992 4. Vingerling JR, Dielemans I, Hofman A et al: The prevalence of age-related maculopathy in the Rotterdam Study. Ophthalmology 102:205-210, 1995 5. Macular Photocoagulation Study Group: Subfoveal neovascular lesions in age-related macular degeneration. Arch Ophthalmol 109:1242-1257, 1991 6. Tholen AM, Meister A, Bernasconi PP, Messmer EP: Radiotherapy for choroidal neovascularization (CNV) in agerelated macular degeneration (AMD). Klin Mbl Augenheilk 216(2):112-115, 2000 7. Reichel E, Berrocal AM, IP M et al: Transpupillary thermotherapy of occult choroidal neovascularization in patients with age-related macular degeneration. Ophthalmology 106:1908-1914, 1999 8. Aisenbrey S, Lafaut BA, Szurman P et al: Macular translocation with 360 degrees retinotomy for exudative agerelated macular degeneration. Arch Ophthalmol 120(4): 451-459, 2002 9. The Treatment of Age-Related Macular Degeneration with Photodynamic Therapy (TAP) Study Group: Photodynamic therapy of subfoveal choroidal neovascularization in agerelated macular degeneration with Verteporfin. Arch Ophthalmol 119:198-207, 2001 10. Verteporfin in Photodynamic Therapy Study Group: Verteporfin therapy of subfoveal choroidal neovascularization in age-related macular degeneration: two-year results of a randomized clinical trial including lesions with occult and no classic choroidal neovascularization. Am J Ophthalmol 131(5):541-560, 2001 11. Engelbrecht NE, Aaberg TM, Lewis LM et al: Neovascular membranes associated with idiopathic juxtafoveal telangiectasia. IOVS 42:800, 2001 12. Potter MJ, Szabo SM, Chan EY et al: Photodynamic therapy for choroidal neovascularization in type 2a idiopathic juxtafoveal retinal telangiectasis. IOVS 42:438, 2001 13. Müller-Velten R, Michels S, Schmidt-Erfurth U, Laqua H: Photodynamische Therapy: Erweiterte Indikationen. Ophthalmologe 2002 (in press) 14. Virgili G, Pece A, Giacomelli G: Photodynamic therapy of choroidal neovascularization associated with angioid streaks. IOVS 43/2002 (in press) 15. Shaikh S, Ruby A, Williams G: Photodynamic therapy using verteporfin for choroidal neovascularization in angioid streaks. IOVS 43/2002 (in press) 16. Saperstein D: Effects over two years of verteporfin therapy on choroidal neovascularization secondary to the ocular histoplasmosis syndrome (OHS). IOVS 43/2002 (in press) 17. Augsburger JJ, Shields JA, Moffat KP: Circumscribed choroidal hemangiomas: longterm visual prognosis. Retina 1:56-61, 1981 18. Anand R, Augsburger JJ, Shields JA et al: Circumscribed choroidal hemangiomas. Arch Ophthalmol 107:1338-1342, 1989 19. Schmidt-Erfurth U, Michels S, Kusserow C et al: Photodynamic therapy for symptomatic choroidal hemangioma: visual and anatomical results. Ophthalmology 2002 (december) 20. Schmidt-Erfurth U, Kusserow C, Barbazetto I, Laqua H: Benefits and complications of photodynamic therapy of papillary capillary hemangiomas. Ophthalmology 109: 12561266, 2002 21. Allison BA et al: Evidence for low-density lipoprotein
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23. 24.
25.
26.
27. 28.
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receptor-mediated uptake of benzoporphyrin derivative. Br J Cancer 69:833-839, 1994 Schmidt-Erfurth U et al: Photodynamic targeting of human retinoblastoma cells using covalent low-density lipoprotein conjugates. Br J Cancer 75:54-61, 1997 Fogelman AM et al: Lipoprotein receptors and endothelial cells. Semin Thromb Hemost 14:206-209, 1988 Gaffney J et al: Differences in the uptake of modified low density lipoproteins by tissue cultured endothelial cells. J Cell Sci 79:317-325, 1985 Ochsner M: Photophysical and photobiological processes in the photodynamic therapy of tumors. J Photochem Photobiol B 39:1-18, 1997 Richter AM et al: Preliminary studies on a more effective phototoxic agent than hematoporphyrin. J Nat Cancer Inst 79:1327-1332, 1987 Kessel D: Sites of photosensitization by derivatives of hematoporphyrin. Photochem Photobiol 44:489-493, 1986 Hasan T, Parrish JA: Photodynamic therapy of cancer. In: Holland JF (ed) Cancer Medicine, pp 739-751. Baltimore, MD: Williams & Wilkins 1997
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215 29. Henderson BW, Dougherty TJ: How does photodynamic therapy work? Photochem Photobiol 55:145-157, 1992 30. Haimovici R et al: Localization of lipoprotein-delivered benzoporphyrin derivative in the rabbit eye. Curr Eye Res 16:83-90, 1997 31. Schmidt-Erfurth U et al: Photodynamic therapy of experimental choroidal melanoma using lipoprotein-delivered benzoporphyrin. Ophthalmology 101:89-99, 1994 32. Fingar VH: Vascular effects of photodynamic therapy. J Clin Laser Med Surg 14:323-328, 1996 33. Schmidt-Erfurth et al: Histopathological changes following photodynamic therapy in human eyes. Arch Ophthalmol 120(6):835-844, 2002 34. Schmidt-Erfurth U, Michels S et al: Photodynamic effects on choroidal neovascularization and physiological choroid. Invest Ophthalmol Vis Sci 43:830-841, 2002 35. Michels S, Barbazetto I, Schmidt-Erfurth U: Changes in neovascular membranes and normal choroid blood vessels after multiple photodynamic therapy treatments. Ophthalmologe 99(2):96-100, 2002
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Controversial aspects of photodynamic therapy Keisuke Mori,1 Darius M. Moshfeghi,2 Gholam A. Peyman3 and Shin Yoneya1 Saitama Medical School, Saitama, Japan; 2Department of Ophthalmology, Cleveland Clinic Foundation, Cleveland, OH, USA; 3Department of Ophthalmology, Tulane University Health Sciences Center, New Orleans, LA, USA
1
Keywords: photodynamic therapy, senile, neovascular, exudative maculopathy, photosensitizers, hydrophilic versus hydrophobic photosensitizers, chemistry, results, pitfalls
Introduction Photodynamic therapy (PDT) is generating intense interest in ophthalmology with the recent introduction of verteporfin (Visudyne; Novartis AG, Bulach, Switzerland) as a treatment for age-related macular degeneration (AMD).1-3 The two-year results of the Treatment of Age-related Macular Degeneration Study indicated that verteporfin therapy is more likely than placebo to result in stabilization of visual acuity in patients with predominantly classic choroidal neovascular membranes secondary to AMD at two years.4 However, several controversial points in PDT require further and detailed discussion. In this chapter, the following four aspects are discussed: selectivity of photosensitizers for targeting choroidal neovascularization (CNV); pitfalls in the application of PDT in the clinical setting; photosensitizer properties; and future trends in photodynamic and antiangiogenic therapy.
pathway of photodynamic damage to cells has been identified as a type II photosensitization reaction: singlet oxygen directly damages the plasma membrane and intracellular membranes (Fig. 1).6,9 Typical values of the decay lifetime of singlet oxygen are 3 µsec in H2O, 30 µsec in deuterium oxide, 12 µsec in ethanol, and 0.2 µsec in living cells.6,9 Therefore, a singlet oxygen molecule can diffuse only about 0.1 µm during its lifetime in tissue, limiting the primary reactions to the initial localization sites.6,9 This is the basis of the selective targeting of PDT within the targeted tissues. Although sensitizers accumulate in neovascular tissue such as CNV membranes, adjacent tissues, especially retinal pigment epithelium (RPE) cells and photoreceptor cells, are also moderately damaged. This finding was demonstrated by a series of experiments using verteporfin and other photosensitizers (Fig. 2). Pitfalls of PDT: clinical application
Photosensitizer selectivity: observations based on histological findings PDT is predicated upon the selective damage of pathological tissue while preserving surrounding normal tissue, and was originally designed to treat solid tumors and their associated neovascular processes.5 PDT consists of local activation of a photosensitizer using laser irradiation of an appropriate wavelength, after that photosensitizer has selectively accumulated in neovascular tissues. This results in activation of singlet and reactive oxygen species and other free radicals that can directly induce vascular cell death or occlude the vascular supply through the activation of the clotting cascade.6-8 The main
PDT is a straightforward treatment. Patients are administered the dye, either by intravenous infusion (hydrophobic photosensitizers, i.e., verteporfin) or intravenous bolus (hydrophilic photosensitizers, i.e., mono-L-aspartyl chlorin e6 (NPe6)), and then seated at the slit lamp, which is coupled to the appropriate diode laser system. A contact lens is placed on the ocular surface with a coupling agent such as Goniosol or Artificial Tears. At the optimal time, the laser is activated for the treatment duration. Care is taken to use the minimum slit-lamp illumination necessary to identify macular landmarks, thereby minimizing the possibility of inadvertent activation of the photosensitizer by the white light.
Address for correspondence: Gholam A. Peyman, MD, Tulane University Health Sciences Center, 1430 Tulane Avenue SL-69, New Orleans, LA 70112-2699, USA. e-mail:
[email protected] Lasers in Ophthalmology – Basic, Diagnostic and Surgical Aspects, pp. 217–227 edited by F. Fankhauser and S. Kwasniewska © 2003 Kugler Publications, The Hague, The Netherlands
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a
b
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d Fig. 1. Electron spin resonance (ESR) spectrum after 664-nm diode laser irradiation to NPe6 solutions with the trapping agent of 2,2,6,6-tetramethyl-4-piperidone hydrochloride (TMP) for the detection of singlet oxygen (a) and with the trapping agent of 5,5dimethyl-1-pyrroline-N-oxide (DMPO) for the detection of superoxide (b) and hydroxyl radical (c). In general, the pathways of photosensitized reactions are divided into two broad categories: types I and II photosensitization (d). Type I photosensitization typically involves electron transfer between photosensitizer molecules and substrate molecules. The type I pathway is more likely for photosensitizer triplet states than for excited singlet states because of the longer lifetimes. Signals of superoxide and hydroxyl radical (b, c) indicate type I photosensitization process. Type II photosensitization generates singlet oxygen, which is identified as the main product of photodynamic damage to cells.
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Fig. 2. Light micrograph obtained seven days after PDT in the normal Macaca monkey retina. Optimal power setting with choroidal vascular occlusion, but without retinal vascular occlusion. The structure of sensory retina was preserved. However, note the shortening of outer segments of photoreceptors, migration of retinal pigment epithelial cells, and cytoplasmic debris in the subretinal space. (Reproduced from Mori et al.38 by courtesy of the publisher.)
There are several factors that can influence the outcome of PDT: intraocular pressure, region of fundus treated, ocular pigmentation, duration of laser exposure, and retreatment.10 Too much pressure when placing the contact lens against the cornea can result in the effect of PDT being diminished.10 Similarly, treatment spots of similar fluence but longer exposure time result in greater angiographic and funduscopic lesions than shorter exposure time and higher power spots.10 Fundus pigmentation has been demonstrated to affect the appearance of treatment spots in pigmented rabbits.10 Additionally, a treatment spot location in the posterior segment is more likely to result in a therapeutic effect than a peripheral lesion.10 Fortunately, unlike in laser photocoagulation, the laser spot does not have to be focused perfectly upon the retina in order to achieve a therapeutic effect.10 If all these factors have been controlled and there does not appear to be an angiographically visible treatment spot within 24 hours, then the possibility of inadvertent subcutaneous injection must be considered. This complication is difficult to miss at the time of treatment, as the clinical signs of infiltration are usually visible in the antecubital region, and can be associated with discomfort at the site. When retreating patients, it is wise to consider that the treatment may result in cumulative damage to the RPE, photoreceptors, and inner retina.10,11
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Selection of photosensitizers: hydrophilic versus hydrophobic sensitizers, mechanisms of dye accumulation, and indocyanine green-guided PDT Photosensitizers were originally thought to be selective to neoplastic tissue because they were noted to accumulate in solid tumors.12,13 Eventually it was discovered that the tumor vasculature was the primary target of PDT.14,15 This property of biodistribution and treatment selectivity may play a beneficial role in treating subfoveal CNV. However, the firstgeneration photosensitizers were limited by two main factors: suboptimal tissue penetration of the maximum exciting wavelength and prolonged cutaneous phototoxicity.16-19 An advantageous sensitizer should have low skin photosensitization and high photosensitizing ability in targeted tissue when exposed to far-red light. Many second-generation sensitizers have been designed to overcome these problems for clinical application. These second-generation sensitizers possess major absorption peaks at wavelengths above 650 nm.20 Verteporfin, a synthetic chlorin-like porphyrin with a light absorption peak at 692 nm, is a hydrophobic sensitizer in a liposomal preparation for administration by intravenous infusion over ten minutes.1,2 The mechanism of verteporfin uptake by neovascular tissue is not known, but it has been suggested that neovascular and neoplastic tissues have increased lipoprotein receptors that may enhance the preferential uptake of verteporfin.21,22 In the pilot examination using experimental CNV, verteporfin was coupled with low-density lipopro-
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b Fig. 3. Light and transmission micrograph obtained two hours after NPe6 PDT. a. Outer retina and choroid of the normal Macaca monkey. Note swollen endothelial cell nuclei of choriocapillaris and vacuolations in retinal pigment epithelial cells. b. The retinal pigment epithelial cells. The architecture of basal infolding and microvilli was approximately intact. Numerous vacuolations appeared in the apical and basal side of retinal pigment epithelial cells. Some mitochondria were swollen, and their inner membranes that formed cristae were often destroyed. The endoplasmic reticulum and nucleus were minimally affected.
tein (LDL) in order to enhance its delivery to neovascular and tumor tissues and its PDT effect.23 CNV may selectively accumulate lipoprotein-associated sensitizers because of increased LDL receptors in rapidly proliferating endothelium and increased LDL transport across the endothelium of permeable vessels.24,25 Liposomal verteporfin was used in the second series of the experiments and clinical trials which demonstrated the relative selectivity in CNV occlusion.1-4,26,27 This indirect evidence suggests that the verteporfin binds to LDL in the blood and then
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accumulates preferentially in neovascular tissue in the choroid.28 Other second-generation sensitizers have been proposed for PDT, and the list continues to grow.29 Representative second-generation sensitizers include phthalocyanine dyes,30,31 purpurins,32 ATX-S10,33 and NPe6, which have all been investigated for the treatment of ocular diseases. Among these dyes, there are significant differences in the time intervals required for peak uptake and retention levels. In a mouse study, peak circulating levels of photosensi-
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c
d Fig. 3. Light and transmission micrograph obtained two hours after NPe6 PDT. c. and d. The endothelial cells of choriocapillaris were more severely damaged. Note the numerous vacuolations with single-unit membrane, which might be lysosomes. Some mitochondria remained approximately normal. Some nuclei were swollen and lost structural integrity of chromatin. (Figs. 3a and b are reproduced from Mori et al.38 by courtesy of the publisher.)
tizer were reached for hematoporphyrin derivative after five to ten hours, for phthalocyanine dyes after 24-48 hours, for verteporfin after three hours, and for NPe6 after two to 60 minutes.29 Levels of singlet oxygen quantum yields were found to be: hematoporphyrin derivative, 0.29; purpurins, 0.67; phthalocyanine dyes, 0.36; and NPe6, 0.77.29 NPe6 is a second-generation photosensitizer that has undergone preliminary clinical trials in humans.34 The advantages of NPe6 are minimal skin photosensitization, a major absorption peak at a far-red
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wavelength, hydrophilic formulation allowing intravenous bolus administration with rapid tissue uptake and clearance, high affinity for neovascular tissues, and a high photosensitizing ability.20,29,35-37 We have previously demonstrated that combining NPe6 with a laser at 664 nm allowed efficient occlusion of choroidal vessels with minimal injury to the overlying sensory retina of normal rabbits and monkeys, and that the primary targeted organelle may be the lysosome (Fig. 3).38 This finding agrees with the reports by Roberts and colleagues39-41 that NPe6
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Fig. 4. Fluorescence microscopy of NPe6. NPe6 accumulated intensively in the choroidal neovascular lesions (arrowhead) and retinal pigment epithelial cells. Some choroidal vascular walls also fluoresced moderately. In contrast, there was trace fluorescence in retinal vascular walls (arrow) and microcapillaries. There was little fluorescence of NPe6 detected in retinal stroma, vascular lumen of the retina and choroid, vitreous cavity and subretinal space. (Reproduced from Mori et al.42 by courtesy of the publisher.)
enters the cell by endocytosis and is localized in lysosomes, whereas other hydrophobic sensitizers enter the cell by diffusion and are located throughout the cytoplasm. These properties on cellular uptake and targeting selectivity might result from the hydrophilic nature of NPe6.39-41 Fluorescence microscopy demonstrated the strong accumulation of NPe6 in CNV and RPE cells,42 possibly indicating active endocytosis in these locations. In contrast, there was trace fluorescence in the retinal vascular walls and the subretinal space, and no fluorescence in the retinal stroma or vitreous cavity (Fig. 4).42 These fluorescence microscopic findings support the hypothesis of the beneficial properties of active accumulation in CNV mediated by the lysosomal uptake, indicating targeting selectivity by NPe6. Biodistribution of NPe6 in the fundus with experimental CNV has also been examined with fundus NPe6 videoangiography, utilizing a scanning laser ophthalmoscope and a 488nm argon laser light that fit the minor absorption peak of NPe6.39 The peak time of dye accumulation in experimental CNV with this system was 20-60 minutes after the dye injection, as evidenced by
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increasing fluorescence intensity of retinal vessels after 20-60 minutes, followed by a decline in intensity (Fig. 5a-c). Although the dye concentration used for NPe6 angiography was higher than that in the PDT experiment, these angiographical findings may provide a rough estimation for the optimal timing of the laser irradiation. An additional finding during the NPe6 angiography experiments was that there was a great similarity in dye-filling, dye accumulation, and retention patterns of NPe6 and indocyanine green (ICG) (Fig. 5a-f). Both NPe6 and ICG have the same biochemical properties: they are hydrophilic, of approximately equivalent molecular weights, and have a high affinity for lipoproteins.38-43 These characteristics may play an important role in the biodistribution of dye, and provide an explanation for the similarity of those angiograms. ICG angiography, which has been applied in a wide spectrum of choroidal diseases (especially AMD), has been proven safe for clinical use,44 and provides an estimation or a simulation of the NPe6 fundus biodistribution in various patients with AMD. This indirect evidence suggests the enhancement of practicality of NPe6 PDT by the guidance of ICG, especially in the determination of optimal timing for laser irradiation. ICG-guided PDT also has advantages (in the detailed parameter settings pointed out in the section on pitfalls in PDT). We may be able to estimate the parameters based on ICG-angiography findings of dye accumulation, fundus pigmentation, and other fundus conditions such as retinal edema, subretinal hemorrhage, and exudates. Future trends in PDT: prevention of CNV recurrence; modulation of CNV with antiangiogenic therapy One of the most important problems with current PDT is the high rate of recurrence of CNV.3,4 In the clinical study of a single treatment of verteporfin PDT, fluorescein leakage reappeared in at least a portion of the CNV one to three months after treatment.3 Increasing the photosensitizer concentration or the light doses did not prevent the recurrence and could lead to undesirable, nonselective damage to the retinal vessels.1 The two-year results of large clinical trials showed decreased rates of moderate vision loss; however, 5.6 treatments were needed during the two-year follow-up period.4 The necessity for multiple PDT sessions can be expected to lead to cumulative damage to the RPE and choriocapillaris, and to possible progressive treatmentrelated vision loss.11,45 PDT per se may result in occlusion of CNV, but is not thought to regulate molecular signals (angiogenic stimuli) that mediate the neovascular process; this failing may be the reason for the common recurrence of CNV. Several antiangiogenic agents are now under investigation for another alternative or additional
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a
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c Fig. 5. NPe6 and indocyanine green angiography of the experimental CNV. NPe6 angiogram at seven minutes 30 seconds (a), 20 minutes (b), and 60 minutes (c). The dye accumulation grew intensive along the time course. Note retinal vessels silhouetted negatively against the background fluorescence of CNV 20 minutes after the dye injection.
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f Fig. 5. NPe6 and indocyanine green angiography of the experimental CNV. Indocyanine green angiogram at seven minutes 30 seconds (d), 20 minutes (e), and 60 minutes (f). Notice the strong resemblance in accumulation and retention patterns of these two dyes. (Reproduced from Mori et al.42 by courtesy of the publisher.)
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Controversial aspects of photodynamic therapy modality for subfoveal CNV. Interferon alpha 2a causes dramatic involution of hemangiomas and inhibits iris neovascularization in a model of ischemic retinopathy.46,47 However, a multicenter, randomized, placebo-controlled trial demonstrated that patients with CNV who received interferon alpha 2a did not have any involution of CNV and, at the close of the study, had worse vision than those treated with a placebo.48 Vascular endothelial growth factor (VEGF) is an endothelial cell mitogen with a central role in ocular angiogenesis.49-51 Inhibition of VEGF, using antiVEGF antibodies or soluble receptors, can prevent the development of experimental iris or retinal neovascularization.52,53 VEGF kinase inhibitors, which block VEGF signalling, prevent the development of retinal neovascularization and CNV.54 These facts suggest that anti-VEGF therapy may play a role in the prevention of CNV or its recurrence after PDT. Phase I clinical trials testing the safety and tolerability of intraocular injections of an aptamer that binds VEGF or an anti-VEGF antibody have been completed, and phase II trials are being planned. Purported endogenous inhibitors of angiogenesis have also been described, including angiostatin,55 endostatin,56 antithrombin III,57 pigment epitheliumderived factor (PEDF),58 etc. However, their role, if any, in the development of retinal neovascularization and CNV is unknown. Recently, Mori et al. used adenoviral vectors to demonstrate that two of these proteins, endostatin59 and PEDF,60 inhibit ocular neovascularization. These studies also provided proof of the concept for the use of gene transfer to treat ocular neovascularization. Gene transfer offers a means for local delivery of a therapeutic agent without systemic side-effects or repeated intraocular injections (because of the sustained release of the protein). Mori et al. have also shown that, in two ocular neovascularization models, transgenic mice expressing VEGF in photoreceptors and a laser-induced CNV model, PEDF gene transfer in eyes with already established neovascularization caused regression of the neovascularization.61 This is the first demonstration of a pharmacological treatment that causes regression of ocular neovascularization, and could potentially be applied to the many patients who present with subfoveal CNV. However, adenoviral vectors have features that may limit their use in humans, including some evidence of toxicity and decreasing transgene expression to low levels over the course of a few months. Patients with AMD are at risk for the development or recurrence of CNV for many years, and long-term treatment is needed. Prolonged intraocular expression has been achieved with adeno-associated viral vectors. Actually, it has also reported that adeno-associated viral vector mediated gene transfer of PEDF inhibits CNV development.62 Phase I clinical trials of gene therapy expressing PEDF are now being planned.63 It is now expected that several drugs or viral vectors for anti-
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225 angiogenic therapy will be able to overcome the problems of PDT and will present an additional or alternative modality for AMD patients. Conclusions Photodynamic therapy offers new hope for patients with AMD and other diseases associated with CNV. The primary mechanism of action is selective occlusion of neovascular vessels following activation with a low-energy laser of the appropriate wavelength. While verteporfin is the only drug approved for treatment in humans, photosensitizers such as ATX-10 and NPe6 have shown great promise. Future trends include the use of selective antiangiogenic agents and gene therapy to control CNV, possibly in conjunction with PDT. References 1. Miller JW, Schmidt-Erfurth U, Sickenberg M et al: Photodynamic therapy with verteporfin for choroidal neovascularization caused by age-related macular degeneration: results of a single treatment in a phase 1 and 2 study. Arch Ophthalmol 117:1161-1173, 1999 2. Schmidt-Erfurth U, Miller JW, Sickenberg M et al: Photodynamic therapy with verteporfin for choroidal neovascularization caused by age-related macular degeneration: results of retreatments in a phase 1 and 2 study. Arch Ophthalmol 117:1177-1187, 1999 3. Treatment of Age-related Macular Degeneration with Photodynamic Therapy (TAP) Study Group: Photodynamic therapy of subfoveal choroidal neovascularization in agerelated macular degeneration with verteporfin: one-year results of 2 randomized clinical trials-TAP report 1. Arch Ophthalmol 117:1329-1345, 1999 4. Treatment of Age-related Macular Degeneration with Photodynamic Therapy (TAP) Study Group: Photodynamic therapy of subfoveal choroidal neovascularization in agerelated macular degeneration with verteporfin: two-year results of 2 randomized clinical trials-TAP report 2. Arch Ophthalmol 119:198-207, 2001 5. Dougherty TJ, Kaufman J, Goldfarb A, Weishaupt K, Boyle D, Mittleman A: Photoradiation therapy for the treatment of malignant tumors. Cancer Res 38:2628-2635, 1978 6. Grossweiner LI: The Science of Phototherapy, pp 27-49, 139-155. Boca Raton, FL: CRC Press 1994 7. Henderson BW, Dougherty TJ: How does photodynamic therapy work? Photochem Photobiol 55:145-157, 1992 8. Reed MWR, Miller FN, Wieman TJ et al: The effect of photodynamic therapy on the microcirculation. J Surg Res 45:452-459, 1988 9. Mori K, Yoneya S, Ohta M, Sano A, Sonoda M, Kaneda A, Sato Y: Potential of photodynamic therapy with a secondgeneration sensitizer: mono-L-aspartyl chlorin e6. J Jpn Ophthalmol Soc 101:134-140, 1997 10. Peyman GA, Kazi AA, Unal M, Khoobehi B, Yoneya S, Mori K, Moshfeghi DM: Problems with and pitfalls of photodynamic therapy. Ophthalmology 107:29-35, 2000 11. Nakashizuka T, Mori K, Hayashi N, Anzail K, Kanail K, Yoneya S, Moshfeghi DM, Peyman GA: Retreatment effect of NPe6 photodynamic therapy on the normal primate macula. Retina 21:493-498, 2001 12. Dougherty TJ, Lawrence G, Kaufman JH et al: Photo-
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32. Peyman GA, Moshfeghi DM, Moshfeghi A, Khoobehi B, Primbs GB, Doiron DR, Crean DH: Photodynamic therapy for choriocapillaris using tin ethyl etiopurpurin (SnET2). Ophthalmic Surg Lasers 28:409-417, 1997 33. Obana A, Gohto Y, Kanai M, Nakajima S, Kaneda K, Miki T: Selective photodynamic effects of the new photosensitizer ATX-S10 (Na) on choroidal neovascularization in monkeys. Arch Ophthalmol 118:650-658, 2000 34. Allen RP, Tharratt RS, Volz W, Senders C, Donald P: The toxicology and efficacy of NPe6 in man with superficial malignancies. Fifth International Photodynamic Association Biennial Meeting (Abstract) 54, 1994 35. Volz W, Allen R: Cutaneous phototoxicity of NPe6 in man. In: Spinelli P, Fante MD, Marchesini R (eds) Photodynamic Therapy and Biomedical Lasers, pp 446-448. Amsterdam: Elsevier 1992 36. Kessel D, Allen R: Determinants of localization by secondgeneration PDT sensitizers. In: Spinelli P, Fante MD, Marchesini R (eds) Photodynamic Therapy and Biomedical Lasers, pp 526-530. Amsterdam: Elsevier 1992 37. Roberts WG, Hasan T: Role of neovasculature and vascular permeability on the tumor retention of photodynamic agents. Cancer Res 52:924-930, 1992 38. Mori K, Yoneya S, Ohta M, Sano A, Anzai K, Peyman GA, Moshfeghi DM: Angiographic and histologic effects of fundus photodynamic therapy with a hydrophilic sensitizer; mono-L-aspartyl chlorin e6. Ophthalmology 106:1384-1391, 1999 39. Roberts WG, Shiau F-Y, Nelson JS, Smith KM, Berns MW: In vitro characterization of monoaspartyl chlorine e6 and diaspartyl chlorine e6 for photodynamic therapy. J Nat Cancer Inst 80:330-336, 1988 40. Roberts WG, Berns MW: In vitro photosensitization I: cellular uptake and subcellular localization of mono-L-aspartyl chlorine e6, chloro-aluminum sulfonated phthalocyanine, and Photofrin II. Lasers Surg Med 9:90-101, 1989 41. Roberts WG, Liaw LHL, Berns MW: In vitro photosensitization II: an electron microscopy study of cellular destruction with mono-L-aspartyl chlorine e6 and Photofrin II. Lasers Surg Med 9:102-108, 1989 42. Mori K, Yoneya S, Anzail K, Kabasawa S, Sodeyama T, Peyman GA, Moshfeghi DM: Photodynamic therapy of experimental choroidal neovascularization with a hydrophilic sensitizer mono-L-aspartyl chlorin e6. Retina 21:499-508, 2001 43. Yoneya S, Saito T, Komatsu Y, Kayama I, Takahashi K, Duvoll-Young J: Binding properties of indocyanine green in human blood. Invest Ophthalmol Vis Sci 39:1286-1290, 1998 44. Yannuzzi LA, Flower RW, Slakter JS: Indocyanine Green Angiography, pp 2-17, 46-49. St Louis, MO: CV Mosby 1997 45. Renno RZ, Delori FC, Holzer RA, Gragoudas, ES, Miller JW: Photodynamic therapy using Lu-tex induces apoptosis in vitro, and its effect is potentiated by angiostatin in retinal capillary endothelial cells. Invest Ophthalmol Vis Sci 41: 3963-3971, 2000 46. Ezekowiyz RAB, Mulliken JB, Folkman J: Interferon alpha2a therapy for life-threatening hemangioma of infancy. N Engl J Med 326:1456-1463, 1992 47. Miller JW, Stinson W, Folkman J: Regression of experimental iris neovascularization with systemic alpha-interferon. Ophthalmology 100:9-14, 1993 48. The Pharmacologic Treatment for Macular Degeneration Study Group: Interferon alpha-2a is ineffective for patients with choroidal neovascularization secondary to age-related macular degeneration: results of a prospective randomized placebo-controlled clinical trial. Arch Ophthalmol 115:865872, 1997
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Controversial aspects of photodynamic therapy 49. Aiello LP, Avery RL, Arrigg PG, Keyt BA, Jampel HD, Shah ST, Pasquale LR, Thieme H, Iwamoto MA, Park JE, Nguyen MS, Aiello LM, Ferrara N, King GL: Vascular endothelial growth factor in ocular fluid of patients with diabetic retinopathy and other retinal disorders. N Engl J Med 331:1480-1487, 1994 50. D’Amore PA: Mechanisms of retinal and choroidal neovascularization. Invest Ophthalmol Vis Sci 35:3974-3978, 1994 51. Tobe T, Okamoto N, Vinores MA, Derevjanik NL, Vinores SA, Zack DJ, Campochiaro PA: Evolution of neovascularization in mice with overexpression of vascular endothelial growth factor in photoreceptors. Invest Ophthalmol Vis Sci 39:180-188, 1998 52. Aiello L, Pierce E, Foley H et al: Suppression of retinal neovascularization in vivo by inhibition of vascular endothelial growth factor (VEGF) using soluble VEGF-receptor chimeric proteins. Proc Nat Acad Sci US 92:1045710461, 1995 53. Adamis AP, Shima DT, Tolentino MJ, Gragoudas ES, Ferrara N, Folkman J et al: Inhibition of vascular endothelial growth factor prevents retinal ischemia associated iris neovascularization in non human primate. Arch Ophthalmol 114:66-71, 1996 54. Seo M-S, Kwak N, Ozaki H, Yamada H, Okamoto N, Fabbro D, Hofmann F, Wood JM, Campochiaro PA: Dramatic inhibition of retinal and choroidal neovascularization by oral administration of a kinase inhibitor. Am J Pathol 154:17431753, 1999 55. O’Reilly MS, Holmgren S, Shing Y, Chen C, Rosenthal RA, Moses M, Lane WS, Cao Y, Sage HE, Folkman J: Angiostatin: a novel angiogenesis inhibitor that mediates the suppression of metastases by a Lewis lung carcinoma. Cell 79:315-328, 1994
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227 56. O’Reilly MS, Boehm T, Shing Y, Fukai N, Vasios G, Lane WS, Flynn E, Birknead JR, Olsen BR, Folkman J: Endostatin: an endogenous inhibitor of angiogenesis and tumor growth. Cell 88:277-285, 1997 57. O’Reilly MS, Pirie-Sheperd S, Lane WS, Folkman J: Antiangiogenic activity of the cleaved conformation of the serpin antithrombin. Science 285:1926-1928, 1999 58. Dawson DW, Volpert OV, Gillis P, Crawford SE, Xu HJ, Benedict W, Bouck NP: Pigment epithelium-derived factor: a potent inhibitor of angiogenesis. Science 285:245-248, 1999 59. Mori K, Duh E, Gehlbach P, Ando N, Takashi K, Pearlman J et al: Pigment epithelium-derived factor inhibits retinal and choroidal neovascularization. J Cell Physiol 188:253263, 2001 60. Mori K, Ando A, Gehlbach P, Nesbitt D, Takahashi K, Goldsteen D et al: Inhibition of choroidal neovascularization by intravenous injection of adenoviral vectors expressing secretable endostatin. Am J Pathol 159:313-320, 2001 61. Mori K, Gehlbach P, Ando A et al: Regression of ocular neovascularization by increased expression of pigment epithelium-derived factor. Invest Ophthalmol Vis Sci 43: 2002 (in press) 62. Mori K, Gehlbach P, Yamamoto S et al: AAV-mediated gene transfer of pigment epithelium-derived factor inhibits choroidal neovascularization. Invest Ophthalmol Vis Sci 43: 1994-2000, 2002 63. Rasmussen H, Chu KW, Campochiaro P, Gehlbach PL, Haller JA, Handa JT, Nguyen QD, Sung JU: Clinical protocol: an open-label, phase I, single administration, dose-escalation study of ADGVPEDF.11D (ADPEDF) in neovascular age-related macular degeneration (AMD). Hum Genet Ther 12:2029-2032, 2001
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Lasers in diabetes Robert A. Stoltz and Alexander J. Brucker Scheie Eye Institute/University of Pennsylvania, Philadelphia, PA, USA
Keywords: diabetes, diabetic retinopathy, rubeosis iridis, proliferative diabetic retinopathy, macular edema, laser treatment, classification, complications
Introduction Since the first direct evidence of retinal vascular abnormalities in diabetic patients was shown by Ashton in 1950,1 there have been continuing efforts to alleviate this visually debilitating condition. In 1952, Luft and coworkers2 carried out a hypophysectomy in the hope of ameliorating the vascular complications of diabetes, and, in 1959, photocoagulation for the treatment of diabetic retinopathy was first reported by Meyer-Schwickerath,3 who used a xenon arc photocoagulator to treat new vessels on the surface of the retina directly. During the 1970s, further treatments were developed (such as argon laser photocoagulation and pars plana vitrectomy) and their efficacy in preserving vision in patients with diabetic retinopathy examined in several landmark clinical trials which spanned the ensuing two decades. The Diabetic Retinopathy Study (DRS) of 1976, the first of such trials, showed that the rate of severe visual loss in high-risk proliferative diabetic retinopathy could be reduced by as much as 60% following the timely application of panretinal laser photocoagulation therapy.4 Subsequently, results from the Early Treatment Diabetic Retinopathy Study (ETDRS) demonstrated that focal laser photocoagulation treatment to the macula could substantially reduce the risk of visual acuity loss in patients with clinically significant diabetic macular edema.5 Diabetic retinopathy: classification Diabetic retinopathy exists as a spectrum of findings ranging from mild to severe (Table 1). Diabetic retinopathy is broadly classified as nonproliferative diabetic retinopathy (NPDR), formerly known as
background diabetic retinopathy (BDR), and proliferative diabetic retinopathy (PDR). The retinal microvascular changes that occur in NPDR are limited to the confines of the retina, whereas PDR is characterized by the growth and extension of new vessels from the retina, beyond the internal limiting membrane, and out onto the posterior hyaloid surface. The earliest funduscopic changes noted in diabetic retinopathy include retinal microaneurysms and dotblot hemorrhages. These are often most pronounced in the temporal macular region. Microaneurysms tend to cluster near zones of capillary nonperfusion, and their rate of formation appears to be an indication of the severity of diabetic retinopathy.6 Retinal hemorrhages are generally a reliable indicator of the severity of nonproliferative retinopathy. Acceleration of retinal capillary abnormalities eventually affects adjacent arterioles, resulting in arteriolar closure and discrete areas of capillary dropout or nonperfusion. Capillary closure with retinal ischemia leads to nerve-fiber layer infarcts (cotton wool spots), intraretinal microvascular abnormalities (IRMA), and venous beading. The DRS defined preproliferative diabetic retinopathy as any three of the following: nerve fiber layer infarcts, intraretinal microvascular abnormalities, venous beading, and at least moderately severe retinal hemorrhages and/or microaneurysmal formation.7 However, data from the ETDRS de-emphasized the significance of nerve fiber layer infarcts, finding that the severity of intraretinal microvascular abnormalities, hemorrhages and/or microaneurysms, and venous beading were the most important factors in predicting progression to proliferative retinopathy.8 Progressive capillary closure and retinal ischemia herald the development of proliferative retinopathy. Retinal neovascularization in diabetic retinopathy originates either from the op-
Address for correspondence: Alexander J. Brucker, MD, Scheie Eye Institute/University of Pennsylvania, 51 North 39th Street Philadelphia, PA 19104, USA
Lasers in Ophthalmology – Basic, Diagnostic and Surgical Aspects, pp. 229–240 edited by F. Fankhauser and S. Kwasniewska © 2003 Kugler Publications, The Hague, The Netherlands
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Table 1. Classification of severity of diabetic retinopathy Severity
Lesions present
Nonproliferative No retinopathy Microaneurysms only Mild NPDR, venous loops, or both Moderate NPDR Severe NPDR
Very severe NPDR Proliferative PDR without HRC PDR with HRC
No retinal lesions No lesions other than microaneurysms Microaneurysms plus retinal hemorrhage, hard exudate Mild NPDR plus cotton wool spots and/or IRMA Presence of one of the following features: microaneurysms plus venous beading and/or H/MS ≥ standard photograph 2A in four quadrants, or marked venous beading in two or more quadrants, or moderate IRMA in one or more quadrants Two or more of the above features described in severe NPDR
New vessels and/or fibrous proliferations; or preretinal and/or vitreous hemorrhage NVD ≥ standard photograph 10A; or less extensive NVD, if vitreous or preretinal hemorrhage is present; or NVE ≥ half disc area, if vitreous or preretinal hemorrhage is present Extensive vitreous hemorrhage precluding grading, retinal detachment involving the macula, or phthisis bulbi or enucleation secondary to a complication of diabetic retinopathy
Advanced PDR
NPDR: nonproliferative diabetic retinopathy; IRMA: intraretinal microvascular abnormalities; H/MA: hemorrhages and/or microaneurysms; PDR: proliferative diabetic retinopathy; HRC: high-risk characteristics; NVD: new vessels on or within one disc diameter of the optic disc; NVE: new vessels elsewhere
a.
b.
c. Fig. 1. Proliferative diabetic retinopathy with high risk characteristics (HRC). a: NVD ≥ one-quarter to one-third disc diameter, without associated preretinal or vitreous hemorrhage. b: Any neovascularization located on or within one disc diameter of the disc (NVD) associated with preretinal or vitreous hemorrhage. c: Neovascularization of the retina elsewhere (NVE) associated with preretinal or vitreous hemorrhage.
tic disc (NVD) or retinal surface elsewhere (NVE) and proliferates on the posterior cortical vitreous (Fig. 1). New vessels proliferate along the posterior hyaloid surface and cause visual loss by either bleeding into the vitreous, or causing contraction of the
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vitreous, thereby resulting in a tractional retinal detachment. Neovascularization can regress rarely spontaneously or following panretinal photocoagulation, forming residual, relatively avascular sheets of fibrovascular tissue along the posterior hyaloid surface.
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Lasers in diabetes The presence of microaneurysms are the earliest signs of diabetic retinopathy. They may be present at any stage of retinopathy from nonproliferative to proliferative. From these microaneurysms, leakage of fluid occurs within the retina resulting in edema. Macular edema is the most frequent cause of visual impairment in patients with diabetic retinopathy. The breakdown of the inner blood-retinal barrier can be associated with both nonproliferative and proliferative diabetic retinopathy. This excessive vascular permeability resulting in the leakage of fluid and plasma constituents, such as lipoproteins into the retina, leads to thickening of the retina. When thickening involves or threatens the center of the fovea, there is a higher risk of visual loss. In the ETDRS, the three-year risk of moderate visual loss (a doubling of the initial visual angle or a decrease of three lines or more on a logarithmic visual acuity chart) was 32%. The ETDRS classified macular edema as ‘clinically significant macular edema’ (CSME) if any
231 of the following features were present: (1) thickening of the retina at or within 500 µm of the center of the macula; (2) hard exudates at or within 500 µm of the center of the macula, if associated with thickening of the adjacent retina; or (3) a zone or zones of retinal thickening one disc area or larger, any part of which is within one disc diameter of the center of the macula (Fig. 2). Epidemiology of diabetes and diabetic retinopathy Diabetes mellitus affects approximately 17-18 million persons in the USA, and the majority (≈ 90%) of those patients have type 2 diabetes. According to the American Diabetes Association, of the roughly 17 million people who have diabetes, 5.9 million are unfortunately unaware that they have impaired glucose tolerance or definite diabetes. Diagnosed diabetes is most prevalent in the middle-aged and elderly
a.
b. Fig. 2. Clinically significant macular edema. a: Retinal thickening at or within 500 µm of the center of the macular. b: Hard exudates at or within 500 µm of the center of the macula with adjacent retinal thickening.
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c. Fig. 2. Clinically significant macular edema. c: Retinal thickening one disc area or larger in size located within one disc diameter of the center of the macula. (Courtesy of ETDRS.)
populations, affecting 6% of people aged 45-64 years and 18.4% of those aged 65 years or older, but only 1.5% of those aged 18-44 years.9,10 The Wisconsin Epidemiologic Study of Diabetic Retinopathy (WESDR), a population-based study of diabetic patients in southern Wisconsin, provides the best data on the epidemiology of diabetic retinopathy and the risk factors associated with the development of diabetic retinopathy. According to the results of this study, retinopathy, either nonproliferative or proliferative, was seen in 13% of the patients whose age at diagnosis of diabetes was less than 30 years, had less than a five-year duration of diabetes, and were taking insulin at the time of the examination (presumably type 1). In contrast, up to 90% of patients with a 10-to-15-year duration of diabetes had some form of diabetic retinopathy. PDR is present in approximately 25% of patients with type 1 diabetes and a 15-year duration of disease.11 For patients with an onset of diabetes at 30 years of age or older (e.g., those with type 2 diabetes) and a duration of diabetes less than five years, 40% of insulin-requiring and 24% of non-insulin requiring diabetics have retinopathy.12 These rates increase to 84 and 53%, respectively, with an increased duration of diabetes of 15-19 years. PDR develops in 2% of patients with type 2 diabetes and a duration of less than five years, and in 25% of patients with a duration of 25 or more years. For diabetic macular edema, again the best epidemiological data was obtained from the WESDR. It was found that the prevalence of macular edema did not vary as much by diabetes type. The prevalence of diabetic macular edema was approximately 18-20% in patients with either type 1 or type 2 diabetes, but increased significantly with the overall severity of diabetic retinopathy and duration of diabetes.13,14
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The WESDR data discussed above suggested that the major factors associated with increased severity of retinopathy include patient age at diagnosis, increasing duration of diabetes, type 1 diabetes, and insulin-dependent diabetes. Other important factors include degree of metabolic control, blood pressure, elevated serum lipid levels, pregnancy and renal disease.
Laser therapy for proliferative diabetic retinopathy: historical background The DRS15-19 was initiated in 1972 and completed in 1979, and provided important information concerning the understanding and treatment of proliferative diabetic retinopathy. This randomized, controlled, clinical trial determined the natural history of the disease without photocoagulation by following randomly selected control eyes, and confirmed the previously suggested beneficial treatment results with xenon arc and argon laser photocoagulation by examining a large number of patients at regular intervals for an extended time period. The DRS follow-up examinations were terminated in June 1979 after the study’s major goals had been achieved.19 In this study, panretinal photocoagulation was applied from around the disc and major arcades to or beyond the equator in one or several treatment sessions. The xenon technique used 400-800, 3° burns or 200-400, 4.5° burns. The argon blue-green technique used 0.1 second-, moderate-to-heavy-intensity retinal burns. The scatter pattern consisted of 800-1600, 500-µm burns, or 500-1000, 1000-µm burns with a spacing of about one burn width apart. Disc neovascularization or elevated neovascularization elsewhere was treated with scatter argon photocoagulation in an attempt to close the new vessels.
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Lasers in diabetes Flat neovascularization elsewhere was directly treated with either argon or xenon photocoagulation. The DRS found that photocoagulation reduced the risk of severe visual loss (vision 20/80 or worse for two consecutive follow-up visits) by at least 50%.19 The overall risk of severe visual loss with proliferative diabetic retinopathy at the two-year follow-up examination was 16% in the control eyes compared to 6% in the treated eyes. With DRS high-risk characteristics, this risk increased to 26% in the control eyes and 11% in the treated eyes. However, when high-risk characteristics were absent, the risk of severe visual loss decreased to 7% of the control eyes and 3% of the treated eyes. In a separate study by Doft et al., the beneficial effect of panretinal photocoagulation persisted through 15 years of followup.21 Although argon and xenon photocoagulation were equally effective in preventing severe visual loss, argon emerged as the preferred treatment modality due to fewer harmful side-effects.19,23 Twenty-five percent of the xenon-treated eyes had visual field loss (Goldmann IV 4e test object) compared to 5% of eyes treated with argon photocoagulation, and 11% of the xenon-treated eyes experienced a permanent visual loss of two or more lines compared to 3% of the eyes treated with argon photocoagulation. Moreover, there was an increase in macular damage from vitreoretinal traction following xenon laser treatment. Eyes receiving panretinal photocoagulation sometimes developed decreased vision, especially eyes with preexisting macular edema or in those eyes receiving heavy xenon panretinal photocoagulation. It was found that panretinal photocoagulation occasionally aggravated macular edema.24 Eyes treated with scatter argon photocoagulation (PRP) were approximately 60% more likely to lose two or more lines of visual acuity after photocoagulation than untreated eyes and xenon-treated eyes were approximately five times more likely to lose this much vision compared to untreated eyes within six weeks of photocoagulation. Therefore, the following recommendations were made to decrease photocoagulation-induced macular edema: (1) treat macular edema with focal photocoagulation prior to initiating panretinal photocoagulation; (2) avoid intense panretinal photocoagulation burns; and (3) divide panretinal photocoagulation into several treatment sessions.24 One arm of the ETDRS was designed to determine the best time to initiate panretinal photocoagulation in patients with diabetic retinopathy. The ETDRS concluded that scatter photocoagulation is not recommended for eyes with mild or moderate non-proliferative diabetic retinopathy, provided that careful follow-up can be maintained. The five-year rate of severe visual loss in this group was 1-3%. On the other hand, scatter photocoagulation should be considered when retinopathy is more severe and usually should not be delayed if the eye has reached the high-risk proliferative stage, as there was a 47% risk of severe visual loss at five years.25 The
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233 ETDRS data also failed to support the concept of staged scatter photocoagulation.25,26 The beneficial effects of panretinal photocoagulation were found to be independent of the wavelength used.22,27-36 Rubeosis iridis Several reports have demonstrated that panretinal laser photocoagulation is effective in preventing the development of rubeosis and in causing the regression of pre-existing iris neovascularization.37-42 The best results are obtained when treatment is given prior to the development of neovascular glaucoma and extensive peripheral anterior synechia. A cyclodestructive and/or seton procedure may be required if panretinal laser photocoagulation fails to control rubeosis and neovascular glaucoma. In 1983, Pavan et al.43 reported on the results of a prospective, randomized, controlled study that evaluated the benefit of panretinal photocoagulation for rubeosis iridis secondary to proliferative diabetic retinopathy. The rubeosis was documented using iris angiography before and five to seven weeks after panretinal photocoagulation. Panretinal photocoagulation was effective in causing the regression of severe rubeosis iridis (at least 0.5 x 0.5 millimeters in area), which improved in 11 (73%) of 15 treated eyes compared with two (18%) of 11 untreated eyes. Laser guidelines Panretinal photocoagulation Panretinal photocoagulation is indicated for any eye with DRS high-risk characteristics, rubeosis iridis, or neovascular glaucoma. Photocoagulation of clinically significant macular edema should be considered before instituting scatter (panretinal) treatment. Scatter photocoagulation burns are placed from just within the vascular arcades to anterior to the equator. Tight (confluent) photocoagulation is applied directly to areas of flat peripheral neovascularization. The laser settings for scatter photocoagulation are given in Table 2.20 All wavelengths appear to be equally effective in inducing regression of proliferative disease. The red and diode wavelengths, however, are better able to penetrate cataracts or vitreous hemorrhage than the shorter wavelengths. The photocoagulation burns are placed approximately one burn width apart. The anterior edge of treatment should extend to or beyond the equator. The posterior edge of treatment includes an oval area that extends 500 µm nasal to the optic disc margin, along the temporal arcades, and two disc diameters temporal to the macular center. Use of a panfunduscopic contact lens, 3-mirror contact lens, or indirect ophthalmoscope laser delivery system is used for panretinal photocoagulation. It should be kept in mind that treatment over retinal
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Table 2. Laser settings for panretinal photocoagulation Wavelength:
argon green, Nd:YAG green, dye yellow, red, or diode the red or diode wavelengths may be useful when cataract, intraretinal hemorrhage, or vitreous hemorrhage are present Duration: 0.1-0.5 seconds Retinal spot size*: 500 µm with three-mirror contact lens; 200 µm with panfunduscopic contact lens Intensity: moderately intense retinal burn *
Laser spot diameter at the retina versus the photocoagulator spot size setting is dependent upon the spot magnification induced by the specific contact lens (adapted from Bloom and Brucker20 by courtesy of the publisher)
hemorrhage, major retinal vessels, or chorioretinal scars should be avoided. Directly treating retinal hemorrhage can result in unnecessary inner retinal damage. When panretinal photocoagulation burns are placed over retinal vessels, vascular occlusion and rupture may potentially occur, albeit this is a rare complication. Overly intense burns can occur when photocoagulation burns are placed over pigmented chorioretinal scars, causing visual field loss.20 Treatment may extend within the vascular arcades within 3000 µm of the macular center for retinal neovascularization. Such localized scatter photocoagulation is performed with 200-µm burns when treating 5001500 µm from the center of the macula. Treatment should not extend closer than 500 µm from the macular center or disc margin. Care should be taken to avoid photocoagulating within the papillomacular bundle (Fig. 3). Panretinal photocoagulation should be completed in several treatment sessions over a three- to sixweek period.21 Furthermore, dividing panretinal laser into multiple sessions decreases the risks of macular edema, exudative retinal detachment, choroidal detachment with secondary angle-closure glaucoma. The order in which the retina is treated is optional. The inferior retina is usually treated first, as vitreous hemorrhage, should it occur, tends to settle inferiorly, making it difficult to later photocoagulate this region. Although the DRS treatment protocol for scatter photocoagulation consisted of 800-1600, 500-µm burns, a complete treatment generally consists of approximately 1800-2200, 500-µm retinal burns. There are no postoperative physical restrictions following panretinal photocoagulation. Patients should be examined within four months of completing panretinal photocoagulation treatment. The indications for additional treatment are multifactorial and must be individualized. Factors that favor additional photocoagulation include enlarging neovascularization, increasing activity of the neovascularization, e.g., the formation of tight vascular networks with a paucity of accompanying fibrous tissue, and in increase in the frequency or extent of vitreous hemorrhage if active neovascularization is present.20 Additional fill-in laser treatment may consist of scatter photocoagulation anterior to, poste-
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a.
b. Fig. 3. a: ‘Full scatter’ panretinal photocoagulation therapy. Laser spots are spaced approximately one spot-width apart. The posterior margin of the treatment is about two disc diameters above, temporal, and below the center of the macula. b: Severe PDR prior to scatter photocoagulation (left), and 18 months following full scatter panretinal photocoagulation with additional ‘fill-in’ laser treatments (right). There is full regression of the neovascularization.
rior to, or between prior laser scars, as well as local photocoagulation to areas of flat peripheral retinal neovascualarization. For this, treatment consists of nearly confluent scatter photocoagulation applied directly to flat peripheral neovascularization using similar laser settings as that used for panretinal photocoagulation (vide supra). Treatment of diabetic macular edema Diabetic macular edema may be caused by either focal or diffuse leakage. Intraretinal edema results from leakage of fluid through damaged retinal endothelial cells. Focal retinal thickening is usually caused by leaking microaneurysms (Fig. 4). Diffuse retinal thickening is caused by a generalized breakdown of the inner blood-retinal barrier, i.e., leaking microaneurysms, intraretinal microvascular abnormalities, or short-capillary segments. In addition, a defective retinal pigment epithelial pump may also contribute to macular edema. Macular ische-mia is sometimes associated with macular edema and is a poor prognostic indicator for visual improvement with or without laser photocoagulation.44-47 Diabetic macular edema may be present at any level of retinopathy and is found in approximately
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a.
b. Fig. 4. a: Focal macular edema: a discrete area of retinal thickening with a hard exudate, temporal to the macular center (one disc area within one disc diameter of the fovea). b: Corresponding fluorescein angiogram showing clusters of microaneurysms and capillary telangiectasia.
10% of all diabetic patients.13 As the severity of the retinopathy increases, the proportion of eyes with macular edema also increases ranging from 3% in eyes with mild nonproliferative diabetic retinopathy to 38% with moderate to severe non-proliferative diabetic retinopathy and 71% with proliferative diabetic retinopathy.48 In clinic-based surveys, macular edema has been reported to be more frequent in older, adult-onset diabetic patients, 49 whereas population studies demonstrate a higher prevalence of macular edema in juvenile-onset diabetics.13 This discrepancy may be explained by the fact that olderonset diabetic patients are almost ten times more common in the general population. Photocoagulation treatment of diabetic macular edema with various wavelengths and treatment techniques has been used for a number of years with varying success.5,50-56 For example, Gupta et al.56 examined the efficacy of various wavelengths in the treatment of clinically significant macular edema. They compared argon green (514 nm), krypton red (647 nm), frequency-doubled Nd:YAG (532 nm), and diode (810 nm) for focal and/or grid laser photocoagulation treatment. Reduction or elimination of clinically significant macular edema was observed
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in 93.3% of argon-treated eyes, 88.5% in krypton red group, 92.9% with frequency-doubled Nd:YAG, and 84.8% with diode laser. Although there was no statistically significant difference between the groups, frequency-doubled Nd:YAG treated eyes appeared to have the advantage of requiring fewer re-treatments. Nevertheless, all studies suggest a beneficial effect in either improving or stabilizing visual acuity with laser photocoagulation. Based on the available literature, particularly the ETDRS results, laser treatment should be considered for eyes with clinically significant macular edema. Treatment guidelines Treatment should be considered for all microaneurysm 500-3000 µm from the macular center, thought to be causing clinically significant macular edema. Treatment is initially optional for leaking microaneurysms within 500 µm of the macular center. If macular edema persists on follow-up examination and if vision is worse than 20/40 with a good perifoveal capillary network, focal treatment up to 300 µm from the center of the macula should be considered.20
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Treatment settings for focal photocoagulation are given in Table 3. The desired endpoint of laser treatment of microaneurysms is a change in coloration (either whitening or blackening). Multiple treatments may be required for the same microaneurysm. For example, large microaneurysms (at least 40 µm in size) can usually be closed with several 50-75-µm burns. Although closure of smaller microaneurysms is often difficult or impossible, it may be facilitated by the following: (1) First place larger burns, i.e., 100-200 µm in size, over the microaneurysms. The induced outer retinal whitening prevents subsequent photocoagulation burns from penetrating into the retinal pigment epithelium. (2) Smaller (50-µm size) burns may then be able to close these microaneurysms. This technique may also be useful for follow-up treatment, as it avoids intense burns at the level of the retinal pigment epithelium caused by excessive uptake of laser energy from pigmented photocoagulation scars. Grid photocoagulation may also be used to treat clusters of small microaneurysms (vide infra). Small red spots seen clinically may be treated optionally if they are thought to be microaneurysms that do not fill on fundus fluorescein angiography. However, these should not be treated if they are thought to represent dot hemorrhages. Areas of diffuse retinal thickening 500-3000 µm from the macular center thought to be causing clinically significant macular edema, should be treated using grid pattern photocoagulation. Grid photocoagulation should be applied to retinal avascular zones and clusters of small microaneurysms 500-3000 µm from the center of the macula that are associated with clinically significant macular edema. The laser setting for grid photocoagulation are given in Table 4. The grid spacing should be as close as one burn width apart. Photocoagulation should remain more than 500 µm from the disc margin. Treatment within the papillomacular bundle is allowed if it remains more than 500 µm from the center of the macula. Photocoagulation is clearly beneficial for all types of clinically significant macular edema as defined by the ETDRS. However, treatment is most beneficial for eyes with more extensive macular thickening and greater degrees of central macular edema (Fig. 5).57 Immediate treatment must therefore be considered for retinal thickening at or within 500 µm of the macular center, as there is a significant risk of severe visual loss. For hard exudates at or within 500 µm of the macular center, if associated with thickening of the adjacent retina, treatment is less urgent if the visual acuity is normal. Laser is recommended if the risk of treatment causing visual loss appears small or if frequent follow-up examinations cannot be ensured. There is relatively low risk of visual loss from retinal thickening at least one disc area in extent, any part of which is within one disc diameter of the macular center. Although photocoagulation has been shown to be beneficial, it may be reasonable to follow these patients for progression before treating. This is particularly impor-
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Table 3. Laser settings for focal photocoagulation for diabetic macular edema Wavelength: Duration:
argon green, Nd:YAG green, dye yellow 0.1 seconds or less, with 0.05 seconds recommended when treating within 500 µm of the center of the macula Retinal spot size: 75-100 µm, with 50-75 µm recommended when treating within 500 µm of the center of the macula Intensity: whitening or darkening of the lesion (Adapted from Bloom and Brucker20 by courtesy of the publisher)
Table 4. Laser settings for focal/grid photocoagulation for diabetic macular edema Wavelength:
argon green, Nd:YAG green, dye yellow, red, or diode Duration: 0.1 seconds Retinal spot size: 100-200 µm (a 50- to 75-µm spot size may be used in highly thickened retina, since this will result in a larger, more diffuse retinal burn) recommended when treating within 500 µm of the center of the macula Intensity: light retinal burn (Adapted from Bloom and Brucker20 by courtesy of the publisher)
tant when the majority of vascular leakage is close to the macular center, increasing the risk of foveal damage from laser burns.20 Patients should be examined three to four months after treatment, and considered for additional focal and/or grid photocoagulation if clinically significant macular edema is present. The decision of whether to retreat must be tempered, however, by other factors. For example, a patient with 20/25 vision and microaneurysms on the edge of the foveal avascular zone causing persistent clinically significant macular edema but who has shown an improvement in vision and a decrease in macular hard exudate and thickening following treatment can probably be followed initially without re-treatment. Complications of laser treatment in diabetic retinopathy Laser photocoagulation of the retina for the treatment of diabetic retinopathy has undergone significant advances and refinements since its initial application by Meyer-Schwickerath in 1959. Xenon arc light coagulation and ruby lasers have been replaced by argon, krypton, and dye, and diode lasers. The new lasers allow for smaller, more precise burns and diverse delivery systems. Photocoagulation is a minimally invasive treatment modality; nevertheless, complications are possible and diverse depending on the procedure employed. Panretinal photocoagulation is a commonly performed and effective treatment for proliferative
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a.
b.
c. Fig. 5. a: Diffuse macular edema: retinal thickening and scattered hard exudates throughout the macula. b: Same eye four months after focal/grid laser photocoagulation. There is significantly less macular edema and exudate. However, typical circinate rings of hard exudate are still present, and repeat focal laser treatment was applied. c: Three months following the second focal laser treatment. There is residual exudate, but the edema has resolved.
diabetic retinopathy. The DRS suggests that photocoagulation for eyes at risk will reduce the risk of severe visual loss from PDR.19 Argon laser photocoagulation usually proceeds without complication, although reports of these are well-documented.58 Complications of PRP include damage to the cornea, iris, or lens;59-61 transient myopia and accommodative paresis;62 uveitis;63 pain; hemorrhage; vision loss; ciliochoroidal effusion/detachment;64 and el-
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237 evated intraocular pressure with or without angleclosure. The DRS Research Group demonstrated the beneficial effects PRP whereby between 800 and 1600, 500-µm argon laser burns were scattered throughout the peripheral retinal in a random pattern approximately one burn width apart. Since the results of that study were released, most retina specialists have noted that it is often necessary to treat patients who have proliferative retinopathy with increased numbers of photocoagulation burns in order to obtain neovascular tissue to regress, and have observed that in some patients with extensive disease it may become necessary to treat the entire peripheral retina. In addition, complications are usually noted to occur more frequently when PRP is performed during a single session versus multiple sessions.65 For example, Gentile et al.64 performed a prospective study using ultrasound biomicroscopy to determine the risk factors for ciliochoroidal effusion after PRP. They found that a threshold for the development of ciliochoroidal effusion after PRP exists and is dependent on three factors: (1) burn intensity, which is, in turn, dependent on the laser output parameters and the absorption qualities of the retina and choroid; (2) burn size and number, which corresponds to the retinal surface area treated; and (3) axial length, which corresponds to the total retinal surface area and when combined with the retinal surface area treated, represents the percentage of the retinal surface area treated. The incidence of some degree of cilioretinal effusion after PRP is between 59% and 90%.66,67 These complications generally resolve within 14 days. Moriarty et al.68 examined the breakdown of the blood-aqueous barrier using laser photometry after argon laser panretinal photocoagulation for proliferative diabetic retinopathy. These authors found a significant increase in aqueous flare at three, 24, and 48 hours following PRP with an argon green laser (2000 burns, 0.1-second exposure, 200-µm spot size). More pigmented irides underwent a greater breakdown of the blood-aqueous-barrier than the blue, paler irides and this is attributed to a possible increased absorption by iris pigment responsible for the increased thermal effects. In fact, there is also an increased absorption and scattering within the anterior segment of shorter wavelengths which increases with age as flavin pigments accumulate in the lens.69 With respect to laser treatment for diabetic macular edema, the ETDRS also showed that focal photocoagulation of clinically significant macular edema substantially reduces the risk of visual loss. Identification of the patient’s fixation point is important to avoid foveal burns. Treatment of diabetic macular edema can be challenging, particularly because the foveal reflex often is obscured by edema, blood or exudates. Fortunately, the adverse effects of photocoagulation for diabetic macular edema thus far have been rare. These complications include inadvertent
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foveal photocoagulation, lipid precipitation in the fovea, subretinal or epiretinal fibrosis, increased foveal ischemia, choroidal or retinal hemorrhages or both, visual field defects/scotomas, expansion of laser scars, perforation of Bruch’s membrane, and choroidal neovascularization. Delivery of laser energy using small spot sizes, short durations, and high power has been proposed to increase the risk of perforation of Bruch’s membrane and development of choroidal neovascular membranes. It is important to keep in mind the treatment protocol of the ETDRS when treating patients for clinically significant macular edema. Excessive energy delivery should be avoided to any area in the macula, especially as the perifoveal area is approached. Foveal burns can be avoided by finding the fixation point by projecting a fluorescein angiogram, and using akinesia if cooperation is poor. Scotomas are best avoided by not treating near the foveal avascular zone, using low energy, and avoiding the use of blue-green wavelengths.53 Enlargement of the burns over time can cause scotomas. However, this is more common in myopic eyes, and a 5% incidence has been described with a krypton grid.70 Precipitation of hard exudates, foveal migration, and subretinal fibrosis can occur in eyes with severe, chronic edema and with aggressive treatment as has been observed in eyes with macular edema secondary to branch vein occlusion.71 In these instances, treatment in multiple sessions is advised to allow gradual fluid resolution. Moreover, increase in macular edema can occur with aggressive treatment and concomitant panretinal photocoagulation. Therefore, treatment of macular edema prior to PRP reduces this risk. Caution should be exercised in treating eyes with enlarged foveal avascular zones, because treatment can compromise the existing capillaries, leading to more ischemia. General recommendations to avoid complications include identifying the fixation point, treating with threshold laser burns sufficient to close microaneurysms, treating in multiple sessions if necessary, and avoiding treating directly over blood or fibrovascular tissue. Conclusions New laser modalities, expanding delivery systems, and novel applications of laser energy have vastly expanded our armamentarium for the treatment of diabetic eye disease in the 50 years since MeyerSchwickerath’s initial description of xenon arc light coagulation. The minimal invasiveness of laser treatment has significant appeal. Nevertheless, complications are possible, but many can be avoided with meticulous technique and experience. References 1. Ashton N: Injection of the retinal vascular system in the enucleated eye in diabetic retinopathy. Br J Ophthalmol 34:38-41, 1950
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2. Luft R, Olivecrona H, Sjögren B: Hypophysectomy in man. Nord Med 7:351-354, 1952 3. Meyer-Schwickerath G: Lichtkoagulation. Buch Augenarzt 33, 1959 4. The Diabetic Retinopathy Research Group: Preliminary report on effects of photocoagulation therapy. Am J Ophthalmol 81:383-396, 1976 5. Early Treatment Diabetic Retinopathy Study Research Group: Photocoagulation for diabetic macular edema: ETDRS report number 1. Arch Ophthalmol 103:1796-1806, 1985 6. Kohner EM, Sleightholm M: The Kroc Collaborative Study Group: Does microaneurysm count reflect severity of early diabetic retinopathy? Ophthalmology 93:586-589, 1986 7. Diabetic Retinopathy Study Research Group: Report 6: design, methods and baseline results. Invest Ophthalmol Vis Sci 21:149-209, 1981 8. Early Treatment Diabetic Retinopathy Study Research Group: Fluorescein angiographic risk factors for progression of diabetic retinopathy. ETDRS report number 13. Ophthalmology 98:834-840, 1991 9. Harris M: Diabetes in America: epidemiology and scope of the problem. Diabetes Care 21(Suppl 3):11-14, 1998 10. Harris M, Flega KM, Cowie CC, Eberhardt MS, Goldstein DE, Littler RR, Wiedmeyer HM, Byrd-Holt DD: Prevalence of diabetes, impaired fasting glucose and impaired glucose tolerance in US adults: the Third National Health and Nutrition Examination Survey, 1988-1994. Diabetes Care 21:518-524, 1998 11. Klein R, Klein BEK, Moss SE, Davis MD, DeMets DL: The Wisconsin Epidemiologic Study of Diabetic Retinopathy. III. Prevalence and risk of diabetic retinopathy when age at diagnosis is 30 or less years. Arch Ophthalmol 102: 520-526, 1984 12. Klein R, Klein BEK, Moss SE, Davis MD, DeMets DL: The Wisconsin Epidemiologic Study of Diabetic Retinopathy. III. Prevalence and risk of diabetic retinopathy when age at diagnosis is 30 or more years. Arch Ophthalmol 102:527-532, 1984 13. Klein R, Klein BEK, Moss SE, Davis MD, DeMets DL: The Wisconsin Epidemiologic Study of Diabetic Retinopathy: IV. Diabetic macular edema. Ophthalmology 91:14641474, 1984 14. Klein R, Moss SE, Klein BEK et al: The Wisconsin Epidemiologic Study of Diabetic Retinopathy. XI. The Incidence of macular edema. Ophthalmology 96:1501-1510, 1989 15. The Diabetic Retinopathy Study Research Group: Preliminary report on effects of photocoagulation therapy. Am J Ophthalmol 81:383-396, 1976 16. The Diabetic Retinopathy Study Research Group: Photocoagulation treatment of proliferative diabetic retinopathy: the second report of the Diabetic Retinopathy Study findings. Ophthalmology 85:82-106, 1978 17. The Diabetic Retinopathy Study Research Group: Four risk factors for severe visual loss in diabetic retinopathy: the third report from the Diabetic Retinopathy Study. Arch Ophthalmol 97:654-655, 1979 18. The Diabetic Retinopathy Study Research Group: Photocoagulation treatment of proliferative diabetic retinopathy: a short report of long-range results. Diabetic Retinopathy Study (DRS) Report Number 4. In: Waldhaus WK (ed) Diabetes 1979: Proceedings of the 10th Congress of the International Diabetes Federation, pp 789-794. Amsterdam: Excerpta Medica 1980 19. The Diabetic Retinopathy Study Research Group: Photocoagulation treatment of proliferative diabetic retinopathy: clinical application of Diabetic Retinopathy Study (DRS) findings, DRS Report Number 8. Ophthalmology 88:583600, 1981
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Lasers in diabetes 20. Bloom SM, Brucker AJ: Laser Surgery of the Posterior Segment, 2nd edn. Philadelphia/New York: Lippincott-Raven Publ 1997 21. Doft BH, Blankenship G: Single versus multiple treatment sessions of argon laser panretinal photocoagulation for proliferative diabetic retinopathy. Ophthalmology 89:772779, 1982 22. Blankenship GW: Fifteen-year argon laser and xenon photocoagulation results of Bascom Palmer Eye Institute’s patients participating in the Diabetic Retinopathy Study. Ophthalmology 3:125-128, 1991 23. The Diabetic Retinopathy Study Research Group: Photocoagulation treatment of proliferative diabetic retinopathy: relationship of adverse treatment effects to retinopathy severity. Dev Ophthalmol 2:248-261, 1981 24. Ferris FL III, Podgor MJ, Davis MD: The Diabetic Retinopathy Study Research Group: macular edema in diabetic retinopathy study patients: Diabetic Retinopathy Study report number 12. Ophthalmology 94:754-760, 1987 25. The Early Treatment Diabetic Retinopathy Study Research Group: Early photocoagulation for diabetic retinopathy: ETDRS report number 9. Ophthalmology 81:766-785, 1991 26. The Early Treatment Diabetic Retinopathy Study Research Group: Titrated scatter treatment in the ETDRS (reply, letter to the editor). Ophthalmology 98:1755-1756, 1991 27. Crick MDP, Chignell AH, Shilling JS: Argon laser vs. xenon arc photocoagulation in proliferative diabetic retinopathy. Trans Ophthalmol Soc UK 98:170-171, 1978 28. Schulenburg WE, Hamilton AM, Blach RK: A comparative study of argon laser and krypton laser in the treatment of diabetic optic disc neovascularization. Br J Ophthalmol 63:412-417, 1979 29. Hamilton AM, Townsend C, Khoury D et al: Xenon arc and argon laser photocoagulation in the treatment of diabetic disc neovascularization: Part I. Effect on disc vessels, visual fields, and visual acuity. Trans Ophthalmol Soc UK 101:8792, 1981 30. Plumb AP, Swan AV, Chignell AH, Shilling JS: A comparative trial of xenon arc and argon laser photocoagulation in the treatment of proliferative diabetic retinopathy. Br J Ophthalmol 66:213-218, 1982 31. Okun E, Johnston GP, Boniuk I et al: Xenon arc photocoagulation of proliferative diabetic retinopathy: a review of 2688 consecutive eyes in the format of the Diabetic Retinopathy Study. Ophthalmology 91:1458-1463, 1984 32. Blankenship GW, Gerke E, Battle JF: Red krypton and bluegreen argon laser diabetic panretinal photocoagulation. Graefe’s Arch Clin Exp Ophthalmol 227:364-368, 1989 33. Brancato R, Bandello F, Trabucchi G et al: Argon and diode laser photocoagulation in proliferative diabetic retinopathy: a preliminary report. Lasers Light Ophthalmol 3:233-237, 1989 34. Brancato R, Bandello F, Trabucchi G, Lattanzio R: Frequency-doubled Nd:YAG laser versus argon-green laser photocoagulation in proliferative diabetic retinopathy: a preliminary report. Lasers Light Ophthalmol 4:97-102, 1991 35. Bandello F, Brancato R, Trabucchi G et al: Diode versus argon-green laser panretinal photocoagulation in proliferative diabetic retinopathy: a randomized study in 44 eyes with a long follow-up time. Graefe’s Arch Clin Exp Ophthalmol 231:491-494, 1993 36. Krypton Argon Regression of Neovascularization Study Group: Randomized comparison of krypton versus argon scatter photocoagulation for diabetic disc neovascularization: the Krypton Argon Regression of Neovascularization Study report number 1. Ophthalmology 100:1655-1664, 1993 37. Little HL, Rosenthal AR, Dellaporta A, Jacobson DR: The effect of pan-retinal photocoagulation on rubeosis iridis. Am J Ophthalmol 81:804-809, 1976
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239 38. Laatikainen L: Preliminary report on effect of retinal panphotocoagulation on rubeosis and neovascular glaucoma. Br J Ophthalmol 61:278-284, 1977 39. Wand M, Dueker DK, Aiello LM, Grant WM: Effects of panretinal photocoagulation on rubeosis iridis, angle neovascularization, and neovascular glaucoma. Am J Ophthalmology 86:332-339, 1978 40. Jacobsen DR, Murphy RP, Rosenthal AR: The treatment of angle neovascularization with panretinal photocoagulation. Ophthalmology 86:1270-1275, 1979 41. Murphy RP, Egbert PR: Regression of iris neovascularization following panretinal photocoagulation. Arch Ophthalmol 97:700-702, 1979 42. Tasman W, Magargal LE, Augsburger JJ: Effects of argon laser photocoagulation on rubeosis iridis and angle neovascularization. Ophthalmology 87:400-402, 1980 43. Pavan PR, Folk JC, Weingeist TA et al: Diabetic rubeosis and panretinal photocoagulation: a prospective, controlled, masked trial using fluorescein angiography. Arch Ophthalmol 101:882-884, 1983 44. Ticho U, Patz A: The role of capillary perfusion in the management of diabetic macular edema. Am J Ophthalmol 76:880-886, 1973 45. Bresnick GH, DeVenecia G, Myers FL et al: Retinal ischemia in diabetic retinopathy. Arch Ophthalmol 93:1300-1310, 1975 46. Tamura T, Tamura M: Perifoveal capillary network and visual prognosis in diabetic retinopathy. Ophthalmologica 185:141-146, 1982 47. Arend O, Wolf S, Harris A, Reim M: The relationship of macular microcirculation to visual acuity in diabetic patients. Arch Ophthalmol 113:610-614, 1995 48. Bresnick GH: Diabetic macular edema, a review. Ophthalmology 93:989-997, 1986 49. Aiello LM, Rand LI, Briones JC, Wafai MZ, Sebestyen JG: Diabetic retinopathy in Joslin Clinic patients with adultonset diabetes. Ophthalmology 88(7):619-623, 1981 50. Marcus DF, Aaberg TM: Argon laser photocoagulation treatment in diabetic cystoid maculopathy. Ann Ophthalmol 9:365-372, 1977 51. Blankenship GW: Diabetic macular edema and argon laser photocoagulation, a prospective randomized study. Ophthalmology 86:69-75, 1979 52. British Multicentre Study Group: Photocoagulation for diabetic maculopathy: a randomized controlled clinical trial using xenon arc. Diabetes 32:1010-1016, 1983 53. Olk RJ: Modified grid argon (blue-green) laser photocoagulation for diffuse macular edema. Ophthalmology 93:938950, 1986 54. Early Treatment Diabetic Retinopathy Study Research Group: Photocoagulation for diabetic macular edema: ETDRS report number 4. Int Ophthalmol Clin 27:265-272, 1987 55. Karacorlu S, Burumcek E, Karacorlu M, Arslan O: Treatment of diabetic macular edema: a comparison between argon and dye lasers. Ann Ophthalmol 25:138-141, 1993 56. Gupta V, Gupta A, Kaur R, Narang S, Dogra MR: Efficacy of various laser wavelengths in the treatment of clinically significant macular edema in diabetics. Ophthalmic Surg Lasers 32(5):397-405, 2001 57. Early Treatment Diabetic Retinopathy Study Research Group: Focal photocoagulation treatment of diabetic macular edema: relationship of treatment effect to fluorescein angiographic and other retinal characteristics at baseline: ETDRS report number 19. Arch Ophthalmol 113:1144-1155, 1995 58. Little HL: Complications of argon laser retinal photocoagulation: a five year study. Int Ophthalmol Clin 16:145-159, 1976 59. Menchini U, Scialdone A, Pietroni C, Carones F, Brancato R: Argon versus krypton panretinal photocoagulation side
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effects on the anterior segment. Ophthalmologica 201:6670, 1990 Bloom SM, Mahl CF, Schiller SB: Lenticular burns following argon panretinal photocoagulation. Br J Ophthalmol 76:630-631, 1992 Lakhanpal V, Schockett SS, Richards RD, Niraukari VS: Photocoagulation-induced lens opacity. Arch Ophthalmol 100:1068-1070, 1982 Lerner BC, Lakhanpal V, Schocket SS: Transient myopia and accommodative paresis following retinal cryotherapy and panretinal photocoagulation. Am J Ophthalmol 97(6): 704-708, 1984 Lee BL, Van Heuven WA: Hypopyon uveitis following panretinal photocoagulation. Ophthalmic Surg Lasers 28(6): 505-507, 1997 Gentile RC, Stegman Z, Liebmann JM, Dayan AR, Tello C, Walsh JB, Ritch R: Risk factors for ciliochoroidal effusion after panretinal photocoagulation. Ophthalmology 103:827-832, 1996 Doft BH, Blankenship GW: Single versus multiple treatment sessions of argon laser panretinal photocoagulation for proliferative diabetic retinopathy. Ophthalmology 89: 772-779, 1982
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66. Yuki T, Kimura Y, Nanbu S et al: Ciliary body and choroidal detachment after laser photocoagulation for diabetic retinopathy: a high-frequency ultrasound study. Ophthalmology 104:1259-1264, 1997 67. Hessemer V, Schmidt KG: Influence of panretinal photocoagulation on the ocular pulse curve. Am J Ophthalmol 123:748-752, 1997 68. Moriarty AP, Spalton DJ, Shilling JS, Ffytche TJ, Bulsara M: Breakdown of the blood-aqueous barrier after argon laser panretinal photocoagulation for proliferative diabetic retinopathy. Ophthalmology 103:833-838, 1996 69. Pomerantzeff O, Kaneko H, Donovan RH et al: Effect of the ocular media on the main wavelengths of argon laser emission. Invest Ophthalmol 15:70-77, 1976 70. Schatz H, Madeira D, McDonald H et al: Progressive enlargement of laser scars following grid laser photocoagulation for diffuse diabetic macular edema. Arch Ophthalmol 109:1549-1551, 1991 71. Christoffersen N, Sander B, Larsen M: Precipitation of hard exudate after resorption of intraretinal edema after treatment of retinal branch vein occlusion. Am J Ophthalmol 126:454-456, 1998
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Retinal photocoagulation with diode lasers
Rosario Brancato1, Pier Giorgio Gobbi2 and Rosangela Lattanzio1 1 Department of Ophthalmology and Visual Sciences; 2Laser Medicine Research; University Hospital San Raffaele, Milan, Italy
Keywords: diode laser, physics of the diode laser, laser tissue interaction, diabetic retinopathy, retinopathy of prematurity, cyclophotocoagulation
Introduction A little physics and technology Semiconductor lasers (also called diode or junction lasers) belong to the family of solid-state lasers, where the active medium responsible for laser emission is a piece of solid material. Unlike their companions (to name a few that have relevance to ophthalmology: the historical ruby laser, the ubiquitous Nd:YAG, and the more exotic Er:YAG), energy excitation (or ‘pumping’) in diode lasers is not optical, e.g., by means of a lamp, but electrical through a current injected through the sample. Another remarkable difference is that the laser medium is not an insulating crystal (like ruby), but rather a material where electrons are only loosely bonded to the lattice atoms, so that current flow is possible (‘semiconductor’ because the electrical properties are intermediate between those of insulators and conductors, such as metals). Pure semiconductors, such as silicon and germanium, are of little or no utility in electronics. For effective operation, a junction has to be realized, that is a boundary region between two sides of a semiconductor slab, each of which is doped with different impurity atoms (for silicon, for example, arsenic and indium can be used). This doped slab realizes a diode, i.e., a device in which, after application of voltage across its ends, current flow is only allowed in one direction and is inhibited in the opposite one. In order for the junction to emit light, silicon and germanium do not work, and more complex semiconductor materials are required, in the form of binary, ternary, or quaternary compounds (with two, three, or four constituents); dopant elements vary accordingly. The simplest structure is the light
emitting diode (LED), which radiates light incoherently in all directions, just like a tiny bulb. LEDs are used in electronics as indicators and displays, and have become so bright that they are used instead of filament lamps in traffic lights and stop lights. In order to realize a laser device, which operates on stimulated rather than spontaneous emission, semiconductor physics imposes severe material selection, because only a few LED materials match the necessary requirements for laser action. Furthermore, an optical resonator comprising two mirrors must be realized in order to provide regenerative amplification and to generate a spatially collimated beam. Finally, the efficiency and lifetime of the device must be maximized in order to allow, not only pulsed, but also continuous wave operation at room temperature, to reduce threshold and operating currents, and to improve the optical quality of the laser beam. All such requirements result in a very complex structure, comprising many semiconductor layers (at least three), of different composition, doping, and thickness, a typical example of which is sketched in Figure 1. The structure is implemented on a single monolithic ‘chip’, with typical dimensions of 0.5 mm length, 0.2 mm width, and some 1 µm thickness. Such a device, capable of emitting several watts of optical power, is placed in contact with massive copper heatsinks, and packaged in metal cases which look like power transistors (see Fig. 2). To increase the power output level, many parallel stripes can be grown in a single chip, leading to monolithic laser diode arrays, however, with degradation of the optical quality of the beam. Higher powers are commonly available if the laser is operated in a repetitively pulsed mode. The laser emission wavelength strictly depends
Address for correspondence: Professor Rosario Brancato, Department of Ophthalmology and Visual Sciences, University Hospital San Raffaele, Via Olgettina 60, 20132 Milano, Italy. e-mail:
[email protected] Lasers in Ophthalmology – Basic, Diagnostic and Surgical Aspects, pp. 241–254 edited by F. Fankhauser and S. Kwasniewska © 2003 Kugler Publications, The Hague, The Netherlands
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Fig. 1. Schematic structure of a high power diode laser. The beam emitted from the junction is in the form of a thin slit, and evolves into an elliptical spot, with unequal divergence upon the two axes.
Fig. 2. View of a 1-W diode laser after removal of the package cover and window. The arrow indicates the location of the semiconductor junction (200 × 0.5 µm). The emission is perpendicular to the package plane.
upon the semiconductor composition: with ternary and quaternary materials, the resulting wavelength varies if the relative concentration of the constituting elements is changed. In general, laser diode wavelengths are concentrated in the infrared portion of the spectrum (up to 2000 nm), and in the visible, only the red part is covered. It is physically and technologically difficult to obtain emission from semiconductor lasers below 600 nm, except at very low powers and possibly at cryogenic temperatures. The laser diodes in use in ophthalmology mostly use the ernary semiconductor GaAlAs, emitting at 810 nm and capable of a few watts of power from a single stripe; the red wavelengths at 670690 nm for Photo-Dynamic Therapy (PDT) treatments are obtained from the quaternary material AlGaInP.
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Diode laser beam The characteristics of the beam coming out of semiconductor lasers markedly differ from those of other types of lasers, particularly of gas lasers (such as argon and helium-neon). In fact, these emit a round, well-shaped beam which stays collimated for long distances: the typical divergence of 1 milliradian implies that the beam size grows by 1 mm after 1 m propagation. Diode lasers, on the other hand, generate noncircular beams which spread out very rapidly. This is mostly due to the fact that the emitting area (the section of the active layer) is slit-shaped, a fraction of a micrometer high and several tens or hundreds of micrometers long. Concurrently, the beam divergence is extremely high, some 10 × 40° typically (i.e., 0.17 × 0.7 radians)
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Retinal photocoagulation with diode lasers parallel and perpendicular to the active layer, respectively (see Fig. 1). External optics are necessary to compensate for this high divergence (collimation), and cylindrical optics to transform the beam shape from oval to round (equalization). Intrinsic to the structure of the semiconductor junction are also two other specific features of diode laser emission, namely the large spectral linewidth (some 100-1000 times greater than in ion gas lasers), and the dependence of the peak wavelength on the operating temperature of the junction (λ tends to increase with temperature). However, there is no doubt that the most distinctive characteristic of diode lasers is their small size combined with their high level of output power. The comparison with argon or krypton ion lasers, capable of emitting the same output power, is dramatic: a few cubic centimeters of solid state material on one side compared to liters and liters of sealed laser tube on the other. The ability of being able to generate watt-level power from such a small volume necessarily implies a relevant efficiency, in terms of output optical power divided by the input electrical power required. Indeed, diode lasers exhibit efficiencies of up to 30-40%, a factor some 1000 times larger than that for an ion laser tube. Since all the input power not converted into useful output must be dissipated as heat in some way, it is clear that the requirements of a semiconductor laser are far less stringent: there is no need for liquid cooling, heat exchangers, or fans. The consequence is that the power supply must provide an electrical power at most three times larger than the optical power output, and thus it can be dimensioned very compactly, with the advantage of working at low voltages (2-3 Volts) and moderate currents (a few Amperes). Another notable feature of diode lasers is that they are monolithic devices, and as such do not require internal adjustments or maintenance. They have shown reliability performances typical of electronic components, with a lifetime exceeding 20,000 hours of continuous operation. A short history The semiconductor laser was conceived by the Nobel prize winner Basov in 1961,1 and operated for the first time in 1962,2 shortly after the invention of the ruby laser. Within a short time, laser action was being demonstrated in many different semiconductor materials, but at the beginning, they could only be operated in pulsed mode, at cryogenic temperatures, and for a short time. It was only in 1970 that a diode laser was able to produce a continuous-wave beam at room temperature. The state-of-the-art in semiconductor lasers continued to advance at an amazing rate. In the mid 1980s, specific technological advancements (molecular beam epitaxy and quantum-well structures) boosted the output power capabilities well beyond the watt
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243 limit, and operation could be maintained for thousands of hours without the risk of self-destruction. Such achievements made the semiconductor laser suitable for applications in the biomedical field, mainly as a replacement for conventional laser sources. In 1984, Pratesi was the first to realize the enormous potential offered by diode lasers, and suggested the possibility of using them in several fields of photomedicine.3 Shortly after, in 1986, the first retinal photocoagulation experiments with diode lasers were performed on animal eyes.4,5 In 1990, the first commercial photocoagulators based on semiconductor laser sources appeared on the market. At the beginning, also because of power output restrictions, the laser beam was directly coupled from the diode source to the spot-forming optics of the photocoagulator: this sometimes resulted in hot spots being present within the irradiated area, as a consequence of a non-uniform emission across the active layer. Later, when increased power became available, the beam was injected into, and transmitted through, an optic fiber in the same way this usually occurred with ion photocoagulators, thus achieving a smooth and uniform power density at the focal spot. The difference with ion lasers is that a larger core fiber must be used (typically, 200 versus 50 µm), as a result of the multimode nature of the beam emitted from semiconductor lasers, which therefore cannot be as tightly focused as single mode laser sources. This is commonly reflected in the smallest spot size selectable, which usually is 100 µm instead of 50 µm with argon and krypton laser photocoagulators. At the present time, several companies produce diode laser photocoagulators commercially that emit at 810 nm. Possibly the greatest merit of these instruments is to be found in their ergonomic advantages (compactness, portability, reliability) and economic savings compared to bulky and expensive ion systems. This made a wide diffusion of laser photocoagulator systems available to general ophthalmologists, not only retina specialists, thus increasing the therapy chances for a number of patients. Laser-tissue interaction The adoption of a near-infrared wavelength for retinal photocoagulation in place of visible radiation brings with it slight modifications to the picture of the physical interaction of light with tissue, and consequently to the clinical protocols to be followed for effective therapeutic effects. In general, the transmission through the refractive ocular media is rather flat all across the visible spectrum, with a broad relative peak at around 850 nm (Fig. 3). The transmittances at diode and krypton laser lines are almost identical (about 97%), and are definitely higher than that for the argon green line (85%). Such a 12% difference is partly
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Fig. 3. Spectral transmission of optical radiation in the human eye (from cornea to retina) and spectral absorption in the retina and choroid. The vertical lines pinpoint the laser wavelengths used more often in retinal photocoagulation: argon blue (488 nm); argon green (514.5 nm); krypton red (647 nm); and semiconductor diode (810 nm).
due to the higher reflectance of the ocular interfaces at short wavelengths, but mostly to the scattering losses suffered by radiation in propagating through the refractive media. Scattering (or diffusion) has a much stronger dependence on wavelength: the scattering effect is six times greater for the argon green wavelength relative to the diode laser, and eight times greater for the argon blue line (only two and a half times for the krypton red line). This markedly different behavior is not usually appreciated in normal subjects, but it can make a difference in mild cataractous eyes, in which the diode laser infrared radiation is definitely superior in concentrating more power at the retinal spot, being less dispersed when passing through the lens. In this way, the diode laser is able to perform retinal photocoagulation in conditions in which an argon laser would be unsuccessful. A similar advantage can be observed in the presence of intravitreal or intraretinal hemorrhages, because the attenuation coefficient of hemoglobin is 25 times higher for argon green than for semiconductor lasers. Thus, the diode laser can be used more safely for performing photocoagulation through mild hemorrhages, in the same way that krypton lasers are preferred to argon ones for this task. The specific therapeutic action of laser radiation is achieved through absorption in the retinal layers and the consequent heating of the surrounding tissue to the coagulation temperature (typically 6070°C). Here, the chromophore responsible for the absorption of light is essentially melanin, whose
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granules are mainly concentrated in the retinal pigment epithelium. The spectral absorbance behavior of melanin is regular, constantly declining from the ultraviolet to the infrared. As a result, at the diode laser wavelength of 810 nm, the overall absorption in the retinal pigment epithelium (Fig. 3) is about one half the value at the argon green line, and 60% that at the krypton red line. This fact has some important consequences. Being less absorbed in the retina, the diode laser wavelength requires a higher dose (energy = power × exposure time) to obtain the same temperature rise in the tissue compared to the argon green wavelength. The threshold dose for photocoagulation was shown to be three to four times higher for the diode laser, as a combined result of reduced absorption and greater absorbing volume. In fact, the penetration depth is definitely higher with the diode laser because the absorption coefficient (absorption per unit length) is markedly lower than at the green wavelength. Since coagulation takes place deeper in the tissue, involving the choroid as well as the retinal pigment epithelium, it is somewhat more difficult to detect it: while, with the argon laser, whitening of the retina (due to protein denaturation) clearly indicates when the irradiation should be stopped, with the diode laser, the effect of coagulation is seen rather more as a slight graying of the area being treated. More caution is therefore necessary, in order to avoid the risk of overexposure. In principle, more power, released deeper in the tissue, should result in more pain during treatment, and indeed discomfort to patients was sometimes
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Retinal photocoagulation with diode lasers reported, particularly in the first years of diode laser photocoagulation. However, it was also shown that, with optimized treatment parameters, no significant difference in pain was experienced by patients undergoing diode compared to argon laser treatment, but rather some preference emerged for the infrared wavelength because of the absence of bright flashes during treatment.6
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Among different medical applications, ophthalmology requires the lowest laser intensity for therapeutic effects and was therefore the first to benefit from the availability of diode lasers. Until then, photocoagulation of the retina had only been carried out using the (blue-)green or red lines of ion lasers; dye lasers were very useful, thanks to their tunability over the entire visible and near-infrared spectrum. Transmission of these wavelengths through the eye was high, and absorption by melanin and hemoglobin great enough to permit coagulation with short exposure times. As outlined above, the infrared radiation emitted by diode lasers appeared to be appealing for chorioretinal photocoagulation, by using either transpupillary irradiation with standard slit-lamps, or contact transscleral irradiation with suitably shaped fiber tips. The first coagulations of the chorioretina with diode lasers (cw/pulsed) were reported in 1986 in connection with endophotocoagulation4 and transpupillary photocoagulation5 of rabbit eyes. Brancato et al.5 used a GaAlAs diode laser emitting at 901 nm; the average input laser power was
30 mW, and the exposure time varied between four and ten seconds. The lesions obtained with a fivesecond exposure were minimal, whitish, and pointlike, while with an irradiation of ten seconds, the chorioretinal photocoagulation appeared white, surrounded by a slightly darker grayish area about 50 µm wide. Histological examination, carried out three days later, showed an alteration mainly located at the pigment epithelium with vacuolization and karyolysis of the outer nuclear layer which was more evident in longer exposure photocoagulations. Diode laser retinal photocoagulations were similar, both ophthalmoscopically and in histological lesions, to the photocoagulations obtained with standard visible lasers at shorter wavelengths.5,7 Similar histological data were reported by Puliafito et al. in 1987 when performing retinal endophotocoagulations using a semiconductor diode prototype laser at 808 nm.8 In 1988, Brancato et al.9 presented their results obtained on rabbit eyes with a transpupillary slit-lamp photocoagulator coupled to a semiconductor laser emitting at 811 nm in the continuous wave (Figs 4-7). In 1989, the histopathology of diode lesions in rabbit retina was reported in a comparative study with the argon laser.10 Lesions ophthalmoscopically similar to those obtained therapeutically in humans were obtained with a diode and an argon laser. Twenty-four hours after treatment, these lesions were studied by light and electron microscopy: argon irradiations resulted in damage to all the retinal layers, but especially the retinal pigment epithelium, while diode laser radiation produced damage to the outer retina and choriocapillaris; most internal retinal structures and the internal limiting membrane were spared. A sufficient fraction of diode laser radiation could propagate into the choroid in-
a.
b.
The pioneering days of diode laser photocoagulation
Fig. 4. Rabbit fundus photographs: a. one hour, and b. 15 days after diode laser photocoagulation.
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Fig. 5. Rabbit retinal fluorescein angiography: a. one hour, and b. 15 days after diode laser photocoagulation.
Figs. 6 and 7. Rabbit optical microscopy after diode laser irradiation with intense coagulation necrosis of the retinal pigment epithelium and photoreceptors cells.
ducing irregular obliteration of the choriocapillaris, a certain degree of disorganization of the choroidal pigment, and/or edema. The minimal structural cellular reaction at the vitreoretinal interface could prevent the postphotocoagulative appearance of proliferative vitreoretinopathy. In the same study,10 Brancato et al. showed that opthalmoscopically similar retinal lesions could be produced with argon and diode lasers at comparable irradiation levels (about 120 W/cm²), although the argon laser exposure time was slightly shorter. This result was rather unexpected, since it was well known that melanin absorption decreases as the radiation wavelength approaches the infrared spectral region. On the other hand, it was reported that, in addition to a radiation range between 800 and 900 nm, low energy levels were required to produce threshold retinal lesions.8,11 Furthermore, diode laser radiation is well transmitted by the ocular optic media, thus reducing the total amount of energy needed to produce a photocoagulative effect. Other histopathological studies were undertaken
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to assess diode laser-induced retinal thermal damage, mainly in rabbit eyes; all the results demonstrated the clinically evident dose-response effect, namely sparing of inner retinal elements with mild burns and full-thickness retinal cell loss with severe burns. Longer irradiation exposure appeared to be a safer way of producing a severe burn than higher power; burns characteristically bloomed in the few seconds following laser application, indicating the deep localization of energy absorption.12-17 Other groups employed fluorescein angiography to control laser-induced thermal damage, using the quantification of fluorescence staining in terms of both intensity and time.18,19 The diode laser was also tested to evaluate its efficacy in obtaining retinochoroidal adhesions.20,21 All these experimental results had the effect of opening a fascinating new era for laser applications in ophthalmology, offering almost ideal laser systems for the treatment of a variety of eye pathologies.
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Retinal photocoagulation with diode lasers Initial experiments to investigate the clinical use of diode lasers for the treatment of chorioretinal diseases were carried out in 1989,22,23 and these opened the way for several extended studies, reported below. Transpupillary photocoagulation Retinal vascular diseases Preliminary results on the efficacy of diode laser photocoagulation in the treatment of retinal vascular diseases were reported in 1989-1992,22,24-31 and confirmed its potential technical and clinical advantages. In 1993, Bandello et al.32 published data on a randomized study of diode versus argon-green panretinal laser photocoagulation in proliferative diabetic retinopathy, with a mean follow-up of two years: the long-term efficacy of the diode laser turned out to be similar to that of the argon laser (100% of new vessel regression versus 91%). The complications resulting from the overdosage of diode energy observed in preliminary reports (Figs. 8, 9) were partially reduced thanks to the lower energy used, but were still high (choroidal detachment, neurotrophic keratopathy, troublesome pain, severe field loss). When performing panretinal photocoagulation at 810 nm radiation, three problems were encountered: (1) it was difficult to judge the intensity of the burn by its appearance; (2) only a narrow power range was shown to produce satisfactory retinal thermal damage; and (3) pigmentation of the treated areas strongly influenced the thermal response. These effects made it possible to produce an unintentional variation in retinal exposure (Fig. 10). These features, together with the complications observed, forced investigators to perform diode
247 laser treatment using lower energy levels than in preliminary reports, in order to obtain non-contiguous gray laser spots only.32 Moreover, the chronic fluorangiographic aspects observed after diode laser irradiation (i.e., marked chorioretinal atrophy larger than acute lesions and greater size and intensity of the chorioretinal atrophy than those produced by the argon laser, see Figs 11-13) deserved some consideration. Initially, these findings were attributed to a higher energy deposit in the choroid because of overdose treatments. However, similar aspects of chorioretinal atrophy were found in eyes treated with lower diode energy; this phenomenon seemed to be wavelengthdependent rather than energy-related: approaching longer wavelengths, absorption by the retinal pigment epithelium decreases and the thermal effect moves to the choroid, as previously demonstrated histologically.10,33 Therefore, the pronounced damage to the choriocapillaris produced by the diode laser in the acute phase contributes to the suffering of the overlying retinal layers, resulting, in the chronic stage, in complete retinal atrophy.10,12 Moreover, Ulbig and Hamilton34 underlined the fluorangiographically-deeper effects of the diode laser in the choroid; compared to the argon laser, their diabetic patients found diode laser treatment more painful, but appreciated the absence of bright flashes during therapy. Altering the pulse configuration resulted in affecting the pain response during diode laser photocoagulation.35 Again, in patients with proliferative diabetic retinopathy, a tendency towards lower decline in color contrast sensitivity and pattern electroretinogram recordings was reported after diode laser compared to argon laser photocoagulation.36,37 The diode laser was also shown to be a viable tool for managing macular edema secondary to diabetes or retinal vein occlusion.22,28,37-42
Fig. 8. Normo- (A) and overdosed (B) diode laser burns in the human retina.
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Fig. 9. Choroidal detachment after an overdose of diode laser photocoagulation.
Retinopathy of prematurity Diode laser photocoagulation for retinopathy of prematurity was applied for the first time in 1992, in nine infants by means of an indirect ophthalmoscopic delivery system;43 the anatomical and functional results were satisfactory and encouraged the following trials.44-52 Some side-effects were re-
ported, such as lens changes, phthisis bulbi, hyphema, and angle-closure glaucoma.53-56 Other authors proposed transscleral diode laser retinal photocoagulation for threshold retinopathy of prematurity, and reported a favorable outcome.57-62 The technique proved to be as effective as transpupillary diode laser photocoagulation, but minor side-effects were noted. However, the technique turned out to be a technically straightforward alternative to cryotherapy. In 1997, Young et al.63 reported histopathology and vascular endothelial growth factor (VEGF) expression in untreated and diode laser-treated eyes of an infant with stage 3 retinopathy of prematurity. In the treated eye, the histopathological results demonstrated the clinically evident dose-response effect, i.e., sparing of inner retinal elements with mild laser burns, and full-thickness retinal cell disruption with severe burns; scleral and ciliary nerve effects were absent. VEGF mRNA was found to be elevated in the peripheral avascular retina of the untreated eye, consistent with the hypothesis that retinal hypoxia stimulates VEGF expression. In the treated eye, VEGF mRNA was not detected in the photocoagulated areas, but increased between laser scars.
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b. Fig. 10. Appearance of diode laser spots: a. immediately, and b. three months after retinal transpupillary photocoagulation.
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Retinal photocoagulation with diode lasers The histopathological study reported by Park et al.64 in diode laser photocoagulated eyes disclosed segmental areas of chorioretinal scarring with retinal atrophy and gliosis, loss of retinal pigment epithelium, and extensive atrophy of the choroid and its vasculature, resembling lesions described after transscleral cryotherapy, but with less severe chorioretinal damage. Although advances in scleral buckling and vitrectomy techniques offer hope for infants suffering from stage 4 or 5 retinopathy of prematurity,
249 prevention of progression to these stages offers the greatest promise for favorable structural and visual outcomes. Technological advances in screening tools and portable diode lasers enable ophthalmologists to successfully manage threshold retinopathy of prematurity. Macular degeneration Diode laser photocoagulation was applied in patients with macular degeneration, and it turned out
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b. Fig. 11. Fluorescein angiography: a. before, and b. after panretinal photocoagulation for proliferative diabetic retinopathy; when the treatment was carried out with an overdose of energy, confluent areas of chorioretinal atrophy could be observed at follow-up.
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b. Fig. 12. Fluorescein angiography: a. before, and b. after panretinal photocoagulation for proliferative diabetic retinopathy.
to be at least as effective as that using conventional lasers; advantages were found when performing indocyanine green dye-enhanced diode laser photocoagulation of new choroidal vessels or feeder vessels.40,65-70 Olk et al.71 reported therapeutic benefits from diode laser macular grid photocoagulation in the prophylactic treatment of non-exudative macular degeneration, significantly reducing drusen levels and improving visual acuity when either visible endpoint burns or subthreshold endpoint lesions were used. Complications were fewer using sub-
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threshold endpoint lesions. Data from this clinical pilot study have been used to design the Prophylactic Treatment of AMD Trial (PTAMD), a multicenter, randomized, prospective, clinical trial currently in progress, which compares subthreshold (invisible) treatment with observation in eyes with nonexudative AMD. Tumors In 1989, a patient with choroidal melanoma, whose eye was about to be enucleated, was briefed about
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b. Fig. 13. Fluorescein angiography: a. three months, and b. one year after panretinal photocoagulation for proliferative diabetic retinopathy; angiograms show the enlargement of scars during the one-year follow-up.
the aim of human experimental diode laser retinal photocoagulation, which first demonstrated the possibility of obtaining therapeutically useful chorioretinal photocoagulation, ophthalmoscopically and histologically similar to that produced by ion lasers.23 The diode laser has proved to be effective in the treatment of choroidal hemangiomas,72 and has been tested in capillary papillary hemangiomas.73
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Transscleral diode laser retinal photocoagulation The transscleral application of semiconductor diode lasers was first reported in 1990;74,75 Peyman et al.75 achieved retinal photocoagulation in rabbit eyes using energy levels of 200-500 mW for 0.5 seconds; chorioretinal scar formation was observed clinically and histologically within two to three weeks. Mouries et al.76 examined laser burns obtained
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with transscleral diode laser photocoagulation in rabbits by light and electron microscopy and found them to be similar to those produced by argon and krypton lasers. Histopathological evaluation of the lesions demonstrated an intact sclera overlying the chorioretinal lesions. The results of experimental studies support the hypothesis that transscleral retinal photocoagulation using the diode laser in selected indications may be a valuable alternative to cryotreatment and diathermy in the human eye. The absence of scleral damage and pigmented epithelium cell dispersion, as well as the decreased breakdown of the blood ocular barrier after transscleral diode laser photocoagulation, were the main advantages of the technique. Various studies showed that transscleral diode laser photocoagulation of retinal breaks is effective and safe;77-80 in particular, Bonnet and Mouries77 demonstrated that it is a valuable alternative to cryotreatment in eyes at high risk of postoperative proliferative vitreo-retinopathy (PVR), and/or when argon laser photocoagulation cannot be used, and/ or in retinal detachments after previous failed surgery. Effective diode laser retinopexy has been reported in retinal detachment surgery.80-82 Transscleral diode laser retinal photocoagulation turned out to be an effective and safe treatment in proliferative sickle cell retinopathy with vitreous bleeding,83,84 and in rubeosis iridis/neovascular glaucoma.85,86 Endo-ocular diode laser retinal photocoagulation The first coagulations of the chorioretina with a diode endolaser were reported in 1986 in rabbits.5 Some years later, diode endolaser photocoagulation was applied in 25 patients with proliferative diabetic retinopathy, proliferative vitreoretinopathy, complex retinal detachments, or retinal breaks.87 Good retinal and retinal pigment epithelium laser uptake was observed in all cases; the clinical appearance of the burns was similar to that with the argon laser, but it was subtly lighter, especially in less-pigmented areas and in eyes. The logistical advantages offered by this system have been confirmed by other studies,38,88 and it is now routinely applied in vitreoretinal surgery. Conclusions The development of diode laser photocoagulators represented a major technological breakthrough relative to conventional ion laser devices, and opened a new era in the field of retinal photocoagulation. The technical and ergonomic advantages of diode laser sources are well evident and certain: compactness, portability, long lifetime, reduced main-
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tenance costs. Their attractiveness for a variety of microsurgical procedures has led to a widespread diffusion of the diode laser in the field of vitreoretinal surgery. In transpupillar retinal photocoagulation, the diode laser fully demonstrated its efficacy and a substantial overlap of its therapeutic results with argon and krypton laser sources, despite a few differences experienced in the course of the treatment. In the last years, the diode laser has been progressively surpassed in the current practice by the diodepumped frequency-doubled Nd:YAG laser photocoagulators, which combine the compactness and reliability of a solid-state laser source with the advantage of a visible wavelength (532 nm), closely resembling the emission line of argon green photocoagulators (514.5 nm). References 1. Basov NG, Krokhin ON, Popov YM: Production of negative temperature states in P-N junctions of degenerate semiconductors. Journal Experimental Theoretical Physics 40: 1320, 1961 2. Hall RN, Fenner GE, Kingsley JD, Soltys TJ, Carlson RO: Coherent light emission from GaAs junctions. Phys Rev Lett 9:366, 1962 3. Pratesi R: Diode lasers in photomedicine. J Quantum Electronics 20:1433-1439, 1984 4. Deutsch TF, Boll J, Puliafito CA, To K: Semiconductor laser photocoagulation of the retina. In: Technical Digest, Conference on Lasers and Electro-Optics, pp 150-151. San Francisco, CA: OSA/IEEE 1986 5. Brancato R, Giovannoni L, Pratesi R, Vanni U: New lasers for ophthalmology: retinal photocoagulation with pulsed and diode lasers. In: Proceedings of the 1986 Conference on Optics, Optical Systems and Applications (ECOOSA), Florence. SPIE 701:365-366, 1986 6. Goebel W, Pfeiffer N, Grehn F: Patient discomfort during laser treatment, a comparison between diode and argon laser. Invest Ophthalmol Vis Sci 35(Suppl):553, 1994 7. Brancato R, Pratesi R: Applications of diode lasers in Ophthalmology. Lasers Ophthalmol 1:119-129, 1987 8. Puliafito CA, Deutsch TF, Boll J, To K: Semiconductor lasers endophotocoagulation of the retina. Arch Ophthalmol 105:424-428, 1987 9. Brancato R, Pratesi R, Leoni G, Trabucchi G, Giovannoni L, Vanni U: Retinal photocoagulation with diode lasers operating from a slit lamp microscope. Lasers Light Ophthalmol 2:73, 1988 10. Brancato R, Pratesi R, Leoni G, Trabucchi G, Vanni U: Histopathology of diode and argon laser lesions in rabbit retina: a comparative study. Invest Ophthalmol Vis Sci 30: 1504-1510, 1989 11. Lund DJ, Beatrice ES, Schuschereba ST: Bioeffects concerning the safe use of GaAs laser training devices. In: Beatrice ES (ed) Combat Ocular Problems, pp 15-29. San Francisco, CA: Letterman Army Institute of Research 1985 12. Smiddy WE, Hernandez E: Histopathologic results of retinal diode laser photocoagulation in rabbit eyes. Arch Ophthalmol 110:693-698, 1992 13. Benner JD, Huang M, Morse LS, Hjelmeland LM, Landers MB III: Comparison of photocoagulation with the argon, krypton, and diode laser indirect ophthalmoscopes in rabbit eyes. Ophthalmology 99:1554-1563, 1992
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Retinal photocoagulation with diode lasers 14. Cho HK, Park YW, Kim YJ, Shyn KH: Histopathologic and ultrastructural findings of photocoagulation lesions produced by transpupillary diode laser in the rabbit retina. J Korean Med Sci 8:420-430, 1993 15. McHugh D, England C, Van der Zypen E, Marshall J, Fankhauser F, Fankhauser-Kwasniewska: Irradiation of rabbit retina with diode and Nd:YAG lasers. Br J Ophthalmol 79:672-677, 1995 16. Richardson PR, Boulton ME, Duvall-Young J, McLeod D: Immunocytochemical study of retinal diode laser photocoagulation in the rat. Br J Ophthalmol 80:1092-1098, 1996 17. Pollack JS, Kim JE, Pulido JS, Burke JM: Tissue effects of subclinical diode laser treatment of the retina. Arch Ophthalmol 116:1633-1639, 1998 18. Desmettre T, Devoisselle JM, Soulie-Begu S, Mordon S: Value of fluorescein angiography in control of retinal thermal damage due to diode laser. J Fr Ophtalmol 22:730-737, 1999 19. Desmettre TJ, Soulie-Begu S, Devoisselle JM, Mordon SR: Diode laser-induced thermal damage evaluation on the retina with a liposome dye system. Lasers Surg Med 24:61-68, 1999 20. Aso S: Retinochoroidal adhesions of diode laser endophotocoagulation lesions. 1. Histopathology of rabbit eyes. Nippon Ganka Gakkai Zasshi 96:479-485, 1992 21. Smiddy WE, Hernandez E: Histopathologic characteristics of diode laser-induced chorioretinal adhesions for experimental retinal detachment in rabbit eyes. Arch Ophthalmol 110:1630-1633, 1992 22. McHugh JD, Marshall J, Ffytche TJ, Hamilton AM, Raven A, Keeler CR: Initial clinical experience using a diode laser in the treatment of retinal vascular disease. Eye 3:516527, 1989 23. Brancato R, Pratesi R, Leoni G, Trabucchi G, Vanni U: Semiconductor diode laser photocoagulation of human malignant melanoma. Am J Ophthalmol 107:295-296, 1989 24. Noyori K, Noyori S, Ryutaro O: Clinical trial of diode laser photocoagulation: a preliminary report. Laser Light Ophthalmol 2:81-87, 1990 25. Brancato R, Bandello F, Trabucchi G, Leoni G, Lattanzio R: Argon and diode laser photocoagulation in proliferative diabetic retinopathy: a preliminary report. Lasers Light Ophthalmol 3:233-237, 1990 26. Puliafito CA: Semiconductor diode laser photocoagulation in retinal vascular diseases. Laser Surg Med 2:S67, 1990 27. Balles MW, Puliafito CA: Semiconductor diode lasers: a new laser light source in ophthalmology. Int Ophthalmol Clin 30:77-83, 1990 28. Balles MW, Puliafito CA, D’Amico DJ, Jacobson JJ, Birngruber R: Semiconductor diode laser photocoagulation in retinal vascular disease. Ophthalmology 97:15531561, 1990 29. Noyori S, Iijima M, Ohki R, Noyori K, Yoneya S: Effects on the retina and choroid of transpupillary diode laser photocoagulation. Nippon Ganka Gakkai Zasshi 95:758766, 1991 30. Brancato R, Bandello F: New laser modalities for posterior segment treatment. Curr Opin Ophthalmol 2:299-305, 1991 31. Buckley S, Jenkins L, Benjamin L: Field loss after pan retinal photocoagulation with diode and argon lasers. Doc Ophthalmol 82:317-322, 1992 32. Bandello F, Brancato R, Trabucchi G, Lattanzio R, Malegori A: Diode versus argon-green laser panretinal photocoagulation in proliferative diabetic retinopathy: a randomized study in 44 eyes with a long follow-up time. Graefe’s Arch Clin Exp Ophthalmol 231:491-494, 1993 33. Puliafito CA, Deutsch TF, Boll J, To K: Semiconductor
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tion in prophylactic treatment of nonexudative age-related macular degeneration: two-year results of a randomized pilot study. Ophthalmology 106:2082-2090, 1999 Lanzetta P, Virgili G, Ferrari E, Menchini U: Diode laser photocoagulation of choroidal hemangioma. Int Ophthalmol 19:239-247, 1995 Garcia-Arumi J, Sararols LH, Cavero L, Escalada F, Corcostegui BF: Therapeutic options for capillary papillary hemangiomas. Ophthalmology 107:48-54, 2000 Jennings T, Fuller T, Vukich JA, Lam TT, Joondeph BC, Ticho B, Blair NP, Edward DP: Transscleral contact retinal photocoagulation with an 810-nm semiconductor diode laser. Ophthalmic Surg 12:492-496, 1990 Peyman GA, Naguib KS, Gaasterland D: Transscleral application of a semiconductor diode laser. Lasers Surg Med 10:569-575, 1990 Mouries O, Vitrey D, Germain P, Bonnet M: Use of diode laser in transscleral retinal photocoagulation. J Fr Ophtalmol 16:108-113, 1993 Bonnet M, Mouries O: Pilot study of transscleral light coagulation of retinal breaks using diode laser. J Fr Ophtalmol 17:739-745, 1994 McHugh DA, Schwartz S, Dowler JG, Ulbig M, Blach RK, Hamilton PA: Diode laser contact transscleral retinal photocoagulation: a clinical study. Br J Ophthalmol 79:10831087, 1995 Loewenstein A, Lamborne AN, Haller JA, de Juan E Jr: Effect of scleral indentation on diode laser transscleral retinal photocoagulation. Ophthalmic Surg Lasers 29:658662, 1998 Duquesne N, Fleury J, Bonnet M: Postoperative proliferative vitreoretinopathy in rhegmatogenous retinal detachment associated with stage B preoperative proliferative vitreoretinopathy: comparative results of transscleral retinopexy with diode laser or transpupillary retinopexy with argon laser. J Fr Ophtalmol 21:555-559, 1998 Haller JA, Blair N, De Juan E Jr, De Bustros S, Goldberg MF, Muldoon T, Packo K, Resnick K, Rosen R, Shapiro M, Smiddy W, Walsh J: Transscleral diode laser retinopexy in retinal detachment surgery: results of a multicenter trial. Retina 18:399-404, 1998 Steel DH, West J, Campbell WG: A randomized controlled study of the use of transscleral diode laser and cryotherapy in the management of rhegmatogenous retinal detachment. Retina 20:346-357, 2000 Seiberth V: Trans-scleral diode laser photocoagulation in proliferative sickle cell retinopathy. Ophthalmology 106: 1828-1829, 1999 Seiberth V: Transscleral and transpupillary laser coagulation in proliferative sickle-cell retinopathy. Ophthalmologe 98:199-202, 2001 Tsai JC, Bloom PA, Franks WA, Khaw PT: Combined transscleral diode laser cyclophotocoagulation and transscleral retinal photocoagulation for refractory neovascular glaucoma. Retina 16:164-166, 1996 Flaxel CJ, Larkin GB, Broadway DB, Allen PJ, Leaver PK: Peripheral transscleral retinal diode laser for rubeosis iridis. Retina 17:421-429, 1997 Smiddy WE: Diode endolaser photocoagulation. Arch Ophthalmol 110:1172-1174, 1992 Sasoh M, Smiddy WE: Diode laser endophotocoagulation. Retina 15:388-393, 1995
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Central serous chorioretinopathy Antonio P. Ciardella, Sheau J. Huang, Danielle L.L. Costa, Irene M. Donsoff and Lawrence A. Yannuzzi LuEsther T. Mertz Retinal Research Department, Manhattan Eye, Ear & Throat Hospital, New York, NY, USA
Keywords: central serous maculopathy, treatment, epidemiology, classification, pathogenic mechanism, complications, angiography, fluorescein angiography, indocyanine angiography
Introduction Some of the clinical features of the idiopathic recurrent detachments of the macula have been known since 1866 when Von Graefe originally described the disorder he named relapsing central retinitis.1 Almost 100 years later, Bennet applied the term central serous retinopathy.2 At the same time, via the use of fluorescein angioscopy, Maumenee noted that the detachment of the macula resulted from a leak at the level of the retinal pigment epithelium (RPE).3 In the following years, numerous articles were published expanding our knowledge of the etiology, natural history, clinical manifestations, and treatment of this peculiar disorder.4-123 Since the disease appears predominantly to involve the choroid and, secondarily, the retina, the disorder is best termed central serous chorioretinopathy (CSC).4 Pathogenesis The pathophysiology of CSC is still not completely understood. The primary site of pathology begins with a disturbance of the RPE. Often, discoloration or elevation of the RPE is noted clinically. In turn, this leads to an alteration of the normally impermeable state of the RPE, leading to serous leakage underneath and through the RPE to produce a detachment of the neurosensory retina. It is difficult to accept that a single isolated disturbance of a few RPE cells may overwhelm the physiological pump of the neighboring normal RPE. Much more plausible is the assumption that there is a more diffuse dysfunction of the choroid to RPE as the basis of the disease.
In a series of patients with acute and chronic CSC studied with indocyanine green angiography (ICGA), Guyer et al. noted diffuse hyperpermeability around active leakage sites that were not appreciated on fluorescein angiography (FA).5 Furthermore, they described additional focal regions of ICG hyperfluorescence in clinically unaffected fellow eyes. Similar findings have also been described by Spaide et al. in a large series of patients with CSC.6 Interestingly, diffuse choroidal hyperpermeability was especially evident in older patients with chronic disease. It was seen bilaterally, even in asymptomatic eyes without discernible clinical or fluorescein angiography abnormalities. The cause of the choroidal abnormality is still unknown, but there are both clinical and experimental models suggesting that this diffuse hyperpermeability could be initiated by an abnormality of the autoregulation of the choroidal blood flow. In 1927, Horniker first suggested the psychogenic-related hypothesis of CSC.10 Clinically, CSC is more frequent among subjects with type A personality.7-9 Type A behavior has been associated with the increased basal levels and release of cathecolamine and cortisol in response to stress.124 Although the role played by those substances in CSC is not well understood, cortisol excess may cause increased capillary fragility and hyperpermeability, which in turn may lead to choroidal circulation decompensation and leakage of fluid into the subretinal space.125 Furthermore, the anti-inflammatory properties of steroids may cause delayed healing of the RPE defect. By suppressing synthesis of extracellular matrix components and inhibiting fibroblastic activity, cortisol may also directly damage the RPE cells or their tight junctions, thus delaying any reparative process in the damaged RPE
Address for correspondence: Lawrence A. Yannuzzi, MD, LuEsther T. Mertz Retinal Research Department, Manhattan Eye, Ear & Throat Hospital, 210 East 64th Street, New York, NY 10021, USA. e-mail:
[email protected] Lasers in Ophthalmology – Basic, Diagnostic and Surgical Aspects, pp. 255–275 edited by F. Fankhauser and S. Kwasniewska © 2003 Kugler Publications, The Hague, The Netherlands
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cells.126-128 Finally, by its direct effect on ion transport, cortisol may be responsible for the reverse polarity of the RPE cells, causing them to secrete ions into the subretinal space. In turn, the ions may cause osmotic fluid attraction and serous macular detachment as a result of increased fluid inflow.129131 This hypothesis, implicating a disturbance in the autoregulation of choriocapillaris blood flow linked to abnormal circulating levels of cathecolamine or cortisol, is supported by experimental CSC animal models.132-137 Another possible theory suggests that a protracted circulatory disturbance in the microcirculation of the choriocapillaris leads to increased fluid leakage in the sub-RPE space. To start with, the RPE cells are able to maintain their integrity and function, as well as being able to pump in a retinal-choroidal direction, thus keeping the subretinal space dry. However, the prolonged excessive stress on the RPE cells ultimately causes generalized damage to the cells and loss of function. A combination of the increased fluid leakage from the choriocapillaris on one side and the impaired function of the RPE cells on the other ultimately leads to pooling of fluid in the sub-RPE space, loss of continuity of the RPE layer, and pinpoint leakage of fluid in the subretinal space. The loss of function of the contiguous RPE cells allows the fluid to accumulate in the subretinal space and causes a neurosensory detachment. When new RPE cells substitute the damaged ones, there is spontaneous healing and reabsorption of fluid. Perpetuating microcirculatory disturbance leads to new damage and recurrence of the process. Further clinical and pathological studies are required to support the above-mentioned hypothesis further; yet, this theory provides an excellent explanation for some of the clinical characteristics of CSC: diffuse choroidal hyperpermeability in areas that appear clinically normal, recurrence of the disease, progressive damage to the RPE layer, and multifocal pinpoint areas of leakage in the more severe and chronic forms of CSC. Clinical features Demographics Greater understanding of the clinical manifestations of CSC has considerably changed our knowledge of the demographics of the disease. Until a few years ago, CSC was considered predominantly a disease of males aged between 30 and 50 years. No case of CSC in a patient under the age of 20 years has been reported in the literature, the only exception being a seven-year-old girl who evidentially have had posterior scleritis.32 The overall incidence in males versus females in numerous reports is approximately eight or nine to one,2,33-40 but the incidence in women is noted to double between the ages of 31 and 40 years compared to between the
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ages of 21 and 30 years.41 Duke-Elder and Dobree felt it was appropriate to diagnose the condition in patients under the age of 50 years, since, in patients over 50 years of age, distinguishing CSC from some forms of age related macular degeneration (AMD) would be extremely challenging.42-44 Very little is known about the age-specific prevalence and the clinical findings of CSC in older adults. In fact, neurosensory macular detachment in an adult over the age of 50 years suggests the presence of choroidal neovascularization (CNV) secondary to AMD, which is the most common cause of blindness in older adults.141 In 1996, Spaide et al. reported a study conducted on 130 older patients with neurosensory macular detachment.6,45 Over half the subjects were 50 years of age or older and 57 were diagnosed with CSC after they had turned 50. The male:female ratio in the study population was 2.6:1.0, with no significant differences among the age groups. Thus, although still typical of young adult males, CSC could also be manifested in females and older adults. Risk factors It has been speculated that psychological factors might play a causative role in CSC.46-51 Type A behavior and its psychological consequences, most notably, a sympathetic discharge, can adversely stimulate the eye as an organ system and the macula as an ultimate target area.7 Werry and Arends used the Minnesota Multiphasic Personality Inventory (MMPI) to compare patients with CSC with an age- and sex-matched group of healthy individuals.52 They found that patients with CSC showed significantly higher values on the hypochondria and hysteria scale. Another risk factor is the refractive state of the eye. Most CSC cases are diagnosed in patients with no refractive errors or a mild degree of hyperopia.7,53,54 There may also be a racial predisposition, with a higher incidence among Caucasians, Hispanics, and possibly Asians, and an extremely low occurrence in African-Americans.7,46,55-57 The more severe form of CSC occurs most frequently in individuals from south Asia or in those of Latin origin.34 Severe forms of CSC have been associated with pregnancy,24-31 end-stage renal disease (ESRD),18,73,74 organ transplant,17,19 increased endogenous cortisol production (Cushing’s disease),11,12 the use of inhaled nasal corticosteroids,20 systemic corticosteroid treatment,13-23 and epidural corticosteroid injection.75 We have seen one patient with ESRD treated with systemic corticosteroid who developed a severe, bilateral variant of CSC, complicated by bilateral giant rips of the RPE and permanent loss of vision. Serous retinal and RPE detachments resembling CSC have also been described in three patients with paraproteinemias and in a patient using methylenedioxymethamphetamine (MDMA) or ‘ecstasy’.76,77
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Central serous chorioretinopathy There is a number of systemic diseases characterized by ischemic disorders of the choroid which may present with PEDs and bullous neurosensory detachments resembling CSC, such as systemic lupus erythematosus (SLE),142-150 polyarteritis nodosa,151,152 Goodpasture’s syndrome,153 Wegener’s granulomatosis,154 accelerated hypertension,155-157 toxemia of pregnancy,158-160 disseminated intravascular coagulopathy (DIC),161-163 and thrombotic thrombocytopenic purpura (TTP).152 It is difficult to judge whether the disease, its treatment, or both, cause CSC. Although these diseases have different pathogenesis and systemic manifestations, they all can produce serous retinal detachments. We think that the final common pathway is ischemia of the choriocapillaris caused either by vasospasm (e.g., malignant hypertension, toxemia of pregnancy), intravascular deposition of clots (e.g., DIC, TTP) or intravascular precipitation of circulating immunocomplexes (SLE and other collagen vascular diseases). The choriocapillaris ischemia may finally lead to damage of the retinal pigment epithelium and subretinal fluid exudation. Symptoms If the neurosensory detachment does not involve the central macula, the patient is usually asymptomatic and the detachment may resolve spontaneously. When the neurosensory detachment involves the fovea, the patient becomes symptomatic. Many patients first notice a minor blurring of vision, followed by the various degrees of metamorphopsia, micropsia, chromatopsia, central scotoma, loss of contrast sensitivity and increasing hyperopia. Usually, the area of metamorphopsia is reproducible on Amsler grid testing. Delay recovery to light stimulation detectable with the Photo Stress Test can also occur. Visual acuity in the acute stages ranges from 20/20 to 20/200.58,59 Vision can usually be improved with a small hyperopic correction. In some patients, the onset of symptoms is preceded or accompanied by migraine-like headaches.33 Examination The anterior segment and vitreous remain normal. Biomicroscopic examination with a fundus contact lens reveals a macular neurosensory detachment. The detached neurosensory retina is usually transparent and of normal thickness. A yellowish discoloration is sometimes discernible in the foveal area, caused by increased visibility of the macular xanthophyll pigment. Multiple white-yellowish dotlike deposits cover the posterior surface of the retinal detachment in some cases (Fig. 1). A small serous pigment epithelium detachment (PED) may be present under the superior half of the neurosensory macular detachment. The PED appears as a round or oval area of detached RPE, yellow or gray in color, usually not larger than one-fourth of a
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257 disc diameter. Long-standing and recurrent PEDs may present with pigment migration or atrophy. The pigment clumping may produce a cruciate or triradiate pigment figure. Although PEDs are usually small in CSC, some patients may develop one as large as a disc diameter, or even more. In these cases, the PED appears as a dome-shaped, well delineated, gray-yellowish elevation. A pink halo may surround the PED, caused by shallow separation of the retina at the edge of the PED. Some clinical findings help in differentiating a larger PED from a neurosensory detachment. Usually, the PED is better delineated, surrounded by a light reflex, and does not allow appreciation of details of the underlying choroid when compared to a neurosensory detachment. Sometimes, two or more PEDs are present, and, in some cases, the PED is noted above the neurosensory detachment, due to gravity forcing the fluid inferiorly. In this situation, there appears to be a lack of continuity between the PED and the neurosensory detachment (Fig. 2). Although the subretinal fluid is usually transparent and allows clear visualization of the underlying RPE and choroidal details in some patients, it may also be cloudy and white-grayish in color, presumably because of the presence of fibrin exudation (Fig. 3).34 There must be significant hyperpermeability of the choriocapillaris, which allows a large molecule (fibrin, 340,000 Daltons) exudate in the extravascular space. It is possible that these patients have a blow out of the RPE at or near the border of the RPE detachment, allowing the abnormal egress of fluid and protein under the retina (Fig. 4).61 The subretinal fluid becomes progressively more opaque with increasing concentration of fibrin. Sometimes, there is a dark spot inside or on the edge of the fibrinous plaque which corresponds with the PED causing the RPE leakage. The fluid leaking from the choriocapillaris has a very low concentration of fibrin molecules, and thus it is transparent. The clear fluid leaking in the subretinal space allows a clear view of the underlying RPE detachment, which is usually reddish or dark in contrast to the surrounding cloudy subretinal fluid. While, in most cases, the fibrin deposit simply disappears, in a few cases, it might stimulate subretinal fibrosis and scar formation. This may cause permanent visual loss and become complicated by subretinal neovascularization and RPE rips.60,61 The conditions predisposing to subretinal fibrin exudation include large and multiple PEDs, chronic and recurrent disease, pregnancy,24 systemic corticosteroid use, organ transplant, diabetes, and male gender. In patients with chronic subretinal fluid, it is also possible to observe subretinal lipid deposition.61,62 Since subretinal lipid deposits are more typical of conditions causing chronic subretinal exudation such as AMD and polypoidal choroidal vasculopathy (PCV), it is important to recognize that they may also be present in CSC, and do not represent an exclusion criterion to the diagnosis of CSC as once was believed.
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Fig. 1. A 34-year old Caucasian male who presented complaining of blurred vision in his left eye of a week’s duration. A. Clinical photograph of the left eye showing serous neurosensory macular detachment. B and C. Fluorescein angiography study demonstrating a pinpoint area of hyperfluorescence in the central macula leading to the characteristic smokestack configuration in the late phase. A pigment epithelial detachment temporal to the fovea increased in hyperfluorescence throughout the examination (white arrows).
Fig. 2. Optical coherence tomography study of pigment epithelium detachment in a patient with CSC, confirming the presence of localized elevation consistent with serous PED.
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Central serous chorioretinopathy Clinically-discernible, peripheral-dependent bullous neurosensory detachments have been described in patients with CSC (Fig. 5).62-70 Yannuzzi et al. first characterized the presence of the RPE atrophic tracts extending inferiorly in the fundus periphery, secondary to the antecedent retinal detachment in patients with CSC.62 Presumably, there is a particularly severe and/or prolonged time leakage of fluid from an RPE defect in the subretinal space at the posterior pole. The subretinal fluid gravitates inferiorly to form a dependent neurosensory detachment in a ‘flask’, ‘teardrop’, ‘dumbbell’, or ‘hourglass’ pattern. Sometimes, the tract of subretinal fluid connecting the macular detachment with the bullous neurosensory detachment in the inferior hemisphere is so shallow that it is very difficult to appreciate. The RPE under the chronic retinal detachment experiences atrophic changes that appear as atrophic RPE tracts connecting the posterior pole with the dependent retinal detachment. The retina itself develops secondary manifestations, including pigment migration, capillary dilatation (telangiectasia) proximally, and capillary nonperfusion (is-
259 chemia) distally, to the area of detached retina. The changes in the RPE consist of both atrophy and pigment clumping in the form of perivascular deposits or bone spicules, the condition described by Gass as “pseudoretinitis pigmentosa-like atypical CSC presentation”.69 Angiography Fluorescein angiography The typical angiographic finding, occurring in about 95% of all cases of CSC, include the presence of one or more leakage point in the RPE. In the majority of these cases, the dye spreads to all sides slowly and evenly, staining the subretinal blister. Even though the initial diffusion of the dye occurs rather quickly, it may not reach the borders of the blister until the late phases of the angiogram. Interestingly, no diffusion or staining of the retina beyond the edges of the detachment is observed. In 7-20% of all cases, the dye enters the blister
Fig. 3. A 36 year-old male who presented with decreased visual acuity in his right eye. A. Color photograph of the right eye showing serous neurosensory detachment (white arrows) in the superior area of the macula with ring-like yellowish subretinal nodular deposits consistent with fibrin surrounding the localized PED (black arrows). B. Early-phase fluorescein angiography study revealing localized area of leakage corresponding to serous PED. C. Late-phase fluorescein angiography demonstrating hyperfluorescence due to pooling beneath the serous PED. D. Late-phase indocyanine green angiogram showing an area of hyperfluorescence corresponding to the serous PED and staining of the subretinal fibrin deposits.
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Fig. 3. A 36 year-old male who presented with decreased visual acuity in his right eye. E and F. The fellow eye, appearing normal both clinically and on fluorescein angiography. G. Late-phase ICG angiogram revealing multiple areas of hyperfluorescence consistent with diffuse choroidal vascular hyperpermeability and bilateral disease.
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Fig. 5. A 47-year-old female with an 18-year history of CSC in both eyes. A. Color photograph composite of the left eye showing bullous dependant detachment of the neurosensory retina inferiorly. B. Fluorescein angiogram composite revealing diffuse decompensation of the retinal pigment epithelium, multiple scattered PEDs, and obliteration of the retinal capillaries in the region of the detachment. Note the presence of early neovascularization at the junction between perfused and non-perfused retina. C. Clinical photograph of the left eye showing PED superior to the optic disc partially surrounded by fibrin deposits.
← Fig. 4. Hypothesis of the pathophysiology of subretinal fibrin deposition in CSC. A. Disturbance in the choriocapillaris blood flow, leading to accumulation of choroidal exudate (green dots). B. PED formation. C and D. RPE leakage through a defect, reaching the subretinal space.
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Fig. 5. A 47-year-old female with an 18-year history of CSC in both eyes. D. Fluorescein angiography study confirming the presence of active leakage from the serous PED. E. Color photograph composite of the same eye two months after laser treatment of the site of leakage revealing partial resolution of the detachment and lipid precipitation.
through a single leakage point, but, instead of spreading evenly to all sides, it first ascends in a so-called ‘smokestack’ phenomenon. Upon arrival at the upper limit of the blister, it expands laterally in a mushroom- or umbrella-like fashion (Fig. 6).36-38 Shimizu and Tobari, who first described this phenomenon, believe that it is due to an osmotic pressure gradient resulting from different protein concentrations in the content of the blister and
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in the liquid entering through the leakage point.78 There are usually one or two leakage points, but there may be as many as seven or more.35,36,38 In some rare cases, the mechanical defect at the margin of the PED may be apparent as a puncture or ‘blow out’. Large PEDs with multiple points of ‘blow out’ at their margin have also been described.79 Although usually the exudate extends into or be-
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Fig. 5. A 47-year-old female with an 18-year history of CSC in both eyes. F. Clinical photograph composite 16 months after laser treatment in the area of leakage showing complete resolution of the detachment and partial reperfusion to the inferior retina.
yond the fovea, it is interesting to note that the leakage point is found in the foveal area in less than 10% of all cases; as a matter of fact, it is more frequently located in a 1-mm-wide ring-like zone immediately adjacent to the fovea.80,81 The incidence rapidly decreases beyond this ring. The greatest overall incidence of leakage point is in the upper nasal quadrant, followed by the lower nasal quadrant, upper temporal quadrant, and lower temporal quadrant.2,38-40 These topographic differences are unexplained, since neither the neurosensory retina nor the neighboring structures, such as RPE, Bruch’s membrane, or choriocapillaris, show any regional differences that might account for this phenomenon. In 18-30% of cases, the leakage point lies in the area of the papillomacular bundle where the incidence above the midline is almost twice that below this line.2,38-40 Interestingly, the exudate tends to extend into and to involve the fovea, even if the leakage point is located eccentrically. This may be explained by the fact that, in the fovea, the rods, and thus the strong adhesion between neurosensory retina and RPE, are missing.80 In some cases, no leakage point is seen under the retinal detachment. In these cases, the area superior to the macular detachment should be studied because gravity may have caused the leaking point to remain outside the detached area. In
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other cases, the leaking point cannot be found because the area has healed. Indocyanine green angiography The application of ICGA to the study of CSC has expanded our knowledge of the disease.5,6,82-95 The common findings in patients with CSC are multifocal areas of hyperfluorescence in the early and mid phases of the study, which tend to fade in the late phases (Fig. 7). Typically, these areas of hyperfluorescence are found, not only in correspondence with the leaking point seen on FA, but also in fundus areas, which appear clinically and angiographically normal, as well as in normal fellow eyes (Fig. 8). The areas of early hyperfluorescence are believed to represent diffuse choroidal hyperpermeability. On wide angle ICGA, we have been able to note multifocal hyperpermeable areas and presumed ‘occult’ PEDs, which extend far beyond the posterior pole.94 Although further pathological correlations are necessary, these findings suggest that PEDs may be much more common in CSC than previously believed. In addition, PEDs may be found in eyes that seem inactive on clinical examination and on FA. Both the multifocal hyperfluorescent spots seen throughout the fundus and the presumed occult PEDs that are only noted on ICGA
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Fig. 6. A 41-year-old male with the diagnosis of CSC in both eyes for ten years presented complaining of a sudden decrease of vision in the left eye. A. Red-free photograph of the right eye revealing pigmentary changes temporal to the fovea. B. Red-free photograph of the left eye showing a well-circumscribed neurosensory detachment of the macula with two areas of focal PED (white arrows). C, D, E, and F. Fluorescein angiography study showing two localized areas of PEDs and typical ‘smokestack’ appearance of the dye leaking under the detached retina. Note the dye expanding in an umbrella-like fashion once it reaches the upper limit of the detachment.
emphasize the fact that this disease may be more diffuse and widespread than previously believed. Optical coherence tomography Optical coherence tomography (OCT) is another
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diagnostic tool for high-resolution cross-sectional imaging of the retina. It is useful for the quantitative examination of patients with central serous chorioretinopathy and for objectively monitoring the clinical course of the serous retinal detachment in this disease.166 Typically, in the acute phase of
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265 central vision loss, and significant superior visual field loss. Other complications noted in these patients are cystoid macular edema (CME) and cystoid macular degeneration (CMD). Intraretinal cystoid spaces detected by OCT but without intraretinal leakage, or CMD, was a common finding in eyes with chronic CSC and reduced central vision after resolution of subretinal fluid. Chronic foveal detachment and antecedent intraretinal leakage were proposed to be the mechanisms for the development of these changes. In conjunction with foveal atrophy, CMD was an important clinical finding to account for the poor visual outcome in patients with CSC.
Fig. 6. A 41-year-old male with the diagnosis of CSC in both eyes for ten years presented complaining of a sudden decrease of vision in the left eye. G. Mid-phase fluorescein angiography study of the fellow eye showing window defect hyperfluorescence corresponding to the RPE tract that extends inferiorly. Note a pinpoint nasal leakage superior to the nerve (white arrow).
Differential diagnosis
CSC, patients present with thickening of the retina and areas of low reflectivity localized within the detached retina. Sometimes, it is also possible to detect a moderately reflective mass bridging the detached neurosensory retina and RPE, which seems to be a fibrinous exudate that accumulates in the subretinal space and infiltrates into the outer retina (Fig. 9).165 In patients who develop chronic CSC, OCT is extremely helpful in detecting shallow detachments and intraretinal cystoid spaces.
Infectious and inflammatory disorders
Natural history When left alone, CSC generally heals spontaneously within 12 weeks, with complete recovery of visual acuity. However, recurrences are frequent and occur in about one-third to one-half of all patients after the first episode of the disease, and 10% of all affected individuals have three or more recurrences.4,33,38 In almost half the patients with recurrent disease, the recurrence occurs within one year of the primary episode, but relapses may occur up to ten years later.96,97 Although long-term follow-up data from patients with CSC are not available, there is evidence that even a single episode may be followed by a chronic, slowly progressing disturbance of the retinal pigment epithelium in the posterior pole. Some eyes may suffer from persistent and progressive macular detachment with the associated visual decline. A small percentage of patients will develop CNV, perifoveal RPE atrophy, choriocapillaris atrophy secondary to the RPE damage in the areas of RPE tracts, and subretinal lipid deposition.71 This severe variant of CSC appears to be more frequent in patients of Latin or Asian ancestry, and is usually associated with frequent recurrences, permanent
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Although the clinical diagnosis of CSC is usually confirmed by fluorescein angiography, several entities should be considered in the differential diagnosis.
Similarly to CSC, Harada’s disease may present with serous macular detachments. The presence of anterior uveitis, vitritis and optic disc hyperemia, the associated systemic manifestations, and the prompt response to anti-inflammatory therapy, are helpful in differentiating Harada’s disease from CSC. Posterior scleritis can also present with exudative neurosensory detachment at the posterior pole, but it can be distinguished from CSC by the presence of scleral thickening, vitreous cells, and pain on ocular movements. The diagnosis can be confirmed on ultrasound examination by the presence of an echolucent area at the posterior pole behind the echo of the sclera (T sign, typical of posterior scleritis). Sympathetic ophthalmia may also present with serous macular detachment, but the associated intraocular inflammation, yellowish cellular detachments of the RPE (Dalen-Fuch’s nodules), and the history of trauma to the fellow eye allows easy differentiation of the two diseases. Two additional conditions, which may present with exudative neurosensory macular detachment in otherwise healthy subjects, are idiopathic uveal effusion syndrome and benign reactive lymphoid hyperplasia.138-140 Tumors Choroidal melanoma, choroidal hemangioma, choroidal metastasis, choroidal osteoma, and leukemic choroidal infiltrates can present with the exudative macular detachments. While clinical examination of the fundus is usually sufficient for recognizing the choroidal tumor, some tumors, especially choroidal hemangiomas, can be confused with large PEDs with associated neurosensory detachment. In
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Fig. 7. A 43-year-old Caucasian female with chronic CSC in the left eye. A, B and C. ICG angiography is essentially normal in the early phase, but reveals multiple patchy areas of hyperfluorescence in the mid phase, which fades in the late phase of the study.
these cases, ultrasound examination and angiography may be helpful in determining the correct diagnosis. Vascular disorders Collagen vascular disease such as SLE, polyarteritis nodosa, scleroderma, dermatomyositis, and relapsing polychondritis, may be associated with the serous detachments of the macula caused by fibrinoid necrosis of the choroidal vessels. In these diseases, CSC may arise as a result of the prolonged systemic use of corticosteroids. Malignant hypertension, toxemia of pregnancy, and disseminated intravascular coagulopathy, can present with a neurosensory detachment secondary to acute multifocal occlusion of the choroidal arteries and choriocapillaris, and necrosis of the overlying RPE (Elshing’s spots).
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Optic nerve pit with serous macular detachment Clinically, a pit should be looked for in all patients with CSC. On FA, there is no pinpoint leakage of the dye or slow filling of the macular detachment as seen with CSC. OCT may be helpful in imaging the schisis-like separation of the retinal layers at the posterior pole and the presence of the optic disc pit or any other optic nerve abnormalities, which can lead to a neurosensory detachment.164 Optic nerve sheath meningocele Optic sheath meningocele is a rare disease. The term ‘optic sheath meningocele’ was recently proposed by Garrity and Forbes to describe primary CSF cysts of the optic nerve sheath, without apical mass or malformation of the cranio-orbital junction.167 Presenting symptoms are often related to involvement of the optic nerve, with a slow or rapid decrease of visual acuity. Sometimes there is also a serous macular detachment, simulating CSC. CT
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Fig. 8. An 87-year-old male with a history of longstanding CSC in the left eye and a visual acuity of 20/100. A. Color photograph composite showing diffuse pigmentary changes at the posterior pole. B. Fluorescein angiogram composite revealing diffuse decompensation of the RPE.
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Fig. 8. An 87-year-old male with a history of longstanding CSC in the left eye and a visual acuity of 20/100. C. Mid-phase ICG angiography illustrating multiple patchy areas of hyperfluorescence in the left eye. Note that the areas of hyperfluorescence on ICG do not correspond to the areas of leakage shown on fluorescein angiography.
Fig. 9. OCT study in a patient with CSC confirming neurosensory detachment of the macula and a shallow PED in the central area (arrow).
and MRI studies reveal a tubular-cystic enlargement of the optic nerve/optic sheath complex, with thickening of the optic nerve. Radiological differential diagnosis should include optic nerve tumors,
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such as gliomas, meningiomas, and arachnoid cysts involving the optic nerve sheath.167
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Central serous chorioretinopathy Age-related macular degeneration The differentiation between age-related macular degeneration (AMD) and CSC is necessary when CSC is present in adults over 50 years of age. While FA is helpful in distinguishing the two conditions when there is well-defined CNV secondary to AMD and a well-defined pinpoint leakage in CSC, it is much less helpful in the presence of diffuse, ill-defined hyperfluorescence, which may be caused by both occult CNV or a diffuse ‘ooze’ of RPE. In such cases, ICGA may be helpful in showing multifocal early hyperfluorescence, fading in the late phase in CSC, and late hyperfluorescence, corresponding to the CNV in AMD.5 Polypoidal choroidal vasculopathy Polypoidal choroidal vasculopathy (PCV) is presumed to be a variant of occult CNV. The typical or classic clinical presentation of PCV poses little challenge. However, an isolated macular variant may have clinical and FA characteristics resembling CSC (Fig. 10). These atypical cases of PCV involve small-caliber vascular abnormalities and may present exclusively with detachment of the neurosensory retina. The polypoidal lesions may clinically and fluorescein angiographically resemble small PEDs. These cases of PCV may masquerade as CSC (Fig. 11).98 Vitelliform macular detachment Patients with cuticular or basal laminar drusen and vitelliform foveomacular or pattern dystrophy may present with visual loss caused by yellow serous exudative detachment of the retina in the macular area in one or both eyes. These neurosensory macular detachments, which resemble the lesions seen in Best’s vitelliform dystrophy, may be mistaken for serous detachments of the RPE. Fluorescein angiography is helpful in differentiating these conditions from CSC. In the early phases of the study, the yellowish subretinal fluid obstructs the fluorescence of the underlying choriocapillaris, while later there is patchy filling of the subretinal fluid with dye through the RPE defects. Treatment Although, thus far, no medication has proven effective in CSC, and the disorder generally resolves spontaneously, a beneficial effect of photocoagulation has been shown in several studies.2,99,103-115 Some authors have also reported that treatment of CSC with photocoagulation reduces the recurrence rate,97,106,114,115 whereas others109,110,116-118 have observed no such difference. Because CSC is usually a self-limiting disease, and there are possible complications associated
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269 with laser treatment, particularly when applied very close to the fovea, the recommendation for photocoagulation therapy can only be based on clinical judgment. As a general rule, we recommend observing any new onset of acute serous macular detachment for the first three months, unless the patient has special occupational reasons, which require rapid improvement of visual acuity, or if the affected eye is the only eye. If the detachment does not resolve itself after the first three months, and the leakage point is remote from the center of the fovea, it is reasonable to treat a symptomatic patient. If the leakage point is within 500 µm from the center of the fovea, six months of observation is recommended before undertaking treatment. Other indications for laser treatment are primary detachment with visual decline in a patient who has experienced permanent visual loss from an untreated macular detachment in the fellow eye and recurrent macular detachment in the eye that has experienced permanent visual loss from the initial episode. Furthermore, laser treatment is indicated in particularly severe forms of CSC, which are known to have poor prognosis if left untreated, such as CSC complicated by multiple serous detachment of the RPE and bullous sensory retinal detachment, dependent neurosensory detachment, epithelial tracts, diffuse RPE decompensation, subretinal deposits of fibrin and lipids, and those associated with secondary CNV. A recent fluorescein angiogram showing the area of RPE leakage should be positioned in a viewer for guidance at the time of laser treatment. A fundus contact lens is used for examination of the posterior pole and for laser treatment. While the laser is in standby, the point of fixation is identified by asking the patient to fixate on a target light and to hold the fixation while the target light is moved (either the smaller light beam of the slit lamp or the laser aiming beam are appropriate). Once the fixation point has been clearly identified, the RPE leakage point to treat is identified with the help of the fluorescein angiography picture (known fundus landmarks, such as foveal avascular zone and vessel crossing sites, are used to define the exact position of the area to be treated).111,113 The spot size should be slightly larger than the leaking point, usually 200 µm. Although, for very small leaks, just one spot may be sufficient, more laser spots (three to five) are usually necessary to complete the treatment. It is better to start the treatment with a low intensity (100 mW) and a relatively long duration (0.2-0.3 seconds), and to use a slight increase of intensity if necessary, looking for a slight gray discoloration of the treated area. The laser beam is focused at the level of the RPE in order to obtain a slow, mild intensity burn, the goal of treatment being a faint gray discoloration of the RPE and not a frank white reaction, as in laser treatment of CNV. After laser photocoagulation treatment of the leak, anatomical resolution of the
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Fig. 10. A 56-year-old Caucasian male who had had three transient episodes of vision disturbance diagnosed as CSC. A. Clinical photograph of the macula is flat, overlying multiple, nummular elevations, suggestive of small serous PEDs. At the center of the lesion, a patch of fibrous metaplasia (white arrow) can be seen. B. Fluorescein angiogram revealing a net of subretinal inner choroidal vessels terminating in aneurismal or polypoidal lesions. C. Late ICG angiogram confirming the presence of polypoidal vascular abnormality.
Fig. 11. A. Clinical photograph of a 62-year-old female with neurosensory retinal detachment in the central macula. B. ICG angiogram revealing the presence of a polypoidal choroidal vascular abnormality in the superior temporal juxtapapillary region.
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Central serous chorioretinopathy macular detachment generally occurs in about two weeks in uncomplicated cases, but it may require up to six weeks in longstanding detachments with turbid subretinal fluid. Complete visual recovery usually requires twice that time. In chronic, severe forms of CSC, there is usually diffuse decompensation of the RPE, which appears on FA as multifocal pinpoint areas of leakage or diffuse ‘ooze’ of the RPE. In such cases, even if its value has not been proved in clinical trials, there is a rationale for applying a grid of laser treatment to cover the area of RPE decompensation completely. Complications Although rare, there are some recognized complications in laser treatment of CSC, the worst of which is inadvertent photocoagulation of the fovea. The patient should always be informed that a persistent scotoma corresponding to the site of laser photocoagulation might be experienced after treatment. Secondary choroidal neovascularization may be induced by the laser treatment, especially when excessive intensity is used.103,109,119,120 An often underestimated complication is the slow but progressive enlargement of the area of RPE atrophy caused by the laser treatment. When the treatment site is close to the center of the fovea, the enlargement of the RPE scar may eventually involve the fovea and cause delayed irreversible visual loss. Whenever possible, the laser treatment of a leaking point in CSC should be avoided within the foveal avascular zone. Recently, there have been case series reports on the use of photodynamic treatment (PDT) with verteporfin for chronic CSC. The rationale behind such a therapeutic approach is that of causing a reduction of blood flow in the choriocapillaris, which is hyperpermeable in CSC.121-123 This therapeutic approach would appear to be promising for chronic and recurrent cases of CSC, especially when there is diffuse decompensation of the pigment epithelium.
Conclusions Although the etiology of central serous maculopathy is not entirely clear, its pathogenic mechanism is supposed to disturb the choriocapillaris and blood flow, leading to a choroidal exudate, and to pigment epithelium detachment, resulting in local disruption of this structure, and to fluid leakage under the sensory retina. The application of indocyanine green shows multifocal areas of hyperfluorescence in the early and mid phases, which tend to fade in the late phase. Optical coherence tomography (OCT) is an important tool for the high resolution of this disease. Although, thus far, no medication has proven effective in central sero ma-
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Central serous chorioretinopathy 154. Kihyoun JL, Kalina RE, Klein ML: Choroidal involvement in systemic necrotizing vasculitis. Arch Ophthalmol 105:939-942, 1987 155. Klien BA: Ischemic infarcts of the choroid (Elshing’s spots): a cause of retinal separation in hypertensive disease with renal insufficiency: a clinical an histopathological study. Am J Ophthalmol 66:1069-1088, 1968 156. Stropes LL, Luft FC: Hypertensive crisis with bilateral bullous retinal detachment. JAMA 238:1948-1949, 1977 157. Venecia G, Jampol LM: The eye in accelerated hypertension. II. Localized serous detachments of the retina in patients. Arch Ophthalmol 102:68-73, 1984 158. Gitter KA, Houser BP, Sarin LK, Justice J Jr: Toxemia of pregnancy: an angiographic interpretation of fundus changes. Arch Ophthalmol 80:449-454, 1968 159. Gass JDM, Pautler SE: Toxemia of pregnancy pigment epiteliopathy masquerading as heredomacular dystrophy. Trans Am Ophthalmol Soc 83:114-130, 1985 160. Menchini U, Lanzetta P, Virgili G, Ferrari E: Retinal pigment epithelium tear following toxemia of pregnancy. Eur J Ophthalmol 5:139-141, 1995 161. Yamaguchi K, Abe S, Shiono T et al: Macular choroidal
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275 occlusion in dysplamogenemia. Retina 11:423-425, 1991 162. Cogan DG: Fibrin clots in the choriocapillaris and serous detachment of the retina. Ophthalmologica 172:298-307, 1976 163. Cogan DG: Ocular involvement in disseminated intravascular coagulopathy. Arch Ophthalmol 93:1-8, 1975 164. Postel EA, Pulido JS, McNamara JA, Johnson MW: The etiology and treatment of macular detachment associated with optic nerve pits and related anomalies. Trans Am Ophthalmol Soc 96:73-88, 1998 165. Iida T, Hagimura N, Sato T, Kishi S: Evaluation of central serous chorioretinopathy with optical coherence tomography. Am J Ophthalmol 129(1):16-20, 2000 166. Hee MR, Puliafito CA, Wong C, Reichel E, Duker JS, Schuman JS, Swanson EA, Fujimoto JG: Optical coherence tomography of central serous chorioretinopathy. Am J Ophthalmol 120(1):65-74, 1995 167. Lunardi P, Farah JO, Ruggeri A, Nardacci B, Ferrante L, Puzzilli F: Surgically verified case of optic sheath nerve meningocele: case report with review of the literature. Neurosurg Rev 20(3):201-205, 1997
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Scanning laser polarimetry of the retinal nerve fiber layer
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Scanning laser polarimetry of the retinal nerve fiber layer in the detection and monitoring of glaucoma Christopher Bowd, Linda M. Zangwill and Robert N. Weinreb Diagnostic Imaging Laboratory, Hamilton Glaucoma Center and Department of Ophthalmology, University of California, San Diego, La Jolla, CA, USA
Keywords: glaucoma, nerve fiber layer, polarimetry, birefringence
Introduction Identifying retinal nerve fiber layer (RNFL) abnormalities and their progression is of vital importance for the diagnosis and monitoring of glaucoma. However, until recently, this task has been subjective, and descriptions of change have been primarily qualitative. Recently, techniques have been developed to objectively measure correlates of RNFL thickness in intact human eyes. One such technique, scanning laser polarimetry (SLP), has rapidly disseminated to the clinics of glaucoma specialists worldwide. Although this technique is relatively easy to use and provides quantitative measurements of the RNFL, it is not yet apparent how the large amount of information provided by this technology should be interpreted clinically. Further, research to determine whether this methodology is effective for monitoring glaucomatous change over time is in its early stages. The modern scanning laser polarimeter is an advanced version of the prototype Fourier ellipsometer developed by Weinreb et al. a decade ago.1 The prototype ellipsometer employed a 514-nm laser source, while the current instrument employs a 780-nm source. SLP exploits the birefringent property of the human RNFL which originates from the parallel organization of the microtubules that support retinal ganglion cell axons.2,3 When light from a polarization modulated laser beam (780-nm diode), with its optic axis parallel to the surface of the birefringent RNFL, is focused on the retina by the optical media of the eye and double-passes the RNFL, the light is split into two beams with different polarization axes travelling at different veloci-
ties. This difference in velocities results in a phase shift, or relative retardation, of the existing beams, resulting in a change in polarization of the reflected light. The polarization axis of the beam is changed (retarded) by an amount dictated by the thickness of the RNFL.4,5 The thicker the birefringent structure, the greater the retardation of transmitted light. Therefore, this technique provides a direct measurement of light retardation that is reportedly linearly related to RNFL thickness in histological preparations of primate retinae.1,6 Retardation values are transformed into indirect measurements of RNFL thickness, based on this known linear relationship. A complete SLP scan measures retardation at 256 × 256 retinal positions, thus acquiring 65,536 data points in less than one second. Three individual scans are usually combined to create a composite mean. The majority of scientific and clinical studies investigating SLP technology have used the Laser Diagnostic Technology (San Diego, CA) GDx or its compatible predecessor, the Nerve Fiber Analyzer (NFA) II. Recently, a more automated, compact device; the Access (Laser Diagnostic Technology), has been introduced. A number of studies have shown significant differences in GDx/ NFA II-measured RNFL thickness between healthy and glaucomatous eyes.7-13 For the clinician’s purposes, a summary GDx printout is available that provides RNFL thickness parameters obtained by the instrument. Information provided on the summary printout includes a reflectance map, a retardation map (also called a thickness map) illustrating RNFL thickness using color coding, a graph depicting RNFL thickness within a
Grant support: NIH EY11008 (LMZ) and Joseph Drown Foundation (RNW). Address for correspondence: Robert N. Weinreb, MD, Hamilton Glaucoma Center, University of California, San Diego, 9500 Gilman Drive, La Jolla, CA 92093-0946, USA. e-mail:
[email protected] Lasers in Ophthalmology – Basic, Diagnostic and Surgical Aspects, pp. 277–284 edited by F. Fankhauser and S. Kwasniewska © 2003 Kugler Publications, The Hague, The Netherlands
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Fig. 1. GDx extended analysis printout from a healthy eye showing the reflectance map, retardation or thickness map, graph depicting RNFL thickness within the measurement ellipse, quadrant specific deviation from normal thickness values, and nerve fiber analysis summary data.
measurement ellipse that circumscribes the optic nerve head, a table that provides quadrant specific deviation from normal thickness values (relative to a large normative database), several summary values (NFA parameters) with corresponding probability values compared to the normative database, and a neural network derived value (The Number) that indicates the likelihood of glaucoma in the examined eye (values for this parameter range from 0 or very unlikely glaucoma, to 100 or very likely glaucoma). Figure 1 shows a GDx printout from a healthy eye.
between operators.14-16 One reproducibility study using the GDx showed that consecutive RNFL thickness measurements obtained from healthy eyes differ by only about 8 µm, or about 10%.17 Values are slightly higher in glaucoma eyes. The relatively low variability observed between operators when imaging the same patients suggests that using different technicians for follow-up examinations is not likely to be a cause for concern. Nonetheless, although GDx images are relatively easy to obtain, adequate training of technicians and quality control assurance are necessary to obtain reliable data.
Reproducibility
Retinal nerve fiber layer thickness in healthy and glaucoma eyes
For a new disease diagnostic/monitoring instrument to be effective, it must provide similar measurements over time in the same subjects. That is, its measurements must be reproducible. Reproducibility of measurements must be good both within and
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Briefly, in primary open-angle glaucoma, the retinal ganglion cell population is decreased, probably due to changes in neurotrophic mechanisms, effects of increased intraocular pressure (or increased sus-
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Scanning laser polarimetry of the retinal nerve fiber layer
Fig. 2. GDx measured average RNFL thickness in 134 healthy eyes. The horizontal line depicts the mean value of 66.9 µm (standard deviation = 11.1 µm). The range of values is 47-100 µm.
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approximately 50-100 µm.22,23 Figure 2 shows an example of the range of GDx RNFL thickness measurements in a group of 134 healthy eyes.23 In glaucoma eyes, the reported average RNFL thickness across several studies ranges from approximately 61-77 µm. Some of this variation is likely due to the difference of glaucoma severity (and other factors) across studies, however, it is clear that many glaucoma eyes fall within the normal range. Figure 3 shows the distribution of GDx RNFL thickness measurements in 75 healthy and 65 glaucoma eyes, illustrating the substantial overlap in RNFL thickness between subjects and patients. This overlap hampers the ability to effectively discriminate between healthy and glaucomatous eyes based on this parameter alone. Because of this, other parameters that describe the relationships between superior and inferior thickness and temporal and nasal thickness (ratio and modulation parameters) are available. Detecting glaucoma with GDx/NFA II
Fig. 3. GDx measured average RNFL thickness in 65 glaucomatous eyes (eyes with repeatable abnormal standard automated perimetry results) and 75 healthy eyes. Diamonds represent 95% confidence intervals. The mean value for glaucomatous eyes is 61.8 µm (standard deviation, 13.88 µm) and the mean value for healthy eyes is 68.1 µm (standard deviation, 12.6 µm). There is a statistically significant difference in RNFL thickness between groups, despite the large overlap in measurements (t (1,138) = -2.83; p = 0.005).
ceptibility to the effects of intraocular pressure), decreased ocular perfusion, or a combination of these factors. This decrease in the retinal ganglion cell population in glaucoma results in a thinner than normal RNFL in glaucoma eyes compared to healthy ones. However, within the healthy population, optic never fiber count varies by a factor of more than two (reported range: approximately 700,000-1.5 million fibers per eye).18-21 Because optic nerve fiber count is correlated with RNFL thickness, RNFL thickness in healthy eyes also varies dramatically. For instance, healthy global RNFL thickness measured using GDx ranges from
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Using the standard clinical GDx printout, the clinician may gain some insight into the health of the RNFL and, therefore, in combination with other sources of information, the probability that the patient has glaucomatous damage. The RNFL thickness map (or retardation map) provides information regarding the thickness of the RNFL across the available field of view. In healthy eyes, the RNFL is thicker in the inferior and superior quadrants than in the nasal and temporal ones, resulting in a sinusoidal ‘double hump’ thickness profile around the optic disc. Furthermore, the thickness in the inferior and superior quadrants is approximately equal and uniform (except in the cases of healthy eyes with split nerve fiber bundles).24 Therefore, any dramatic deviation from this pattern may be cause for concern. The RNFL thickness graphs provide similar thickness modulation information. Information comparing quadrant specific thickness values relative to a normative database is also informative, assuming the normative database includes accurate measurements from subjects with a wide variety of racial backgrounds, spanning all ages. Figure 4 shows a GDx printout from a glaucomatous eye. In this case, the patient is a 90year-old female glaucoma patient. The standard automated perimetry result closest to the GDx imaging date shows a superior arcuate defect extending from the blind spot to the nasal step region, and a less deep inferior arcuate defect. The Glaucoma Hemifield Test result is outside normal limits, and mean deviation is -10.5 dB (p < 0.5%). The GDx retardation map suggests decreased RNFL thickness overall, with a decrease in thickness amplitude around the disc. This information is mirrored by the RNFL thickness graph. The quadrant-spe-
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Fig. 4. GDx extended analysis printout from a glaucomatous eye. See text for description.
cific deviation from normal values information suggests RNFL thinning in the superior and inferior quadrants. Several NFA parameters are outside normal limits, particularly those that describe the amplitude of the RNFL thickness profile (ratio and modulation parameters). The GDx ‘Number’ is within the range indicative of glaucoma. Figure 5 shows an alternate GDx printout that compares the right and left eyes of the above patient in order to display possible asymmetries in RNFL thickness measurements and their deviations from normal. This type of analysis may be valuable because glaucoma often initially presents in one eye only, or is asymmetric in severity. Sensitivity for detecting glaucoma with GDx/ NFA II Despite the large and overlapping range of RNFL thickness measurements in both healthy and glaucoma eyes, some studies have reported relatively good sensitivities for detecting eyes with glaucomatous damage, using information available to the
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clinician. Choplin and Lundy reported that experienced observers provided with masked GDx printouts were able to successfully identify 83% of 42 known glaucoma patients.25 In a similar study, also using clinical printouts, Sanchez-Galeana et al. reported that three experienced observers identified between 72% and 82% of 39 known glaucoma patients.26 Other studies have reported better sensitivities for discriminating glaucoma eyes from healthy ones using individual parameters or combinations of parameters in statistical models.10,12,27,28 However, in some cases, reported sensitivities have been worse.7,13 Differences in reported sensitivities across studies are probably affected by study methodology, glaucoma severity in the different study populations, and possibly by image quality affected by calibration of the instrument and operator skill. Moreover, sensitivity for detecting glaucoma cannot always be compared across studies because reported sensitivities may be dependent upon different specificity cut-offs for individual continuous variable GDx parameters or combinations thereof.
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Fig. 5. GDx symmetry analysis printout from the same eye shown in Figure 4.
Other sources of birefringence in the human eye For laser light from the GDx to measure RNFL thickness, it must first pass through two structures other than the RNFL that might provide birefringent information that could affect measurements: the cornea and the lens. To address the possible artifactual contribution of these structures to RNFL thickness measurements, the GDx incorporates a corneal polarization compensator that assumes an axis of corneal polarization of 15° nasally downward. Recently, Greenfield et al. determined that there is a large variation in the corneal polarization axis in normal eyes.29 Because of this, it is likely that polarization artifacts from the cornea influence GDx RNFL measurements in a substantial number
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of eyes in which corneal polarization compensation is not complete, thus making the measurements inaccurate. Evidence suggests that birefringence of the lens has little or no effect on GDx measurements.30,31 Addressing corneal polarization effects A recent prototype instrument designed for research incorporates a variable corneal compensator with the GDx to address the effect of the considerable range of corneal polarization axes within the population. Using this device, RNFL thickness is measured at the macula where no ‘double-hump’ thickness profile is expected. The variable compensator
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Fig. 6. GDx serial analysis printout depicting progressive RNFL thinning in a glaucomatous eye over the course of 18 months. See text for description.
is adjusted until a flat thickness profile is obtained, and the resulting corrected polarization axis value is used in measuring RNFL thickness in the peripapillary region. Zhou and Weinreb found that individualized anterior segment compensation could be achieved so that the measured birefringence largely reflects the RNFL birefringence.32 Recent studies using this device have shown an improvement in discrimination between healthy and glaucomatous eyes, using variable compensation compared to fixed axis compensation for some but not all parameters.33,34 These results suggest that using patient-specific polarization axis compensation will likely improve the glaucoma diagnosis ability of the GDx. Following glaucoma with GDx/NFA II A recent study by Greenfield and Knighton showed that the axis of corneal polarization in healthy eyes is probably stable over time. This information suggests that, even if baseline measurements are inaccurate because of corneal contributions to RNFL thickness measurements, changes in RNFL
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measurements over time are not likely to be effected by changes in corneal polarization axis and that, although an inaccurate baseline measurement may exist, changes in RNFL thickness over time in glaucoma eyes can probably be attributed to changes in the disease state. Therefore, it appears that SLP technology, even with fixed corneal polarization axis compensation, may be promising for glaucoma monitoring and follow-up, although this claim has not yet been sufficiently tested. For the monitoring of glaucomatous progression, a serial examination report printout is available to the clinician. This printout includes color-coded deviation from baseline thickness values for each follow-up examination. Significant thickness changes in serial images are determined by comparison to the variability in the three scans that make up the composite baseline image. This printout also includes retardation images for each examination, quadrant-specific deviation from normal values, select parameter values, and graphs depicting RNFL thickness within the measurement ellipse for baseline and follow-up images. Figure 6 shows a serial analysis printout depicting progressive RNFL thinning over the course of 18 months.
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Scanning laser polarimetry of the retinal nerve fiber layer At the time of the baseline image, this female patient was 68 years old. The standard automated perimetry result closest to the imaging date shows a superior arcuate defect extending from the blind spot to the nasal step region. The Glaucoma Hemifield Test result is outside normal limits, and the mean deviation is -7.2 dB (p < 0.5%). The serial retardation images show thinning (decreased brightness) in the superior, nasal, and inferior quadrants compared to baseline. This information is reflected in the deviation from normal figures, deviation from reference figures (shown as significant superior temporal thinning in the first follow-up image, and significant superior, nasal, and inferior nasal thinning in the second follow-up image), parameter values, and RNFL thickness graphs. It is possible that eyes which have undergone corneal laser refractive procedures, such as laser assisted in situ keratomelusis (LASIK), will provide a challenge to accurate follow-up using GDx, because of the potential change in the polarization axis of the lens as a result of these procedure. In these cases, it may be necessary to change the baseline image for serial analysis to a post-treatment examination date. Some studies have shown a significant decrease in SLP measured RNFL thickness after LASIK.36-38 However, one study showed no effect of excimer laser photorefractive keratectomy (PRK) on GDx parameters.39 In none of these studies was the effect of the procedures on the axis or magnitude of corneal polarization investigated directly. The issue of possible corneal polarization changes following laser refractive procedures might be avoidable when using a variable corneal compensation device. In this case, changes in corneal polarization axis and magnitude could be controlled for prior to image acquisition. Population-based glaucoma screening using GDx/NFA II The majority of GDx-related studies are conducted in glaucoma clinics, which are often referral practices that may not accurately portray the demographic and disease characteristics of the general population. Moreover, referral patients may be more cooperative. For these reasons, a new glaucoma screening technique should be tested in a real-world environment. Two population-based studies have addressed the success of GDx as a screening tool. Vitale et al. reported a sensitivity and specificity of 69% and 82%, respectively, using the GDx/NFA II neural network ‘Number’ in a population of 280 healthy eyes and in 98 eyes with definite or probable glaucoma in the Baltimore Eye Survey FollowUp study.40 Using the same parameter (GDx ‘Number’), Yamada et al.41 reported an optimal sensitivity and specificity (determined from receiver operating characteristic (ROC) curves) of 68% and 90%, respectively, in 122 healthy and 22
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glaucomatous eyes of attendees at a two-day public glaucoma screening program. These results are similar to those obtained in clinical settings. Conclusions SLP is a promising technique for assessing RNFL thickness in vivo by measuring polarization changes in laser light caused by the birefringent structure of the RNFL. Measurements using the GDx are reproducible, and differences in RNFL thickness between healthy and glaucomatous eyes are well substantiated. The GDx provides informative summary information which allows experienced clinicians to discriminate between healthy and glaucomatous eyes with reasonable accuracy. Furthermore, serial analysis information may allow the clinician to detect changes in nerve fiber layer thickness over time in diseased eyes. However, some GDx measurements can be inaccurate due to inadequate compensation of corneal polarization in some eyes. Moreover, some corneal laser refractive procedures may make comparisons between pre- and post-procedure measurements uninformative. Individually compensating for corneal birefringence will likely enhance the performance of this technology. References 1. Weinreb RN, Dreher AW, Coleman A, Quigley H, Shaw B, Reiter K: Histopathologic validation of Fourierellipsometry measurements of retinal nerve fiber layer thickness. Arch Ophthalmol 108:557-560, 1990 2. Dreher AW, Reiter K: Scanning laser polarimetry of the retina nerve fiber layer. SPIE Proceedings 1746:34-38, 1992 3. Knighton RW, Huang X, Zhou Q: Microtubule contribution to the reflectance of the retinal nerve fiber layer. Invest Ophthalmol Vis Sci 39:189-193, 1998 4. Knighton RW, Jacobson SG, Kemp CM: The spectral reflectance of the nerve fiber layer of the macaque retina. Invest Ophthalmol Vis Sci 30:2392-2402, 1989 5. Knighton RW, Huang XR: Directional and spectral reflectance of the rat retinal nerve fiber layer. Invest Ophthalmol Vis Sci 40:639-647, 1999 6. Morgan JE, Waldock A, Jeffery G, Cowey A: Retinal nerve fibre layer polarimetry: histological and clinical comparison. Br J Ophthalmol 82:684-690, 1998 7. Bowd C, Zangwill LM, Berry CC et al: Detecting early glaucoma by assessment of retinal nerve fiber layer thickness and visual function. Invest Ophthalmol Vis Sci 42:19932003, 2001 8. Choplin NT, Lundy DC, Dreher AW: Differentiating patients with glaucoma from glaucoma suspects and normal subjects by nerve fiber layer assessment with scanning laser polarimetry. Ophthalmology 105:2068-2076, 1998 9. Hoh ST, Greenfield DS, Mistlberger A, Liebmann JM, Ishikawa H, Ritch R: Optical coherence tomography and scanning laser polarimetry in normal, ocular hypertensive, and glaucomatous eyes. Am J Ophthalmol 129:129-135, 2000 10. Horn FK, Jonas JB, Martus P, Mardin CY, Budde WM: Polarimetric measurement of retinal nerve fiber layer thick-
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ness in glaucoma diagnosis. J Glaucoma 8:353-362, 1999 11. Lauande-Pimentel R, Carvalho RA, Oliveira HC, Goncalves DC, Silva LM, Costa VP: Discrimination between normal and glaucomatous eyes with visual field and scanning laser polarimetry measurements. Br J Ophthalmol 85:586591, 2001 12. Weinreb RN, Zangwill L, Berry CC, Bathija R, Sample PA: Detection of glaucoma with scanning laser polarimetry. Arch Ophthalmol 116:1583-1589, 1998 13. Zangwill LM, Bowd C, Berry CC et al: Discriminating between normal and glaucomatous eyes using the Heidelberg Retina Tomograph, GDx Nerve Fiber Analyzer, and Optical Coherence Tomograph. Arch Ophthalmol 119:985993, 2001 14. Hoh ST, Ishikawa H, Greenfield DS, Liebmann JM, Chew SJ, Ritch R: Peripapillary nerve fiber layer thickness measurement reproducibility using scanning laser polarimetry. J Glaucoma 7:12-15, 1998 15. Kook MS, Sung K, Park RH, Kim KR, Kim ST, Kang W: Reproducibility of scanning laser polarimetry (GDx) of peripapillary retinal nerve fiber layer thickness in normal subjects. Graefe’s Arch Clin Exp Ophthalmol 239:118121, 2001 16. Waldock A, Potts MJ, Sparrow JM, Karwatowski WS: Clinical evaluation of scanning laser polarimetry: I. Intraoperator reproducibility and design of a blood vessel removal algorithm. Br J Ophthalmol 82:252-259, 1998 17. Colen TP, Tjon-Fo-Sang MJ, Mulder PG, Lemij HG: Reproducibility of measurements with the nerve fiber analyzer (NfA/GDx). J Glaucoma 9:363-370, 2000 18. Jonas JB, Muller-Bergh JA, Schlotzer-Schrehardt UM, Naumann GO: Histomorphometry of the human optic nerve. Invest Ophthalmol Vis Sci 31:736-744, 1990 19. Jonas JB, Schmidt AM, Muller-Bergh JA, SchlotzerSchrehardt UM, Naumann GO: Human optic nerve fiber count and optic disc size. Invest Ophthalmol Vis Sci 33:20122018, 1992 20. Jonas JB, Schmidt AM, Muller-Bergh JA, Naumann GO: Optic nerve fiber count and diameter of the retrobulbar optic nerve in normal and glaucomatous eyes. Graefe’s Arch Clin Exp Ophthalmol 233:421-424, 1995 21. Mikelberg FS, Drance SM, Schulzer M, Yidegiligne HM, Weis MM: The normal human optic nerve: axon count and axon diameter distribution. Ophthalmology 96:1325-1328, 1989 22. Poinoosawmy D, Tan JC, Bunce C, Membrey LW, Hitchings RA: Longitudinal nerve fibre layer thickness change in normal-pressure glaucoma. Graefe’s Arch Clin Exp Ophthalmol 238:965-969, 2000 23. Bowd C, Zangwill L, Blumenthal E et al: Imaging of the optic disc and retinal nerve fiber layer: the effects of age, optic disc area, refractive error, and gender. J Opt Soc Am A 19:197-207, 2001 24. Colen TP, Lemij HG: Prevalence of split nerve fiber layer bundles in healthy eyes imaged with scanning laser polarimetry. Ophthalmology 108:151-156, 2001 25. Choplin NT, Lundy DC: The sensitivity and specificity of scanning laser polarimetry in the detection of glaucoma in a clinical setting. Ophthalmology 108:899-904, 2001
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26. Sanchez-Galeana C, Bowd C, Blumenthal EZ, Gokhale PA, Zangwill LM, Weinreb RN: Using optical imaging summary data to detect glaucoma. Ophthalmology 108:18121818, 2001 27. Trible JR, Schultz RO, Robinson JC, Rothe TL: Accuracy of scanning laser polarimetry in the diagnosis of glaucoma. Arch Ophthalmol 117:1298-1304, 1999 28. Essock EA, Sinai MJ, Fechtner RD, Srinivasan N, Bryant FD: Fourier analysis of nerve fiber layer measurements from scanning laser polarimetry in glaucoma: emphasizing shape characteristics of the ‘double-hump’ pattern. J Glaucoma 9:444-452, 2000 29. Greenfield DS, Knighton RW, Huang XR: Effect of corneal polarization axis on assessment of retinal nerve fiber layer thickness by scanning laser polarimetry. Am J Ophthalmol 129:715-722, 2000 30. Collur S, Carroll AM, Cameron BD: Human lens effect on in vivo scanning laser polarimetric measurements of retinal nerve fiber layer thickness. Ophthalmic Surg Lasers 31:126130, 2000 31. Park RJ, Chen PP, Karyampudi P, Mills RP, Harrison DA, Kim J: Effects of cataract extraction with intraocular lens placement on scanning laser polarimetry of the peripapillary nerve fiber layer. Am J Ophthalmol 132:507-511, 2001 32. Zhou Q, Weinreb RN: Individualized compensation of anterior segment birefringence during scanning laser polarimetry. Invest Ophthalmol Vis Sci 43:2221-2228, 2002 33. Weinreb RN, Bowd C, Greenfield DS, Zangwill LM: Measurement of the magnitude and axis of corneal polarization with scanning laser polarimetry. Arch Ophthalmol 120:901906, 2002 34. Weinreb RN, Bowd C, Zangwill L: Glaucoma detection using scanning laser polarimetry with variable corneal polarization compensation. Arch Ophthalmol. In press, 2002 35. Greenfield DS, Knighton RW: Stability of corneal polarization axis measurements for scanning laser polarimetry. Ophthalmology 108:1065-1069, 2001 36. Gurses-Ozden R, Pons ME, Barbieri C et al: Scanning laser polarimetry measurements after laser-assisted in situ keratomileusis. Am J Ophthalmol 129:461-464, 2000 37. Gurses-Ozden R, Liebmann JM, Schuffner D, Buxton DF, Soloway BD, Ritch R: Retinal nerve fiber layer thickness remains unchanged following laser- assisted in situ keratomileusis (1). Am J Ophthalmol 132:512-516, 2001 38. Tsai YY, Lin JM: Effect of laser-assisted in situ keratomileusis on the retinal nerve fiber layer. Retina 20:342345, 2000 39. Choplin NT, Schallhorn SC: The effect of excimer laser photorefractive keratectomy for myopia on nerve fiber layer thickness measurements as determined by scanning laser polarimetry. Ophthalmology 106:1019-1023, 1999 40. Vitale S, Smith TD, Quigley T et al: Screening performance of functional and structural measurements of neural damage in open-angle glaucoma: a case-control study from the Baltimore Eye Survey. J Glaucoma 9:346-356, 2000 41. Yamada N, Tomita G, Yamamoto T, Kitazawa Y: Changes in the nerve fiber layer thickness following a reduction of intraocular pressure after trabeculectomy. J Glaucoma 9:371375, 2000
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The glaucomatous optic nerve staging system with confocal tomography
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The glaucomatous optic nerve staging system with confocal tomography Roberto Sampaolesi1 and Juan Roberto Sampaolesi2 Department of Ophthalmology, Faculty of Medicine, University of Buenos Aires; 2Department of Ophthalmology, Faculty of Medicine, ‘UCES’ (University of Business and Social Sciences); Buenos Aires, Argentina 1
Keywords: confocal tomography, optic nerve, glaucoma, Heidelberg tomograph
History of optic nerve examination Since the first images of the fundus were obtained with an ophthalmoscope, created by Helmholtz1 145 years ago, examination methods have continued to improve. First, examination with direct ophthalmoscopy was used, followed by examination with binocular indirect ophthalmoscopy, optic disc drawings, optic disc stereophotographs, planimetry, Takamoto and Schwartz’s stereophotogrammetry, neuroretinal rim measurements, Airaksinen and Tuulonen’s evaluation of the retinal nerve fiber layer, optic disc pallor measurements, angiofluoresceinography of the optic disc, image analyzer (Rodenstock Optic Nerve Head Analyzer, Topcon Analyzer), laser scanning ophthalmoscopy, and Lotmar and Goldmann’s stereochronoscopy. All these methods are extensively explained in the worthwhile book The Optic Nerve in Glaucoma, by Rohit Varma and George Spaeth.2 Of these methods, we still use retinofluoresceinography, because of its particular usefulness. We no longer use stereoscopic photographs of the optic nerve in adults, due to the significant interobserver variation in interpretation of the results, as reported in the literature.3-5 But stereoscopic photographs are useful for optic nerve examinations in children under two years of age because, due to their flat corneas, with confocal tomography we have so far failed to obtain good images. In Boston, we learned about Schwartz and Takamoto’s stereophotogrammetry,6 a very reliable method, the results of which are consistent with those obtained with the Heidelberg Retina Tomograph (HRT). However, it is a time-consuming method. We later performed neuroretinal rim measurements using Airaksinen et al.’s method,7 which we could put into practice
thanks to Dannheim and Airaksinen’s personal communications. This turned out to be the most useful method. We also tried Lotmar and Goldmann’s optic disc stereochronoscopy.8 All these methods require pupil dilation. For the above-mentioned methods, and particularly for measuring optic disc parameters (area, cup area, neuroretinal rim), the formula introduced by Littman in 1982 for the first time allowed us to obtain the dimensions (length, surface) of any observable object in the fundus (exudates, tumors, foreign bodies, optic discs, vessels, etc.). Images of these elements can be observed with considerable magnification produced by the ocular system. This morphometric magnification was corrected in order to find real values with the Littman formula.9 Littman was an engineer who worked for Zeiss and whom we had the chance of meeting in Buenos Aires. Since 1982, thanks to his formula, we have been able to obtain real measurements, in mm or mm², of a body or structure on the retina. For this formula, corneal curvature, which is measured with an ophthalmometer, axial length, which is measured by echometry, and refraction, are very important. Corneal thickness, its posterior curvature, lens face curvatures, depth of the anterior chamber, and lens thickness, are not required. This is because, even if they varied, their influence on the measurement would be minimal. This formula does not apply to aphakia, pseudophakia, and refraction changes due to opacity of the lens. Finally, in 1990, the HRT was introduced by Heidelberg Engineering (Burk et al.10, etc.), and it was with this device that we started our extensive research in 1991, ten years ago, with more than 7300 patients having been examined to date. In this chapter, we will deal with what we be-
Address for correspondence: Professor Roberto Sampaolesi, Parana 1239 – 1A, 1018 Buenos Aires, Argentina. e-mail:
[email protected] Lasers in Ophthalmology – Basic, Diagnostic and Surgical Aspects, pp. 285–301 edited by F. Fankhauser and S. Kwasniewska © 2003 Kugler Publications, The Hague, The Netherlands
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lieve is a very important topic, the HRT parameters used for optic nerve staging in glaucoma, as well as the follow-up of optic nerve damage.11 The tomographic classification is mainly based on the volumes of the said structures, and only secondarily on surfaces and other parameters. This is due to the possibility of stereometric and tridimensional analyses. The advantage provided by volume measurements over area measurements, is that the former are raised to the third power, while the latter are only raised to the second power (whenever change occurs, no matter how slight, there is a greater variation if the value is raised to the cube than if it is raised to the square). Parameters used for glaucoma staging As already stated, the main parameter used for the classification was volume, followed by area, thickness, and slope. The volumes taken into consideration are neuroretinal rim volume, and cup volume. The most important is, in fact, neuroretinal rim volume, but cup volume is also taken into account since a decrease in neuroretinal rim volume produces an increase of cup volume; this is a causeeffect relationship. The same occurs with area, the cup area (red) increases as the rim area (green and blue) decreases. The most important parameters used for the classification are listed in Table 1 and are illustrated in Figure 1, where the parameters belonging to the optic disc are separated from those that are analyzed in the contour line graph. These latter parameters are: mean retinal nerve fiber layer (RNFL) thickness, height variation of the contour line, and area enclosed by the contour line and the reference plane in the contour line height variation diagram (RNFL cross-sectional area). As we all know, the
Table 1. Parameters used for classification Rim volume Cup volume Rim area Cup area Cup shape measure Mean RNFL thickness Height variation of contour line Area between curve and plane
reference plane, far from being just another parameter, is the limit on which most parameters strictly depend, and with which they have a close relationship. Due to this, its position must always be verified and, when performing a longitudinal study, it must be checked to see that it always remains at the same level. Concept and limits of normality The concept of normality is based on the fact that all the optic disc parameters are normal. Nevertheless, in clinical practice, sometimes, for one reason or another, the fact that one or two parameters are not within the limits of normality or normal range does not indicate pathology. The concept and limit of normality are outlined in Figure 2. The limits of normality were obtained in a study of 108 normal volunteers.12 Table 2 lists the most important limits, for example, for neuroretinal rim volume, the normal inferior limit (3.20 mm3) and not the normal superior limit, is given, since this is mainly used to differentiate a large neuroretinal rim from an optic disc edema. In some patients, a neuroretinal rim smaller than 320 µm3 may be found during the first tomography, which makes it fall within the classification of borderline. Nevertheless, it must be taken into account
RETINAL SURFACE
Fig. 1.
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that the evolution of these patients sometimes does not involve optic nerve damage, but remains stable for years, which indicates a physiological or normal decrease of the neuroretinal rim in this group of patients. Evolution phases Fig. 2. Table 2. Limits of normality Rim volume Cup volume Rim area Cup area Cup shape measure Mean RNFL thickness Height variation of contour line Area between curve and plane
min. 3.20 mm3 max. 0.12 mm3 min. 1.37 mm² max. 0.60 mm² max. -0.15 min. 0.17 mm min. 0.27 mm min. 0.87 mm²
Table 3. Normal optic discs Rim volume Cup volume Rim area Cup area Cup shape measure Mean RNFL thickness Height variation of contour line Area between curve and plane Normal visual field
> < > < < > > >
0.32 mm3 0.12 mm3 1.37 mm² 0.60 mm² -0.15 0.17 mm 0.27 mm 0.87 mm²
Glaucomatous optic disc evolution was classified into the following groups: normal borderline phase 1 phase 2 phase 3 phase 4 With the exception of borderline optic discs, the remaining phases are separated from one another by more than two standard deviations, rendering separation into the various groups more significant. Only parameters meeting this requirement are mentioned, since most of those remaining have no significant differences between the various evolution phases. Normal optic discs (Table 3) are characterized by a just visible Elschnig’s ring, except in the temporal area. Both poles have important fiber bundles, which correlate with the two camel humps displayed by the contour line diagram (Fig. 3).
Fig. 3.
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Table 4. Borderline optic discs Rim volume Cup volume Rim area Cup area Cup shape measure Mean RNFL thickness Height variation of contour line Area between curve and plane
Table 5. Phase 1 > < > < < > > >
0.32 mm3 0.12 mm3 1.37 mm² 0.60 mm² -0.15 0.17 mm 0.27 mm 0.87 mm²
Normal visual field
In borderline optic discs (Table 4), neuroretinal rim volume is normal on measuring the entire disc, but if we analyze the rim volume by octants and quadrants, there is a decreased value in one of these sectors. The decrease of neuroretinal rim volume in this phase does not affect the whole optic disc. It corresponds with Burk’s pseudonormal optic discs, in which the humps remain unchanged and there is slight neuroretinal rim loss. With the exception of cup increase, no parameters are altered (this is seen less frequently) (Fig. 4). Phase 1 optic discs (Table 5) are characterized by a generalized decrease in retinal thickness that can be seen in the contour line diagram as an approximation between the contour line proper and the reference plane. At the same time, a decrease in the height of the camel humps, which correlates with the loss of fiber bundles (see tridimensional presentation), and with the fact that Elschnig’s ring is more visible than before, can be seen (Fig. 5). Phase 2 optic discs (Table 6) already have a loss
Rim Cup Rim Cup Cup
volume volume area area shape measure
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mm3 mm3 mm² mm²
Normal visual field
of up to 50% of their total retinal nerve fibers. The disappearance of both humps, which correlates with a huge cup increase that invades the superior and inferior poles, can be seen. The mean RNFL thickness, preventing the contour line from approaching the reference plane, remains unchanged (Fig. 6). Phase 3 optic discs (Table 7) are characterized by a huge decrease in mean RNFL thickness, which causes the contour line to approach towards the reference plane (when localized defects occur, the contour line reaches the reference plane in the damaged areas). The summation image allows the bottom of the cup and Elschnig’s ring clearly to be seen in their full extent. The cup surface covers Table 6. Phase 2 Rim Cup Rim Cup Cup
volume volume area area shape measure
Beginnings of visual field defects
Fig. 4.
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0.30-0.20 mm3 0.24-0.48 mm3 1.20-0.80 mm² 1.00-1.50 mm² -0.12- -0.12
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Fig. 5.
Fig. 6.
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Table 7. Phase 3 Rim Cup Rim Cup Cup
Table 8. Phase 4 0.20-0.10 mm3 0.48-0.96 mm3 0.80-0.40 mm² 1.50-1.80 mm² -0.70- -0.20
volume volume area area shape measure
Rim Cup Rim Cup Cup
volume volume area area shape measure
Visual field defects
Visual field in stage 3 (terminal)
almost the entire optic disc surface. The neuroretinal rim persists like a thin halo around it (Fig. 7). Phase 4 optic discs (Table 8) are characterized by a final decrease in retinal thickness, where the contour line is parallel to the reference plane and, in the places where there is no neuroretinal rim left, the contour line height variation diagram is below the reference plane. This fact correlates with the appearance of white areas in analysis of the surfaces and with the presence of absolute visual field defects (Fig. 8). All six phases are summarized in Figure 9. In normal optic discs, as well as in borderline discs, Elschnig’s ring can only be seen in the temporal sector, whereas in all other phases, it can be seen almost to its full extent, due to fiber atrophy. The bottom of the cup can be more clearly seen from phase 2 onwards. When the brightness of the retina is observed in each section, it can be seen that this decreases steadily from normality to phase 4. Cup shape measures change rapidly. In phase 2, the cup slope is almost perpendicular, while in phases 3
and 4, bayonet-shaped vessels are revealed. The small vessels become more and more evident and their contours are more visible as they become more definite (this is due to atrophy of the retinal nerve fibers). Nevertheless, at first sight, the condition of the optic disc in phase 4 may seem better than in phase 3. Also, the time elapsing between normal and borderline optic discs, or between borderline and phase 1 optic discs may seem the same. This is easily solved with stereometric analysis of the surfaces. Figure 10 shows the six phases together in the ‘measure’ menu, with the color-coded analysis of the surfaces. In normal optic discs, the cup is surrounded by a large neuroretinal rim and not centered in the optic disc. This occurs in normal conditions due to the huge infiltration of fibers at the superior and inferior poles. In borderline optic discs, it is possible to see how the surface of the cup increases at the expense of decrease of the rim area. Simultaneously, the cup becomes central and its area invades the tilted neuroretinal rim area, thus reducing its separation from the flat neuroretinal
Fig. 7.
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Fig. 8
Fig. 9.
rim. In phase 1 optic discs, the cup continues to increase and gets closer to the flat neuroretinal rim, leaving a thin separation covered by the tilted neuroretinal rim. The total surface of the neuroretinal rim decreases markedly. In Phase 2 optic discs, the cup increases considerably and starts to become
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slightly eccentric, and the tilted rim disappears completely in these regions. Consequently, the cup surface borders the flat rim surface. This fact can sometimes cause localized defects and, together with the diffuse atrophy of the rest of the retina, it correlates with the onset of the visual field defects
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Fig. 10.
in this phase. In Phase 3 optic discs, the cup surface almost covers the complete optic disc region. The tilted rim has almost completely disappeared. Only a thin flat rim margin separates the cup from the external optic disc margin. This small volume of neuroretinal rim keeps the visual function at the same level; this is correlated with the rapid visual field loss produced when the remaining neuroretinal rim is damaged. In phase 4 optic discs, the cup occupies almost all the optic disc surface and, in some sectors, where the neuroretinal rim has been completely destroyed, the cup touches the external optic disc margin, making the total absence of the neuroretinal rim evident in that sector. White regions can occur in phase 4, which are due to the fact that the retinal surface is below the level of the reference plane in the most badly damaged sectors. These lesions produce absolute optic disc defects that have a bad prognosis. The role of confocal tomography (Heidelberg Retina Tomograph) Classification of clinical cases Case 1 A 64-year-old male (left eye). Best-corrected visual acuity: sph: -1; cyl: -0.25 in 0°. Diagnosis: openangle glaucoma (eight years earlier). The diurnal pressure curves performed during those eight years revealed an intraocular pressure (IOP) of 33 mmHg at the 7 a.m. measurement, while at 9 a.m., and at
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1, 3, 6, 9 and 12 p.m., all the readings were below 19 mmHg. Medical therapy succeeded in lowering the morning peak to 24-28 mmHg. The visual field was normal according to computerized perimetry, and HRT revealed a phase 1 optic nerve head (ONH) (Figs. 11 and 12). Case 2 A 72-year-old male. Diagnosis: chronic narrowangle glaucoma. Under hypotensive therapy for 15 years. IOP without medication: 31 mmHg; and with medical therapy: 24 mmHg. Peripheral iridectomy with a YAG laser was performed, later supplemented by argon-laser trabeculoplasty, which regulated the IOP to: mean: 17 mmHg; variability: 1.5, according to the diurnal pressure curve. HRT revealed a phase 2 ONH (left eye) associated with a borderline visual field (Figs. 13 and 14). Case 3 A 57-year-old female (left eye). Diagnosis: openangle glaucoma and myopia. Visual acuity: 10/20 with sph: -5.00. The diurnal pressure curve showed a mean of 23 mmHg and a variability of 1.6. Computerized perimetry revealed a stage 3 visual field, with an MD of 18.9 and a CLV of 96.4. HRT revealed a left phase 3 ONH (Figs. 15 and 16). Case 4 A 65-year-old female. Diagnosis: open-angle glaucoma (20 years earlier). Trabeculoplasty performed five years earlier. IOP monitoring was performed by another ophthalmologist by means of single-spot
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The glaucomatous optic nerve staging system with confocal tomography
Fig. 11.
Fig. 12.
Fig. 13.
Fig. 14.
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Fig. 15.
Fig. 16.
checks and not by diurnal pressure curves, with readings of 25 mmHg being taken. Current bestcorrected visual acuity: 18/20 (with sph: 0.75). Computerized perimetry revealed a stage 3 visual field, with an MD of 23.2, and a CLV of 60.5. HRT showed a phase 4 ONH (Figs. 17 and 18). Follow-up of optic nerve damage One of the more interesting applications of HRT is the follow-up of optic nerve damage. There are three periods of evolution in glaucoma: the hypertensive period, when there is neither optic nerve damage nor visual field defects, but with ocular hypertension as the only sign; the preperimetric period is characterized by the presence of either borderline, or phases 1 or 2 optic nerve damage, with the visual field still appearing normal. In the perimetric or final period, the optic nerve falls into either phase 2, 3, or 4, and the visual field is either borderline or belongs to stage 1, 2, or 3. We have designed our own medical records which include a chart on the first page, where the
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corresponding phase, stage, and, finally, period of each case can be ticked. Once the evolution stage of the patient has been determined, if duly recorded on the chart (one for each eye) with the dates of the examinations at the beginning of the patient’s record, the ophthalmologist is spared valuable time when the patient comes back for a follow-up, since he can identify the patient’s stage of evolution at a glance, and there is no need to check the entire records and confocal tomography and visual field examination printouts (Fig. 19). It should be stressed here that IOP should not be measured in single spot-checks, but rather by means of diurnal pressure curves, with the first measurement being made at 6 a.m. with the patient still in bed, using applanation tonometry. The following clinical histories are examples of this routine procedure, by which the evolution period of the disease is identified according to whether ocular hypertension has been controlled or not, thus enabling us to prescribe adequate medical therapy, change the current therapy, or suggest surgery.
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Fig. 17.
Fig. 18.
Fig. 19. The use of optic nerve head tomography in daily practice. The stages in glaucoma evolution. Glaucoma is a disease that can be divided into three periods: hypertensive, when there is high IOP; preperimetric, when optic nerve damage starts to develop; and, perimetric, when optic nerve damage is in phase 2 and visual field defects start.
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R. Sampaolesi and J.R. Sampaolesi Record # 6369, J.D., 56-year-old male
Fig. 20. Pseudo-low-tension glaucoma, with progression of optic nerve damage and apparently normal IOP. Based on the positive result of the diurnal pressure curve, the diagnosis of low-tension glaucoma was replaced by chronic narrow-angle glaucoma in the perimetric period. If the patient is refractory to medical therapy, surgery is indicated.
Case 1
Case 3
Record No. 6369, JD, a 56-year-old male The patient presented for consultation in 1997, when he was in the hypertensive evolution period of glaucoma with normal optic nerve and visual field. He was diagnosed with chronic narrow-angle glaucoma. In 1998, he underwent laser iridectomy for narrow-angle, which regulated his IOP. His follow-up consultations were carried out abroad, where he was diagnosed as having low-tension glaucoma. However, when he returned home in 1999, HRT was performed and visual field examinations and the diurnal pressure curve revealed severely pathological IOP values at 6 a.m. (Figs. 20 and 21).
Record No. 4457, DOJC, a 29-year-old male At his first consultation in 1991, the patient was diagnosed with pigmentary glaucoma in the hypertensive period (with normal visual field and optic nerve). However, IOP readings were elevated, with single spot-checks reaching as high as 24-26 mmHg, resulting in pathological diurnal pressure curves. Medical therapy was initiated and he is still being followed up, in particular by IOP monitoring by several diurnal pressure curves (Figs. 24 and 25). Some of the considerations to be taken into account with regard to diagnosis and follow-up are: In the vast majority of cases, there is consistency between the time elapsed with untreated ocular hypertension, the evolution phase of optic nerve damage, and the evolution stage of the visual field defect. Here, we will mainly deal with the correlation between the two latter parameters: optic nerve and visual field. It is during phase 2 in the evolution of optic nerve damage that the visual field typically starts to deteriorate. Whenever there is no correlation between the optic nerve and visual field condition, i.e., when the optic nerve is in either phase 3 or phase 4 while the visual field is normal or slightly abnormal, either of the following situations should be suspected: firstly, the presence of a scotoma that goes undetected by the visual field, in which case the number of stimuli examined is not enough. Another program with a greater number of stimuli – such as a macular program – should therefore be used, and, if a scotoma is detected, there is a correlation; otherwise, a disease other than glaucoma should be considered. Secondly, a normal optic nerve may be associ-
Case 2 Record No. 5001, LDC, a 68-year-old female At her first consultation in 1993, the patient’s right eye was in the perimetric evolution period, with the optic nerve in phase 2 and the visual field in stage 2. Visual acuity was 0.4 in OD and 0.8 in OS. The patient was diagnosed as having chronic narrowangle glaucoma. The first surgical procedure was performed in June 1993, and the IOP was successfully regulated according to the diurnal pressure curves. In 1998, the IOP became higher, with peaks of 26 mmHg detected in 2000 in the right eye, and the diurnal pressure curve was seriously pathological (mean: 22.86 and SD: 95), but normal in the fellow eye. Evolution of the condition of the optic nerve and visual field progressed because the patient failed to comply with her medical therapy for a year, when the IOP started to rise (Figs. 22 and 23).
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The glaucomatous optic nerve staging system with confocal tomography
Fig. 21. a. Evolution of optic nerve damage revealed by confocal tomography, which maps the neuroretinal rim.
Fig. 21. b. Evolution of the visual field.
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R. Sampaolesi and J.R. Sampaolesi thickness and neuroretinal rim. They also supplemented the examination of their cases with blueyellow visual fields, and concluded that the four above-mentioned parameters are the most important for differentiating normal cases from glaucomatous ones. Blue-yellow visual field examinations are only useful when they are corrected according to the autofluorescence (transmission properties) of the lens, which changes with age. The only difficulty in this regard is that the device for measuring autofluorescence is very expensive. Conclusions
Fig. 21. c. IOP values according to the diurnal pressure curve. During the day, and particularly at 9 a.m., noon, and 3 and 6 a.m., the IOP values are normal, while at 6 a.m., with the patient still in bed, they are 28 mmHg in OD and 26 in OS. Adequate medical therapy is prescribed, which, if refractory, will lead to surgery.
ated with an evident visual field defect. In this case, another disease still to be identified should be suspected. However, a few cases may be diagnosed as ‘vasospastic low-tension glaucoma’, since a visual field defect can be accompanied by a normal optic nerve. Finally, it should be borne in mind that the data supplied by the HRT are valuable for discs with an area within the normal range. However, they are not accurate in the case of small discs or megalopapillas. Vihanninjoki et al. have recently published a study on this,13 stating that the most important parameters for diagnosis and follow-up are, firstly, the cup-shape measure and, thereafter, mean RNFL
Optic nerve evaluation with confocal tomography permits the diagnosis of glaucomatous optic disc neuropathy in its early stages of evolution, when ophthalmoscopy is not yet able to detect any changes. Moreover, optic disc measurement is just as important as IOP and visual field measurement, since it is possible to stage the disease in each individual case using these three techniques. It is very difficult to differentiate between glaucoma in the hypertensive period and preperimetric glaucoma without the aid of confocal tomography. Furthermore, if it is difficult to stage a patient with optic disc biomicroscopy, it will be even more difficult to determine whether glaucomatous optic disc neuropathy has progressed in any way. In our opinion, optic nerve confocal tomography is a vital tool in the evaluation of glaucomatous patients, their accurate early diagnosis, staging, and follow-up, as well as for judging the efficacy of their medical and/or surgical treatment.
RECORD # 5001, L.D.C., B. 68-year-old female
Stage: advanced perimetric period Course of action: medical and/or surgical therapy
Fig. 22. Previously undetected chronic narrow-angle glaucoma undergoing filtering surgery which successfully regulated IOP for six years, with subsequent gradual and progressive destabilization due to lack of follow-up checkups and non-compliance with medical therapy. Glaucoma is in the advanced perimetric evolution period. Surgical therapy will be indicated if medical therapy is ineffective.
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Fig. 23. Pre- and postoperative optic nerve condition in 1993 and follow-up confocal tomography examinations performed in 1996 and 1999, showing the marked changes in the neuroretinal rim from the postoperative examination to the one performed in 1999. The visual field shows consistent clear variations in the MD, which increases from 8.9 to 9.9, as well as in the CLV, which changes from 70.1 to 51.1.
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R. Sampaolesi and J.R. Sampaolesi RECORD # 4457, D.O.J.C., 29-year-old female
Fig. 24. Pigmentary glaucoma in the preperimetric period with early detection of ocular hypertension, which was regulated by timolol 0.5% b.i.d. All measurements of the diurnal pressure curve are normal. The patient is still being followed up annually.
References 1. Helmholtz HV: Beschreibung eines Augen-Spiegels zur Untersuchung der Netzhaut im lebenden Auge. In: Engelking D (ed) Dokumente zur Erfindung des Augenspiegels durch Herrmann von Helmholtz im Jahr 1850. Munich: Bergmann Verlag 1950 2. Varma R, Spaeth G: The Optic Nerve in Glaucoma. Philadelphia, PA: JB Lippincott Co 1993 3. Leydhecker W, Krieglstein GK, Colloni EV: Observer variation in applanation tonometry and estimation of the cup disc ratio. In: Krieglstein GK, Leydhecker W (eds) Glaucoma Update: International Glaucoma Symposium, Nara, Japan, 1978, pp 101-117. Berlin/Heidelberg/New York: Springer Verlag 1979 4. Lichter PR: Variability of expert observers in evaluating the optic disc. Trans Am Ophthalmol Soc 74:532-572, 1976 5. Varma R, Steinmann WC, Scott IU: Expert agreement in evaluating the optic disc for glaucoma. Ophthalmology 99:215-221, 1992 6. Takamoto T, Schwartz B: Photogrammetric measurements of the optic disc in glaucoma. Int Arch Photogrammetry 23(B5):732, 1980 7. Airaksinen PJ, Drance SM, Douglas GR, Schulzer M: Neuroretinal rim areas and visual field indices in glaucoma. Am J Ophthalmol 99:107, 1985
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8. Goldmann H, Lotmar W: Rapid detection of changes in the optic disc: esterochronoscopy. Graefe’s Arch Clin Exp Ophthalmol 202:87-90, 1977 9. Littman H: Zur Bestimmung der wahren Grobe eines Objektes auf dem Hintergrund des lebenden Auges. Klin Mbl Augenheilk 180:286, 1982 10. Burk R, Konig J, Rohrschneider K, Noack H, Volcker HE, Zinser G: Analysis of three-dimensional optic disk topography by laser scanning tomography: parameter definition and evaluation of parameter inter-dependence. In: Nasemann J, Burk ROW (eds) Scanning Laser Ophthalmoscopy and Tomography, pp 161-176. Munich: Quintessenz 1990 11. Sampaolesi R, Sampaolesi JR: Confocal Tomography of the Retina and the Optic Nerve Head. Heidelberg: CityDruck 1999 12. Sampaolesi JR, Sampaolesi R: Lecture: Study of normality in the optic nerve head with HRT, presented at Curso y Simposio Argentino de Glaucoma, July 1995, Buenos Aires, Argentina 13. Vihanninjoki K, Teesalu P, Burk ROW, Läärä E, Tuulonen A, Airaksinen PJ: Search for an optimal combination of structural and functional parameters for the diagnosis of glaucoma: multivariate analysis of confocal scanning laser tomograph, blue-on-yellow visual field and retinal nerve fiber layer data. Graefe’s Arch Clin Exp Ophthalmol 238:477-481, 2000
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Fig. 25. The optic nerve has remained unchanged, as has the visual field, over eight years.
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Principles of photodisruption Joel M. Krauss New York, USA
Keywords: capsulotomy, cornea, Nd:YAG laser, Nd:YLF laser, photodisruption
Outline
Laser-tissue interactions Photodisruption Laser principles optical breakdown and plasma formation mechanisms of damage Instrumentation Clinical applications posterior capsulotomy iridectomy and other anterior segment applications posterior segment New laser techniques Conclusions and future developments Laser-tissue interactions The effect of laser radiation on a particular target depends on the properties of both the laser and the target. The most important laser output parameters are wavelength, duration, and power. Wavelength is a function of the laser cavity’s excited medium, which is a gas in argon, krypton, and excimer lasers, a liquid in dye lasers, a semiconductor in diode lasers, and a solid state material in the neodymium:yttrium aluminum garnet (Nd:YAG) and other lasers that will be examined in greater depth in this chapter. According to the principle of wave-particle duality, radiation is propagated in the form of both waves and discrete quanta, or photons. As such, radiation of a given wavelength is associated with photons of a corresponding energy, such that E = hν = hc/λ, where h = Planck’s constant, ν = frequency, c = speed of light, and λ = wavelength. Thus, frequency and energy increase as
wavelength decreases. The visible spectrum extends approximately from 380-760 nm. The first law of photochemistry (Grotthus-Draper) states that photons must be absorbed by a target in order to initiate a chemical reaction.1 A chromophore is a molecule, or a portion thereof, that absorbs a photon of a particular energy. Depending on the photon’s energy, a chromophore can undergo bond-breaking, ionization, or various types of molecular excitation. The ability of a target, which may or may not be of a homogeneous composition, to absorb radiation is measured by the attenuation in incident radiation after a certain length of the material has been traversed. The absorbence, A, of a material is defined as A(d) = log[I0/I(d)] = εcd, where I0 = initial intensity, I(d) = intensity at distance d, ε = absorptivity of the material, and c = molarity of the material. Transmission is that fraction of the incident energy that is not absorbed after traversing a particular target thickness. It is usually written in the form of Beer’s law, T(d) = 10-A(d) = e-αd, where αd = 2.3A defines the absorption coefficient α. α is generally given in units of cm-1 and represents the fraction of incident energy that is absorbed per unit length of target material. Absorption length is defined as α-1, or that distance at which e-1 = 0.368 of incident energy is transmitted, corresponding to 63.2% absorption. The thermal susceptibility of the irradiated tissue is denoted by the thermal relaxation time, τ, which gives an indication of the time required for the irradiated tissue to carry away heat energy from the target site. It is wavelengthdependent, and is proportional to 1/4α2κ, where κ measures the tissue diffusivity in cm2/sec. The absorption maximum of a compound is that wavelength in a given portion of the spectrum which has the highest probability of absorption. A plot of
Adapted from Krauss JM: Contemporary ophthalmic lasers. In: Rosen A, Rosen H (eds) New Frontiers in Medical Device Technology, pp 155-251. New York, NY: John Wiley & Sons 1995. Address for correspondence: J.M. Krauss, MD, 175 E 96th St., New York, NY 10128, USA. e-mail:
[email protected] Lasers in Ophthalmology – Basic, Diagnostic and Surgical Aspects, pp. 303–313 edited by F. Fankhauser and S. Kwasniewska © 2003 Kugler Publications, The Hague, The Netherlands
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absorption versus wavelength yields an absorption spectrum that is characteristic of the chemical composition of that compound. Quantum yield is a measure of the efficiency with which absorbed radiation produces chemical changes, while an action spectrum is a plot of the relative efficiency of the photoreaction versus wavelength. Photodisruption Photodisruption is the use of high peak-power ionizing laser pulses to disrupt tissue. Energy is concentrated in space and time to create optical breakdown, or ionization of the target medium, with formation of a plasma, seen as a spark. The use of optical radiation to produce a plasma became possible only after the development of lasers capable of emitting high power through very brief radiation pulses. Although the first lasers were too weak to achieve optical breakdown, in 1962 Hellwarth developed the method of Q-switching, which allowed the creation of very brief but large ruby laser pulses over 10-50 nanoseconds (nsec, 10-9 sec), with maximum powers in the tens of megawatts.2 In 1972, Krasnov reported the first use of clinically desirable intraocular photodisruption, treating the trabecular meshwork of eyes with open-angle glaucoma.3 To emphasize the relative importance of nonthermal acoustic mechanisms in creating these tissue effects, he used the term ‘cold laser’, which ignores the fact that plasma formation causes very localized temperature increases greater than 10,000EC. Further work demonstrated that, because of the ruby laser’s high-order mode structure (which limits the minimal spot size that can be achieved) it is not the ideal source for a clinically practical ophthalmic photodisruptor. Although by 1979 Gaasterland succeeded in building a Q-switched ruby laser capable of membranectomy, the required energy and waiting time between laser shots were too great to make this technique clinically useful. However, the increasing popularity of extracapsular cataract extraction (ECCE) and the pioneering research of Aron-Rosa et al.4 and Fankhauser et al.5 in the early 1980s with Nd:YAG lasers, soon combined to make laser photodisruption a reality. Laser principles Laser power can be increased by either increasing energy or, more practically, by decreasing the period over which the energy is delivered. The two principal means of compressing the laser output in time to achieve high-peak power are mode-locking and Q-switching. Mode-locking is comparable to the audible summation of musical tones with similar frequencies, known as beating, which is heard as a periodic surge in intensity. The phase relationships in lasers are synchronized by a shutter near one of the cavity mirrors. For ophthalmic applications, the most common shutter is a saturable dye, employed in a proc-
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ess known as passive mode-locking. The dye absorbs low-power radiation pulses, but becomes transparent on exposure to high-power ones. The Q-switch is an intracavity shutter which requires an active medium that allows atoms to remain in the high-energy state for a relatively long time to create high-peak power. Solid state media such as Nd:YAG are particularly well-suited for this process. At the appropriate time, the Q-switch shutter is opened, exposing the mirror. Oscillation and stimulated emission follow quickly, with emission of a single brief high-power pulse. Methods of Q-switching include saturable dyes, rotating mirrors, and acoustooptic modulators. Pockel’s cell, an electro-optic modulator that is the most common Q-switch, applies voltage across a crystal to vary polarization. Polarity can be rapidly changed by 90E, making the cell either opaque or transparent to the polarized laser beam. The ‘Q’ refers to the quality factor of the laser cavity, which is defined as the energy stored in the cavity divided by the energy lost per cycle. Rapid extraction of high power is accomplished as the Q-switch changes the quality factor of the cavity from a high to a low Q. Whereas typical mode-locked laser output consists of a train of seven to ten 25-picosecond (psec, 10-12 sec) pulses, at intervals of 5 nsec and contained within a 35-50 nsec envelope, Q-switched laser output generally consists of a single 2-30 nsec pulse. The total energy required for a single Q-switched pulse and a train of mode-locked pulses is the same, but the peak power necessary to cause avalanche ionization must be 100-1000 times greater for mode-locked than Qswitched lasers.6,7 Maximum outputs of most ophthalmic models are approximately 10-30 mJ and 5 mJ for Q-switched and mode-locked lasers, respectively, but because of the greater control and relative safety of the Q-switch, those models employing mode-locking have largely fallen out of favor. Optical breakdown and plasma formation When a target is heated by absorbing radiant energy, the effect is linearly proportional to the cause. In contrast, nonlinear effects are sudden all-or-nothing phenomena. Optical breakdown, a nonlinear reaction, occurs when the laser output is sufficiently condensed spatially and temporally to achieve high irradiance. It is manifested by a spark and accompanied by an audible snap, producing dramatic target damage. When focused to a small spot, typically less than 50 µm in diameter, short-pulsed Nd:YAG lasers can produce enough irradiance, usually 1010-1011 W/cm2, to induce optical breakdown, dissociating electrons from their atoms and creating a plasma. Q-switched pulses cause ionization mainly by focal target heating, in a process called thermionic emission, whereas modelocked ones rely primarily on multiphoton absorption.8 In either case, once the initial free electrons have been generated, plasma expands via electron avalanche or cascade if the irradiance is adequate to cause rapid ionization. Plasma absorbs and scatters incident
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Principles of photodisruption radiation, thereby shielding underlying structures. Plasma radiation absorption and growth both occur through inverse bremsstrahlung, the process of photon absorption and electron acceleration in the presence of an atom or ion. Mechanisms of damage In biological systems, thermal denaturation of protein and nucleic acids is theoretically confined to a radius of 0.1 mm for a 1 mJ pulse.9 As such, although high local temperatures exist briefly, total heat energy is low, and significant clinical photocoagulation does not occur. Several mechanisms combine to generate pressure waves radiating from the zone of optical breakdown, the foremost of which is the rapid plasma expansion that begins as a hypersonic wave.10,11 A secondary source of hypersonic and sonic waves is stimulated Brillouin scattering, in which the laser light generates the pressure wave that scatters it.12 The focal heating may lead to vaporization, melting, and thermal expansion, generating acoustic waves.13 If sufficiently strong, the radiation’s electric field will deform a target through electrostriction, which causes simple Brillouin scattering, and radiation pressure induced by momentum transfer from photons to atoms in inverse bremsstrahlung. The shock wave begins immediately with plasma formation, and expands at a hypersonic velocity of 4 km/sec, falling to sonic velocity within 200 µm. The acoustic transient lasts 50 nsec at a distance of 300 µm from the focal point, while the pressure falls from 1000 to 100 atm within 1 mm.14 The next process is cavitation, or vapor bubble formation. This begins within 50-150 nsec after breakdown in water, expands rapidly for the first 20 µsec, reaches a maximum size of approximately 0.6 mm at 300 µsec, and collapses within 300-650 µsec.10,11 Cavity propagation velocity is about 20 m/sec at 300 µm from the breakdown.15 Many shock waves may be generated along the laser beam’s path as impurities are encountered.16 Damage zone size depends on the irradiance and total energy, the plasma’s duration, and the mechanical properties (including density, mass, tensile strength, and elasticity) of the target tissue.17-20 For many years, most cataract operations have included the insertion of intraocular lenses (IOLs), which were mostly made initially of glass, then polymethylmethacrylate (PMMA), and most recently are typically composed of silicone or other foldable materials. These lenses can affect the intraocular use of lasers, especially for posterior capsulotomy (see below) where occasional damage takes different forms characteristic of the IOL material.21 Such damage is typically more significant for rigid IOLs, which can develop microcracks, melted voids, and large pulverized regions. Unlike the situation in liquids, optical breakdown in rigid lenses may be associated with selffocusing and self-trapping with both nsec and psec pulses.22 The damage threshold for glass is approxi-
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305 mately 100 times greater than that for PMMA, but once glass damage occurs, it tends to be more extensive.23 As damage tends to be cumulative, IOLs may be harmed more by bursts of laser shots than by single pulses. Various rigid IOL designs, including the use of spacers to increase the separation between the IOL and the posterior lens capsule, have been created in the attempt to minimize damage from photodisruption. The threshold for retinal laser injury is inversely proportional to the laser’s wavelength and pulse duration and directly proportional to spot size.24 Since clinical applications of Nd:YAG photodisruption involve energies significantly above retinal damage thresholds, it is important to consider how the retina is protected during these laser procedures. Beam divergence, i.e., the angle formed by the cone of light converging on and diverging from the laser system’s focal point, is the most important element in retinal protection. The border of the laser beam is described as either the 1/e or 1/e2 points of the solid angle. Commercial ophthalmic Nd:YAG lasers usually broaden the laser beam with an inverse galilean telescope and then employ a large-diameter, high-power final focusing lens to achieve the desired combination of cone angle, minimal spot size, and comfortable working distance. As such, for retinal injury to occur during Nd:YAG laser posterior capsulotomy, 96 mJ, some 20 times the energy clinically used, would have to be incident on the cornea. Plasma formation is a secondary factor in retinal protection during photodisruption in the pupillary plane. It absorbs and scatters incident radiation, thereby diminishing the transmission of radiant energy along the beam path. Plasma shielding assumes a more important role in retinal protection during vitreous photodisruption. Nevertheless, the pressure waves still propagate unattenuated, and may cause retinal or choroidal damage even in the absence of suprathreshold radiation levels. Instrumentation Although photodisruption is possible with other lasers and at other wavelengths, including some Nd:YAG harmonics, the fundamental Nd:YAG output at 1064 nm is virtually the only one used in commercial ophthalmic photodisruptors (Fig. 1). Most clinical Nd:YAG lasers employ the fundamental TEM00 mode, so that the spot size, and consequently the energy required for optical breakdown, can be minimized. Beam divergence is generally 0.5-3.0 mrad. The lasers are cooled by ambient air or internally recirculated water, and require only standard 110 V outlets. Whereas 5 mJ is sufficient for most applications, many ophthalmic Q-switched Nd:YAG lasers are capable of producing up to 30 mJ. Higher energies may be needed to cut very dense material and in cases of hazy media, such as corneal edema or blood in the anterior chamber.
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Joel M. Krauss focused high-irradiance pulses can damage the mirrors. Special Nd:YAG gonioscopic lenses are now available that enable better angle visualization and treatment within a short working distance (Fig. 2C). Compact laser contact lenses with anterior curvatures ranging from flatter than the cornea to planoanterior are used for mid- and deep-vitreous work (Fig. 2D). Clinical applications Posterior capsulotomy
Fig. 1. YC-1600 portable Nd:YAG laser. (Reproduced by courtesy of Nidek, Inc., Fremont, CA.)
An aiming beam is required to guide the pulsed, invisible Nd:YAG output. This is achieved with a continuous wave helium:neon (He:Ne) laser that produces 632.8 nm output coaxial with the Nd:YAG’s and below the retinal injury threshold. Since high-peak power pulses cannot be satisfactorily transmitted via fiberoptics, ophthalmic Nd:YAG lasers employ fixed mirrors to guide the output to the patient, who is generally seated opposite the surgeon at a specially configured slit-lamp biomicroscope. The larger the solid cone angle, the lower the energy required for optical breakdown and the risk of IOL or retinal damage, but the greater the chances of beam vignetting during some applications. The slit-lamp design limits the angle to approximately 20°, and most systems employ one of 16E. A ‘heads-up’ feature, available on most current Nd:YAG laser systems, displays important information about treatment parameters. While contact lenses are generally not required for simple posterior capsulotomy, they may be helpful to stabilize the eye, prevent blinking, and maintain a regular optical surface. A suitable lens may have a radius of curvature similar to the cornea’s (8-12 mm), with an additional central high plus power button lens to enhance beam convergence and magnify the surgeon’s view (e.g., the central Abraham lens with a 66D button, Fig. 2A). However, such lenses also provide a smaller field of view through the button lens and limit illumination. A peripheral Abraham button lens (Fig. 2B) is helpful for peripheral iridectomy, since it enhances visualization and control of the photodisruption. Standard mirror lenses such as the Goldmann three-mirror contact lens can be used for gonioscopy treatments, although pupillary treatment with the planosurface would eliminate the beam convergence of the cornea and hence risk retinal injury. Moreover, improperly
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Cataract surgery has largely changed from intracapsular cataract extraction (ICCE), in which the entire lens capsule is removed together with the opaque lens, to extracapsular cataract extraction (ECCE) and phacoemulsification, in which the posterior lens capsule is left in place. This serves to both reduce the incidence of postoperative vitreoretinal complications, such as cystoid macular edema (CME), and to provide support for posterior chamber IOLs. Unfortunately, this membrane often also opacifies over a period of months to years, diminishing the improved vision afforded by removal of the primary cataract.25 Several factors may contribute to optical compromise of initially clear posterior lens capsules. In the early postoperative period, fibrosis can manifest as a gray-white band or plaque-like opacity. Months to years following surgery, opacity may be caused by migration of epithelial cells with formation of small Elschnig’s pearls and bladder cells. Capsular wrinkling can develop as either broad undulations or fine wrinkles. It has been suggested that secondary cataract formation is somewhat less with silicone than with PMMA IOLs,26 although whether this is due to differences in the IOL material or the initial surgical technique remains to be determined. Until the advent of photodisruption, an opacified posterior lens capsule was ruptured by introducing a needle into the eye with the patient seated at the slit-lamp, with all the attendant risks of any invasive procedure. The Nd:YAG laser has proven so successful at sectioning these membranes that it has completely replaced traditional surgical discission. Posterior capsulotomy is associated with a high degree of visual improvement, but is not without occasional complications. These may include IOL marking, retinal detachment, rupture of the anterior hyaloid face, and bleeding from diabetic rubeosis iridis and neovascular glaucoma in diabetics. The heparin coating with which many IOLs are now made may also be compromised.27 However, by far the most commonly encountered difficulty is a transient rise in intraocular pressure (IOP),28 probably caused by impaired aqueous outflow resulting from capsular debris, acute inflammatory cells, and high-molecularweight proteins.29 In an early study on the safety and efficacy of Nd:YAG laser posterior capsulotomy, Keates and as-
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Fig. 2. A: Abraham capsulotomy YAG Laser Lens: a 10-mm diameter with a 66D Magnifying button in the center of the lens enhances visualization and allows precise laser focus on the posterior capsule. B: Abraham Iridectomy YAG Laser Lens: a 10-mm diameter with a 66D magnifying button in the anterior surface of the lens is positioned over the peripheral iris to give a clear view of the iridectomy site and increase laser efficiency. C: Magna View Gonio Laser Lens: used for gonioscopy and laser trabeculoplasty, this lens contains a 62E mirror. The tilted anterior surface corrects image and laser beam astigmatism. D: Peyman Wide Field YAG Laser Lens: an 18-mm lens designed for mid-vitreous work with the Nd:YAG laser. Also available are a 12.5-mm lens for the anterior vitreous and a 25-mm lens for the posterior vitreous. The convex anterior surface of each lens optimizes image magnification and laser performance in the area of interest. (Reproduced by courtesy of Ocular Instruments, Inc., Bellevue, WA.)
sociates30 found that among patients followed for at least six months, 90.1% of pseudophakic individuals had improved visual acuity following laser capsulotomy versus only 69.7% who underwent the traditional surgical discission. Only 3.3% of the laser-treated group experienced decreased vision after capsulotomy, compared with 14.8% of the surgical group. Persistent CME was significantly less in the laser (0.2%) than in the surgical (1.9%) group. In addition to improved visual acuity, patients often experience greater contrast sensitivity and less glare following Nd:YAG laser posterior capsulotomy.31 The brief IOP spike following laser posterior capsulotomy is far more common among patients with aphakia or anterior chamber IOLs than in those with posterior chamber implants. Moreover, the pressure elevation B which usually develops within several hours after treatment, often within one hour in patients whose IOP eventually rises more than 10 mmHg B
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is greatest in individuals with high pre-laser IOP (above 20 mmHg) and low facility of outflow.30,32 Steinert and colleagues33 retrospectively analyzed long-term complications in 897 patients who underwent Nd:YAG laser posterior capsulotomy approximately two years after ECCE. There was no correlation between the development of complications and any discernible factor common to affected patients. Eleven patients (1.2%) developed CME at an average of 4.8 months after laser therapy. Five patients were treated with corticosteroids, two with systemic indomethacin, and one (with pre-existing diabetic retinopathy) with photocoagulation at 577 nm. These individuals may have experienced a greater inflammatory response to their original cataract surgery than do most individuals, as evidenced by postoperative reports and the use of steroids for an average of 2.4 months following surgery. Retinal detachment occurred in eight patients (0.9%) at an average of 13.5
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months following laser capsulotomy. Only three patients had at least 3D of myopia, and all eight were treated with scleral buckle. That three patients’ detachments developed more than a year after laser treatment implies that the laser did not directly cause the detachments and highlights the need for long-term follow-up. Seven patients (0.8%) developed new, persistent glaucoma following laser capsulotomy. Mean IOP was 18.7 mmHg before ECCE and 19.6 mmHg before capsulotomy, but reached a mean maximum of 25 mmHg during the first post-laser week. Mean IOP at six months was 22.3 mmHg on medication, and at an average last follow-up of 28.3 months it was 20.4 mmHg. Laser trabeculectomy was necessary in one patient. A posterior chamber IOL may be associated with a lower incidence of transient IOP rise, but does not necessarily protect against persistent IOP elevation. All prospective laser capsulotomy patients require a thorough medical and ophthalmological history and physical. Direct ophthalmoscopy is the most reliable technique for assessing capsular opacity. While secondary cataracts can substantially diminish vision, their removal does not guarantee normal vision. Especially in older patients, concomitant ocular pathology (such as macular degeneration or CME) may impair vision even after capsulotomy. The laser interferometer and potential acuity meter can penetrate mild to moderate capsular opacity and assess macular function. In cases where decreased visual acuity is disproportionate to capsular opacity, a useless or even deleterious procedure may be avoided. Dilatation facilitates laser capsulotomy, but may be omitted in the absence of a miotic pupil. Before dilatation, it may be useful to place a marker shot, to aid subsequent identification of the true visual axis. To prevent iris capture of a posterior chamber IOL, a single drop of 2.5% phenylephrine, together with a drop of 0.5-1% tropicamide if necessary, is advisable. Topical anesthesia is usually required only when using a contact lens. A retrobulbar injection may be needed for akinesia in rare cases such as patients with nystagmus. In most instances, a posterior capsule can be opened with minimal pulse energies of 1-2 mJ from the Nd:YAG laser. Shots are placed along tension lines, as indicated by capsular wrinkles, to create the most efficient opening. Except in patients at high risk for retinal detachment or others in whom a smaller capsulotomy opening may be safer, the opening should be as large as the pupil in ambient light. Figure 3 shows a posterior capsulotomy behind a posterior chamber IOL. Several steps can be employed to minimize IOL damage, including using the lowest possible energy settings, stabilizing the eye with a contact lens, beginning treatment in areas of capsule-IOL separation, and in some cases focusing in the anterior vitreous and allowing the anterior radiation of the shockwave to rupture the capsule. Some surgeons have reported displacement of foldable IOLs into the vitreous following capsulotomy.34 In aphakic patients, focusing
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Fig. 3. Secondary cataract after Nd:YAG laser capsulotomy. The IOL edge and haptic are visible. (Reproduced by courtesy of Roger Steinert, MD.)
anterior to the capsule may help preserve the anterior hyaloid membrane. Post-laser pharmacotherapy varies widely, but typical practice includes the use of a topical steroid (e.g., prednisolone acetate 1% or dexamethasone 0.1%) and a pressure-reducing medication (e.g., β-blocker or apraclonidine) immediately following treatment. Patients with an increase of at least 5 mmHg above baseline IOP or those with pre-existing glaucoma require more aggressive prophylaxis and frequent IOP checks. However, prophylactic treatment (e.g., 1% apraclonidine35 or 0.5% levobunolol36) prior to laser capsulotomy offers the best insurance against IOP spikes. Pre-laser oral (acetazolamide) and topical (2% dorzolamide) carbonic anhydrase inhibitors have been shown to provide comparable prophylaxis.37 Fankhauser and associates5 were the first to use the Q-switched Nd:YAG laser to cut pupillary membranes. Though less common than posterior capsulotomy, membranectomy can optically clear the pupil in eyes with significant pathology that are poor candidates for surgery or would otherwise require major operations. In contrast to posterior capsules, membranes have little or no elastic properties and may have to be chipped away at the edges with pulses of 4-12 mJ. Dense membranes may need multiple treatment sessions. Iridectomy and other anterior segment applications Argon laser iridectomy has largely supplanted the traditional surgical approach, but there remain many instances in which its use is problematic. The argon laser relies on coagulation, vaporization, and necrosis to cut through tissue, and light blue or gray irises may not absorb sufficient energy. Conversely, the strong absorption by dark brown irises may generate a char that impedes further penetration. Since photodisruption does not depend on target pigmentation, the short-pulsed Nd:YAG laser represents an attractive alternative for creating iridectomies. This was verified in clinical studies, which demonstrated that the Q-switched Nd:YAG laser can create openings in the
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Principles of photodisruption iris that, unlike some openings created with the argon laser, do not gradually close.38 The Fankhauser group5 performed the first Qswitched Nd:YAG laser iridectomy. Some phakic patients experienced opacification for 1-2 mm in the underlying anterior lens capsule, but this response was transient, and there was no cataract formation. Aside from small self-limited hemorrhages that are occasionally encountered with the Nd:YAG laser,39 complications with that laser are usually of equal or less severity than with the argon laser. However, as with all photodisruption procedures, the Nd:YAG would be unable to coagulate any significant bleeding that might occur. Unlike cutting through the iris sphincter, where hemorrhage is expected, peripheral iridectomy seldom causes bleeding. Full-thickness iridectomy is achieved more quickly B usually with just a single shot B and reliably with the Nd:YAG laser than with the argon laser. Both lasers are associated with a small IOP rise that resolves within several hours. Nd:YAG laser iridectomy has proved effective in cases in which the argon laser has failed,40 while use of the two lasers together on dark irises may allow less energy to be employed than with either laser alone.41,42 To facilitate laser iridectomy, the iris is drawn taught by instilling miotic drops, such as pilocarpine 4%. Following administration of topical anesthetic, an Abraham lens with a peripheral 66D button lens is applied. Treating the basal iris is safest as it is not directly apposed to the crystalline lens. Penetration is facilitated and bleeding avoided by targeting a thin area, such as a crypt. Openings are often achieved with a laser energy of 4-8 mJ; although a single shot is desirable, up to four are often needed. A burst setting of more than several pulses risks injury to the underlying crystalline lens capsule. With the proper technique, the loss of corneal endothelial cells overlying the treatment site can be minimized.43 Strong topical steroids four times a day are begun immediately following the laser iridectomy. If significant inflammation or bleeding is present and the iridectomy is definitely patent, intermittent short-term dilatation may help prevent synechia formation. Corneal edema and haze, anterior chamber reaction, and iris congestion may make argon laser iridectomy impossible in aphakic and pseudophakic patients with acute angle-closure glaucoma. Photodisruption with the Nd:YAG laser is the treatment of choice, even in instances in which the argon laser can penetrate the iris. The Nd:YAG laser’s success in creating an ‘anterior hyaloidectomy’ to cure malignant (ciliovitreal block) glaucoma demonstrates the pathophysiological role of the anterior hyaloid face.44 To ensure the long-term patency of at least one iridectomy as the iris bombé is relieved, at least three iridectomies should be made. The laser is then aimed at the anterior vitreous through the iridectomy or pupil. Cycloplegia, mydriasis, and intensive topical steroid treatment is then begun. The Q-switched Nd:YAG laser is useful in
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309 coreoplasty, although rarely in phakic patients. Treatment is generally begun in the peripheral iris stroma, with numerous 6-10 mJ shots progressing toward the sphincter and pupil (and hence likely causing hemorrhage) only at the end of the treatment session. Additional anterior segment applications of Nd:YAG laser photodisruption include severing of retained anterior capsule tags following ECCE, creating corneal stromal puncture in cases of refractory traumatic recurrent erosion,45 zonulysis of subluxated crystalline lenses (e.g., in individuals with Marfan’s syndrome), 46 synechialysis (especially in relatively unpigmented tissue that would not be amenable to therapy with the argon laser), and anterior vitreolysis in patients with CME and Irvine-Gass syndrome.47,48 Posterior segment While technically more demanding and potentially more dangerous, Nd:YAG laser photodisruption may also be applicable to pathology in the posterior segment. Relatively avascular vitreous membranes associated with significant retinal traction can be cut with this laser.49,50 Experimental vitreous membranes in rabbits have been successfully sectioned with 4 mJ pulses as close as 4 mm to the retina, without retinal injury.51 Since these membranes may be complex and fibrous, hundreds or thousands of pulses, often in multiple sessions, may be necessary. Nd:YAG laser photodisruption of the anterior surface of preretinal hemorrhages has allowed absorption of blood into the vitreous.52 Aside from the risk of retinal or choroidal hemorrhage, which rises exponentially with proximity to the retina, a lens (crystalline or IOL) may also be damaged if work is performed too close to its posterior surface. However, despite initial concern that photodisruption of the posterior lens capsule or vitreous may cause liquefaction and other vitreous disturbance, Krauss et al.53 employed MRI and other techniques to demonstrate that this process does not significantly affect the structural integrity of the normal vitreous body. New laser techniques The Q-switched Nd:YAG laser is the quintessential ophthalmic photodisruption laser. However, there are several other photodisruption lasers that have been used in the lab or clinically whose mechanisms of tissue interaction are less consistently described. Such nebulous terminology may be a matter of semantics, but it does have some basis in biophysics. While the argon and other lasers with msec or longer output have almost exclusively thermal (photocoagulation) effects, the ultrashort-pulsed Nd:YAG laser with nsec or shorter output has mechanical (photodisruption) effects, and the excimer laser (which also has nsec output) vaporizes tissue via ablation, there are other lasers with intermediate or variable pulse durations and unique properties whose efficacy may indeed be based
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on a combination of mechanisms. Two lasers in the last category are the erbium:yttrium aluminum garnet (Er:YAG) and neodymium:yttrium lithium fluoride (Nd:YLF). The Er:YAG laser at 2.94 µm and output in the µsec range has been tested on virtually every portion of the eye. Although it is sometimes described as causing photodisruption, the Er:YAG is more commonly said to effect ablation. Its various applications B most notably vitreolysis and ‘photophacoablation’ of cataractous lenses B are described elsewhere in this volume. The Nd:YLF laser, with primary emission at 1053 nm and pulses ranging from psec to fsec, has similarly been said to cause ablation, but its interaction with tissue is more accurately described as photodisruption, and hence this laser will be considered here. While the ArF excimer laser at 193 nm remains the paramount refractive laser, researchers continue to examine other lasers that may offer certain clinical or technical advantages. One early focus of such work was intrastromal treatment, since before widespread adoption of LASIK, excimer laser photorefractive surgery necessarily involved removal of the epithelium and obliteration of Bowman’s membrane. Since there is rapidly increasing corneal transmission beginning at about 300 nm, ultra-short-pulsed lasers with emission between approximately 400 and 1500 nm would be able to non-invasively cause photodisruption at any specifically targeted corneal level, rather than relying on absorption by corneal chromophores.54 Indeed, Nd:YAG laser photodisruption with nsec pulses can create intrastromal vacuoles, albeit on a scale of negligible clinical value.55 Although Nd:YAG laser psec pulses have an optical breakdown threshold of only about 20 µJ and appear to cause correspondingly less shockwave emission and cavitation bubble expansion than nsec pulses do,56,57 they are still incapable of equalling the precision of the excimer laser.58 The Nd:YLF is another solid state laser that has been investigated for a variety of ophthalmic applications. Nd:YLF has a larger fluorescence bandwidth than does Nd:YAG and hence can produce pulses of shorter duration. Some researchers used early Nd:YLF laser prototypes experimentally for iridectomy, pupilloplasty, posterior capsulotomy, and vitreolysis, but results were generally not as good as with other laser or conventional surgical techniques.59,60 As such, most studies using the Nd:YLF laser have continued to concentrate on potential corneal applications. Initial results using the Nd:YLF laser operating in the psec domain for intrastromal photodisruption in animal and cadaver eyes were unsatisfactory.61-66 Given the even further reduction in shockwave and cavitation effects B and hence localized tissue damage B caused by photodisruption pulses in the fsec (10-15 sec) range,67 studies with the Nd:YLF laser have increasingly focused on fsec technology. A group at Saint Louis University has extensively studied corneal applications of Nd:YLF laser photodisruption.
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They found that while the Nd:YLF laser producing psec pulses (with energies of 25 µJ and spot separations of 10-20 µm) could serve as a non-invasive microkeratome and create corneal flaps for LASIK with the excimer laser,68 the Nd:YLF laser operating in the fsec range (with pulse energies of only 4-8 µJ and spot separations of 10-15 µm) produced far smoother lamellar surfaces that required no additional mechanical dissection.69 A couple of companies now manufacture fsec Nd:YLF lasers primarily as alternatives to the traditional surgical microkeratome. Figure 4 shows the Pulsion FS laser from IntraLase Corp., a company founded in 1997 by two University of Michigan researchers. As of mid-2002, this device has received 510(k) clearance from the Food and Drug Administration for testing in 13 sites in the USA. The laser produces fsec pulses focused to a spot size of approximately 3 µm. The device attaches to the patient’s eye via a suction ring, but IOP is only elevated to about 30-40 mmHg. A scanning system strings the pulses together at a rate of 10,000 Hz, creating intrastromal lamellar incision planes, and ultimately a flap, in 3060 sec in preparation for excimer LASIK (Fig. 5). Flap creation is one of the most problematic steps in conventional LASIK, the most common complication of which is irregular astigmatism. The fsec laser virtually eliminates the risk of flap-associated complications such as decentered flaps, epithelial ingrowth, epithelial abrasion, or perforations. In principle, flap thickness, diameter, edge bevel, and hinge location can all be adjusted by customizing the laser pulse pattern. The Pulsion FS laser has also been used in
Fig. 4. Pulsion FS femtosecond laser. (Reproduced by courtesy of IntraLase Corp., Irvine, CA.)
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Fig. 5. Schematic of intrastromal photodisruption with the Pulsion FS femtosecond laser creating a corneal flap for excimer laser ablation. A: The 1053-nm transmission of the Nd:YLF laser is readily transmitted by the cornea and can be targeted at any portion thereof. B: With the laser focused in the stroma to a spot size of 3 µm, the beam is scanned at a repetition rate of 10,000 Hz, creating a lamellar dissection plane. C: The resulting corneal flap is peeled back, exposing the underlying stroma for excimer LASIK. D: Following the ablation, the flap is replaced. (Reproduced by courtesy of IntraLase Corp., Irvine, CA.)
early clinical trials to create channels for placement of intrastromal corneal ring segments, affording greater precision and flexibility than is possible with manual dissection using a diamond knife. Conclusions and future developments Photodisruption with the Nd:YAG laser is now a relatively mature procedure, although techniques continue to be refined. No current laser is likely to supplant the Nd:YAG for its full range of applications. Diode lasers, aided by advances in endoprobes and indirect ophthalmoscopy, are already being used to treat a variety of vascular diseases, and are becoming the treatment of choice for retinopathy of prematurity. Diode laser technology is constantly improving, with output at higher energies and shorter pulse durations and wavelengths. New diode lasers may continue to replace larger, more expensive ophthalmic lasers, conceivably someday including the Nd:YAG photodisruptor. Laser photorefractive surgery has obvious considerable clinical and commercial value, and many lasers at a variety of wavelengths and pulse durations have been touted as potential alternatives to the ArF
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excimer. With further refinement in equipment and technique, Nd:YLF intrastromal photodisruption may represent a single-step alternative to traditional excimer LASIK by directly reprofiling the cornea. Another solid state fsec laser that has been the subject of recent pre-clinical studies for intrastromal photodisruption is the titanium-sapphire laser, with output at 780 nm and 200 fsec.70,71 Laser technology in general, and ophthalmic lasers in particular, are dynamic fields that are the subjects of active basic and clinical research. New laser instruments and techniques are continually being developed, and few go untested for potential ophthalmic use. The pace and scope of progress suggest that the field of ophthalmic lasers, including photodisruption, will continue to evolve.
References 1. Longsworth JW: Photophysics. In: Regan JD, Parrish JA (eds) The Science of Photomedicine, p 43. New York, NY: Plenum 1982 2. McClung FJ, Hellwarth RW: Giant optical pulsating from ruby. J Appl Phys 33:828-831, 1967 3. Krasnov M: Laser-puncture of the anterior chamber angle in glaucoma. Vestn Oftalmol 3:27-31, 1972
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4. Aron-Rosa D, Aron JJ, Greisemann J et al: Use of the neodymium-YAG laser to open the posterior capsule after lens implant surgery: a preliminary report. J Am Intraocul Implant Soc 6:352-354, 1980 5. Fankhauser F, Roussel P, Steffen J et al: Clinical studies on the efficiency of a high power laser radiation upon some structures of the anterior segment of the eye. Int Ophthalmol 3:129-139, 1981 6. Steinert RF, Puliafito CA, Trokel S: Plasma formation and shielding by three ophthalmic Nd-YAG lasers. Am J Ophthalmol 96:427-434, 1983 7. Fradin DW, Bloembergen N, Letellier JP: Dependence of laser-induced breakdown field strength on pulse duration. Appl Phys Lett 22:635-637, 1973 8. Ready JF: Effects of High-Power Laser Radiation, pp 133143, 215-217. New York, NY: Academic Press 1971 9. Hu CL, Barnes FS: The thermal-chemical damage in biological material under laser irradiation. IEEE Trans Biomed Eng 17:220-229, 1970 10. Felix MP, Ellis AT: Laser-induced liquid breakdown: a stepby-step account. Appl Phys Lett 19:484-486, 1971 11. Lauterborn W: High-speed photography of laser-induced breakdown in liquids. Appl Phys Lett 21:27-29, 1972 12. Brewer RJ, Rieckhoff KE: Stimulated Brillouin scattering in liquids. Phys Rev Lett 13:334-336, 1964 13. Cleary SF, Hamrick PE: Laser-induced acoustic transients in the mammalian eye. J Acoust Soc Am 46:1037-1044, 1969 14. Van der Zypen E, Fankhauser F, Bebie H et al: Changes in the ultrastucture of the iris after irradiation with intense light. Adv Ophthalmol 39:59-180, 1979 15. Fujimoto JG, Lin WZ, Ippen IP et al: Time-resolved studies of Nd:YAG laser induced breakdown: plasma formation, acoustic wave generation, and cavitation. Invest Ophthalmol Vis Sci 26:1771-1777, 1985 16. Carome EF, Carreira EM, Prochaska CJ: Photographic studies of laser-induced pressure impulses in liquids. Appl Phys Lett 11:64-66, 1967 17. Mainster MA, Sliney DH, Belcher CD et al: Laser photodisruptors: damage mechanisms, instrument design, and safety. Ophthalmology 90:973-991, 1983 18. Taboada J: Interaction of short laser pulses with ocular tissues. In: Trokel S (ed) YAG Laser Ophthalmic Microsurgery, pp 15-38. Norwalk, CT: Appleton-Century-Crofts 1983 19. Smith WL, Liu P, Bloembergen N: Superbroadening in water and deuterium by self-focused picosecond pulses from a neodymium doped YAG laser. Phys Rev (A) 15:2396-2403, 1977 20. Anthes JP, Bass M: Direct observation of the dynamics of picosecond-pulse optical breakdown. Appl Phys Lett 31:412414, 1977 21. Dick B, Schwenn O, Pfeiffer N: Extent of damage to different intraocular lenses by neodymium:YAG laser treatment: an experimental study. Klin Mbl Augenheilk 211:263-271, 1997 22. Ashkinadze BM, Vladimirov VI, Likhachev VA et al: Breakdown in dielectrics caused by intense laser radiation. Sov Phys JETP 23:788-797, 1966 23. Loertscher H: Laser-induced breakdown for ophthalmic applications. In: Trokel S (ed) YAG Laser Ophthalmic Microsurgery, p 39. Norwalk, CT: Appleton-Century-Crofts 1983 24. Gibbons WD, Allen RG: Retinal damage from suprathreshold Q-switch laser exposure. Health Phys 35:461-469, 1978 25. Sinskey RM, Cain W: The posterior capsule and phacoemulsification. J Am Intraocul Implant Soc 4:206-207, 1978 26. Pradella SP, Taumer R: Frequency of Nd:YAG capsulotomy
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Principles of photodisruption 45. Geggel HS: Successful treatment of recurrent corneal erosion with Nd:YAG anterior stromal puncture. Am J Ophthalmol 110:404-407, 1990 46. Tchah HW, Larson RS, Nichols BD et al: Neodymium:YAG laser zonulysis for treatment of lens subluxation. Ophthalmology 96:230-234, 1989 47. Katzen LE, Fleischman JA, Trokel S: YAG laser treatment of cystoid macular edema. Am J Ophthalmol 95:589-592, 1983 48. Steinert RF, Wasson PJ: Neodymium:YAG anterior vitreolysis for Irvine-Gass cystoid macular edema. J Cataract Refract Surg 15:304-307, 1989 49. Jagger JD, Hamilton AM, Polkinghorne P: Q-switched neodymium-YAG laser vitreolysis in the treatment of posterior segment disease. Arch Clin Exp Ophthalmol 228:222225, 1990 50. Berglin L, Stenkula S, Crafoord S et al: A new technique of treating rhegmatogenous retinal detachment using the Qswitched Nd:YAG laser. Ophthalmic Surg 18:890-892, 1987 51. Puliafito CA, Wasson PJ, Steinert RF et al: Nd-YAG laser surgery on experimental vitreous membranes. Arch Ophthalmol 102:843-847, 1984 52. Iijima H, Satoh S, Tsukahara S: Nd:YAG laser photodisruption for preretinal hemorrhage due to retinal macroaneurysm. Retina 18:430-434, 1998 53. Krauss JM, Puliafito CA, Miglior S et al: Vitreous changes after neodymium-YAG laser photodisruption. Arch Ophthalmol 104:592-597, 1986 54. Krauss JM, Puliafito CA, Steinert RF: Laser interactions with the cornea. Surv Ophthalmol 31:37-53, 1986 55. Taboada J, Poirier RH, Yee RW et al: Intrastromal photorefractive keratectomy with a new optically coupled laser probe. Refract Corneal Surg 8:399-402, 1992 56. Vogel A, Busch S, Jungnickel K et al: Mechanisms of intraocular photodisruption with picosecond and nanosecond laser pulses. Lasers Surg Med 15:32-34, 1994 57. Vogel A, Capon MR, Asiyo-Vogel MN et al: Intraocular photodisruption with picosecond and nanosecond laser pulses: tissue effects in cornea, lens, and retina. Invest Ophthalmol Vis Sci 35:3032-3044, 1994 58. Vogel A, Asiyo-Vogel M, Birngruber R: Intrastromal refractive corneal surgery with pico-second Nd:YAG laser pulses. Ophthalmologe 91:655-662, 1994 59. Cohen BZ, Wald KJ, Toyama K: Neodymium:YLF picosecond laser segmentation for retinal traction associated with
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proliferative diabetic retinopathy. Am J Ophthalmol 123:515523, 1997 Geering G, Roider J, Schmidt-Erfurt U et al: Initial clinical experience with the picosecond Nd:YLF laser for intraocular therapeutic applications. Br J Ophthalmol 82:504-509, 1998 Remmel RM, Dardenne CM, Bille JF: Intrastromal tissue removal using an infrared picosecond Nd:YLF ophthalmic laser operating at 1053 nm. Lasers Light Ophthalmol 4:169173, 1992 Brown DB, O’Brien WJ, Schultz RO: Nd:YLF picosecond laser capabilities and ultrastructure effects in corneal ablations. Invest Ophthalmol Vis Sci 34(Suppl):1246, 1993 Itoi M, Bassage S, Del Cerro M et al: Corneal incisions utilizing the 1053 nm picosecond Nd:YLF ophthalmic laser. Cornea 15:2-8, 1996 Krueger RR, Quantock AJ, Juhasz T et al: Ultrastructure of picosecond laser intrastromal photodisruption. J Refract Surg 12:607-612, 1996 Ito M, Quantock AJ, Malhan S et al: Picosecond laser in situ keratomileusis with a 1053-nm Nd:YLF laser. J Refract Surg 12:721-728, 1996 Gimbel HV, Coupland SG, Ferensowicz M: Review of intrastromal photorefractive keratectomy with the neodymium: yttrium-lithium-fluoride laser. Int Ophthalmol Clin 37:95102, 1997 Juhasz T, Kastis GA, Suarez C et al: Time-resolved observations of shock waves and cavitation bubbles generated by femtosecond laser pulses in corneal tissue and water. Lasers Surg Med 19:23-31, 1996 Krueger RR, Marchi V, Gualano A et al: Clinical analysis of the neodymium:YLF picosecond laser as a microkeratome for laser in situ keratomileusis. J Cataract Refract Surg 24:1434-1440, 1998 Kurtz RM, Horvath C, Liu HH et al: Lamellar refractive surgery with scanned intrastromal picosecond and femtosecond laser pulses in animal eyes. J Refract Surg 14:541548, 1998 Lubatschowski H, Maatz G, Heisterkamp A et al: Application of ultrashort laser pulses for intrastromal refractive surgery. Graefe’s Arch Clin Exp Ophthalmol 238:33-39, 2000 Heisterkamp A, Ripken T, Lutkefels E et al: Optimizing laser parameters for intrastromal incision with ultra-short laser pulses. Ophthalmologe 98:623-628, 2001
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Ultrastructural effects of laser irradiation at the anterior chamber angle E. Van der Zypen Institute of Anatomy, University of Bern, Bern, Switzerland
Keywords: laser trabeculoplasty, selective trabeculoplasty, trabeculopuncture, laser cyclodialysis, trabecular meshwork, mechanisms related to trabeculoplasty, histology, ultrastructure
History The transparency of the cornea of the eye provides an opportunity for intraocular surgical interventions with a laser beam without opening the bulbus. In 1973, two groups of scientists described methods to reduce the intraocular pressure (IOP) by irradiation of the trabecular meshwork using a laser. In 1973, Krasnov1 described a ‘microrupture’ technique using a Q-switched ruby laser in open-angle glaucoma patients. His technique attempted to provide a direct opening pathway into Schlemm’s canal. A temporary reduction of IOP in all patients, accompanied by improved outflow facilities, was noted. In that same year, Worthen and Wickham2 described the morphological effects of argon laser irradiation of the trabecular meshwork in the monkey. Later on, Wise and Witter3 accomplished lowering of the IOP in primary open-angle glaucoma patients by irradiation of the trabecular meshwork without perforating Schlemm’s canal, using an argon laser. This method of argon laser trabeculoplasty (ALT) is now used world-wide in the treatment of open-angle glaucoma (for a review, see Reiss et al.4 and Brilakis et al.5). Wavelength dependence Argon (wavelength 488-514 nm), Nd:YAG (523 nm), diode (810 nm), holmium, erbium:YAG, and other laser wavelengths are used for irradiation of the trabecular meshwork.7-14 CW lasers (10 ms - 0.5 s) produce coagulation by heat which is visible by an increased intensity of staining in histological sections (Figs. 1 and 2).
The argon laser does not cause very deep tissue damage (Fig. 1). Ultrastructural analysis shows that the collagen fibrils disintegrate into subfibrillar fragments, which lose their periodicity. Such changes correspond to an exposure of about 200°C, as systematic comparative analysis of heat-induced phenomena has shown.15,16,24 It has been found that penetration of heat induced by the Nd:YAG laser is six times deeper than that induced by the argon laser (Fig. 2).16 Nevertheless, in general, it can be assumed that trabecular photocoagulation is not a process that depends upon wavelength.17
Effects of dependence on pulse duration There are significant differences regarding irradiation with cw- or Q-switched lasers. Using a cw Nd:YAG laser for irradiation of the iris, a predominantly thermal coagulation effect is induced (Fig. 3). Collagen fibrils are damaged and degenerate into subfibrillar fragments, loosing their periodicity. The deepness of the crater generated depends on exposure time and mean power. Using a Q-switched laser, an eruption of the tissue may be observed (Fig. 4), predominantly caused by physical forces such as shock waves. The most conspicuous ultrastructural changes are vacuolation, disruption, and fragmentation of melanin granules.15 From consideration of these different morphological effects, it can be assumed that the micropuncture technique described by Krasnov,1 using a Q-switched ruby laser, corresponds to a microexplosion effect with mechanical destruction of the trabecular meshwork, including the inner wall of Schlemm’s canal. In contrast ALT, conceived by Wise and Witter,3 is based on a coagulation effect and does not injure Schlemm’s canal.
Address for correspondence: Prof. Dr. med. E. Van der Zypen, Abt. für Angewandte und Topographische Anatomie, Anatomisches Institut, Universität Bern, Bühlstrasse 26, Postfach, CH-3000 Bern 9, Switzerland. e-mail:
[email protected] Lasers in Ophthalmology – Basic, Diagnostic and Surgical Aspects, pp. 315–331 edited by F. Fankhauser and S. Kwasniewska © 2003 Kugler Publications, The Hague, The Netherlands
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Fig. 1. Trabecular meshwork of a human autopsy eye, treated by argon laser (exposure duration: 0.5 s; pulse energy: 2.2 W). The increased intensity of staining with azure-II-methylenblue indicates the heat-damaged area. Semithin section; negative magnification ×100.
The deepness of the coagulation effect depends on exposure time and mean power. The short- and long-term effects of both methods of irradiation of the chamber angle have been proved by our group in monkeys. Origins of glaucoma
Fig. 2. Trabecular meshwork of a human autopsy eye near Schwalbe’s line, treated by free-running Nd:YAG laser (exposure duration: 10 ms; pulse energy: 2 J). The penetration of heat induced by the Nd:YAG compared to the argon laser is five times greater. Semithin section stained with azure-IImethylenblue; negative magnification ×64. (Reproduced from Van der Zypen and Fankhauser16 by courtesy of the publisher.)
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Up to the present, the origins of the primary openangle glaucoma are not well understood. An accumulation of intercellular material within the juxtacanalicular region of the trabecular meshwork usually is found in open-angle glaucoma.18 It is thought that the large accumulation of basement membrane-like material with heparin sulfate-type proteoglycans could be one of the causes of the increase of IOP in glaucoma.19 The increase of collagen IV, laminin, and fibronectin found in the basement region of the inner wall and at the juxtacanalicular tissue may be another reason for the increase in IOP.5,20 The aim of laser treatment of the chamber angle in primary open-angle glaucoma may be degradation of the extracellular juxtacanalicular proteoglycans or stimulation of the turnover of the extracellular matrix. ALT is also thought to stimulate the degradation of juxtacanalicular material. In organ cultures of the trabecular meshwork, stromolysin is thought to degrade the trabecular proteoclycans. The expression and secretion of stromolysin is thought to be stimulated by ALT.21,22
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Fig. 3. Diagram of the central section of a non-perforating lesion produced by argon laser irradiation in the iris of a pigmented rabbit. Around of the mouth of the crater is an elevated rim composed of displaced damaged material. The wall of the crater lumen consists of a vacuolated, homogeneous mass of collagen (insert bottom center). Granular structures can be seen 30 µm away from the crater lumen, which are the remains of disintegrated collagen fibers (insert bottom right). The melanin granules emanating from disrupted melanocytes appear morphologically normal (insert bottom left). Exposure duration: 300 ms; pulse energy: 225 J. (Reproduced from Van der Zypen et al.15 by courtesy of the publisher.)
Fig. 4. Diagram of the central section of a lesion produced by Q-switched ruby laser irradiation of the iris of a pigmented rabbit. At both the anterior and posterior surfaces of the iris, extrusion of tissue can be seen which presumably arise from the dissipation of forces generated by microexplosions. The collagen fibers are fragmented, but the fragments retain their normal periodicity (insert bottom center). In these lesions (unlike those after argon laser irradiation), the isolated melanin granules from ruptured melanocytes are shattered (insert bottom left). Further disorganization arises through the traumatic interruption of blood vessels (insert bottom right). Exposure duration: 30 ns; pulse energy: 200 mJ. (Reproduced from Van der Zypen et al.15 by courtesy of the publisher.)
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Morphological effects of trabeculopuncture By irradiating the chamber angle of monkey eyes (Macaca speciosa) using a Q-switched Nd:YAG laser (exposure duration 30 ns), mechanical destruction of the trabecular meshwork, including the opening of Schlemm’s canal in the anterior cham-
ber, may be produced (Fig. 5). Then a widely gaping of Schlemm’s canal is created.23,52 Similarly, the endothelial cells of the outer wall are destroyed to a large extent (Fig. 6). However, the scleral spur usually remains intact. The width of the opening between the lamellae results from retraction of the disrupted trabecular lamellae at the ciliary muscle.
Fig. 5. Chamber angle of a human cadaver eye in meridional section. Free passage between Schlemm’s canal (S) and the anterior chamber is produced by a Q-switched Nd:YAG laser (exposure duration: 50 ns; pulse energy: 200 mJ). T: trabecular meshwork; SS: scleral spur; Sc: sclera; O: outflow canal. Scanning electron microscopic (SEM) picture; negative magnification ×160. (Reproduced from Van der Zypen and Fankhauser52 by courtesy of the publisher.)
Fig. 6. Open communication between Schlemm’s canal (S) and the anterior chamber. The damage to the trabecular meshwork and the inner wall of Schlemm’s canal was produced by a Q-switched Nd:YAG laser (pulse duration: 30 ns; pulse energy: 110 mJ; theoretical focus diameter: 15 µm). Macaca speciosa. The border of the sagittal section is marked by a dark line. From the dark line to the right, the structures can be seen in frontal view. Sp: scleral spur; Cf: ciliary folds; C: cornea; SEM micrograph; negative magnification ×50. (Reproduced from Van der Zypen et al.23 by courtesy of the publisher.)
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Fig. 7. Trabecular meshwork (T) of a monkey (Macaca speciosa) 18 months after irradiation with a Q-switched Nd:YAG laser (exposure duration: 35 ns; pulse energy: 200 mJ). A continuous layer of corneal endothelial cells covers the irradiated area. J: iris; SEM micrograph; negative magnification ×100. (Reproduced from Van der Zypen and Fankhauser25 by courtesy of the publisher.)
Fig. 8. Stronger magnification of Figure 7. A continuous layer of hexagonal corneal endothelial cells (Co) covers the irradiated area of the trabecular meshwork (T) 18 months after irradiation with a Q-switched Nd:YAG laser. SEM picture; negative magnification ×400. (Reproduced from Van der Zypen and Fankhauser53 by courtesy of the publisher.)
Collapse of Schlemm’s canal can be observed at the border of the irradiated area, and the inner wall of the canal is in close contact with the outer wall. In trabeculopuncture of the monkey eye using a Q-switched laser, damage to the corneal endothelial cells and Descemet’s membrane inevitably occurs. This injury to the peripheral cornea leads to stimulation of the corneal endothelial cells, which grow into the damaged area. Eighteen months after
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opening Schlemm’s canal in Macaca speciosa eyes, the original site of perforation is covered by an continuous layer of corneal endothelial cells (Figs. 7 and 8). No gaps were seen in this tight cell layer.53 On transmission electron microscopy, pigment granules were observed in some of these endothelial cells, which covered tight collagen scar tissue.24 In human eyes, Schlemm’s canal is located in a more posterior direction towards the equator. For
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this reason, it could be possible that, in trabeculoplasty of human eyes, the corneal endothelium is better preserved than in monkey eyes, in which Schlemm’s canal is located near Schwalbe’s line (Fig. 9). In glaucomatous human eyes, no scar formation or endothelium proliferation was seen up to three years after laser trabecular ablation with removal of trabecular tissue and opening of Schlemm’s canal using an erbium:YAG laser (pulse energy: 5-7 mJ; pulse duration: 200 sec).14 Morphological effects of cyclodialysis By aiming the laser beam at the center of the uveal trabecular meshwork between the iris root and the posterior rim of the cornea (Fig. 9), cyclodialysis is created (Fig. 10). However, Schlemm’s canal is not opened with this technique.25 Histological investigations have shown that the sclera only appears to be slightly damaged. Bundles of collagen fibers are formed the bottom of Schlemm’s canal. In long-term studies (16 months after cyclodialysis), it has been observed that the cyclodialysis canal remains open (Fig. 11). The border of the canal is covered by a monolayer of connective tissue cells (Fig. 12). Collagen fibers are seen among the cell processes of the fibroblasts, which means that the cellular wall of the cyclodialysis canal is discontinuous, which enables the aqueous humor to enter deeper tissue compartments. Three routes for the aqueous humor to flow out,
Fig. 9. Semi-schematic view of the angle of an anterior chamber of a monkey (Macaca speciosa). T: uveal trabecular meshwork; S: Schlemm’s canal with collector channel; P: scleral spur with insertion of the outer portion of the ciliary muscle (M); A: sclera; B: supraciliary space; F: ciliary folds; I: iris. Two sites of impact of the laser beam, either opening Schlemm’s canal or the supraciliary space can be seen. (Reproduced from Van der Zypen and Fankhauser25 by courtesy of the publisher.)
propagating from a cyclodialysis canal, can be seen when HgS particles are used as a tracer substance.24,25,26 Tracer substances leak between the collagen fiber bundles of the sclera and are absorbed by branches of the vortex veins (Fig. 13). Thirty percent of the aqueous humor normally leaves the
Fig. 10. Open communication leading from the anterior chamber to the supraciliary space created by a Q-switched, Nd:YAG laser (exposure duration: 35 ns; pulse energy: 300 mJ). The eye of a 45-year-old woman, which was enucleated three days after irradiation because of a malignant melanoma in the choroid. The broken line indicates the position of Schlemm’s canal. J: iris; T: trabecular meshwork; SEM micrograph; negative magnification ×100. (Reproduced from Van der Zypen and Fankhauser25 by courtesy of the publisher.)
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Fig. 11. Anterior chamber of the eye of a monkey (Macaca speciosa) 18 months after opening the supraciliary space (cyclocialysis) with a Q-switched Nd:YAG laser (exposure duration: 35 ns; pulse energy: 200 mJ). The newly formed canal has remained open. C: cornea; T: trabecular meshwork; J: iris; SEM micrograph; negative magnification ×100. (Reproduced from Van der Zypen and Fankhauser25 by courtesy of the publisher.)
Fig. 12. Same region as in Figure 11 at a greater magnification. An incomplete monocellular layer of connective tissue cells covers the outer wall of the opened supraciliary space. Connective tissue fibers (K) can be seen among the cell processes of the fibrocytes. SEM micrograph; negative magnification ×780. (Reproduced from Van der Zypen and Fankhauser25 by courtesy of the publisher.)
anterior chamber via a uveoscleral pathway.38,39 The second route leads to the veins of the ciliary muscle. Therefore, HgS particles can be found between the bundles of the ciliary muscle (Fig. 14). The third route is of some interest, but is speculative. HgS particles can be found in the supraciliary space, which is continuous with the suprachoroidal space up to the optic nerve head. The optic nerve outside the
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bulbus is surrounded by the three layers of the meninges (Figs. 15a and b). The subarachnoid space containing cerebrospinal fluid extends from the brain and passes around the optic nerve up to the ocular bulbus. Tracer substances are found in the subrachoroid space and the tissue of Elschnig which borders the optic nerve head (Figs. 15a and b). Tracer substances are also found in the subarachnoid space
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Fig. 13. HgS particles (P) in an intrascleral vein demonstrating an aqueous humor outflow route after cyclodialysis with a Qswitched Nd:YAG laser (exposure duration: 35 ns; pulse energy: 200 mJ). Transmission electron microscopic (TEM) micrograph; negative magnification ×3400.
Fig. 14. HgS particles (P) between the muscle bundles (M) of the ciliary muscle of monkey eye (Macaca speciosa), demonstrating an aqueous humor outflow route after cyclodialysis with a Q-switched Nd:YAG laser (exposure duration: 35 ns; pulse energy: 200 mJ). TEM micrograph; negative magnification ×2800. (Reproduced from Van der Zypen and Fankhauser16 by courtesy of the publisher.)
of the optic nerve (Fig. 16). Finally, tracer substances can be identified within the granulations of Pacchioni and pass from here into the superior sagittal sinus of the cranial dura mater. This pathway from the supraciliary via the suprachoroid into the subarachnoid space in enucleated human eyes can also be marked in the opposite direction. Tracer substances injected at high pressure into the subarachnoid space
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of the optic nerve of human eyes can also be found in the suprachoroid space. Morphological effects of trabeculoplasty Trabeculoplasty performed using a cw ALT, freerunning Nd:YAG laser light, or a diode laser shows
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a.
b. Figs. 15a and b. Human optic nerve and its sheaths. O: optic nerve; P: pia mater; A: arachnoid mater; D: dura mater; Ss: subarachnoid space; S: sclera; Od: optic disc; E: tissue of Elschnig; SEM micrograph; negative magnification: a. ×20; b. ×80.
similar morphological reactions.6-14 In clinical applications, when comparing argon and Nd:YAG lasers, the final reduction of IOP using either method was found to be identical.27 However, the effect of argon laser light is more superficial compared to Nd: YAG laser light. Irradiation with an argon laser (pulse power 0.5 W; pulse duration: 0.1 s; focus spot diameter: 50 µm; Fig. 17) or a free-running Nd:YAG laser (energy: 200 mJ; pulse duration; 10 ms; focus spot diameter: 50 µm; Fig. 18) produces superficial coagulation of the inner layers of the uveal trabecular meshwork.8,28 Some of the inner trabecular sheets are disrupted, and the solid cell processes of the trabecular endothelial cells broken up.28-30 Collagen fibrils of the trabecular cores disintegrate into the subfibrillar structures, loosing their periodicity (Fig. 19). In analogy, if collagen fibrils of the iris
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are heated up, it can be assumed that temperatures of up to 70-90°C will be reached within the center of impact during irradiation with the Nd:YAG laser (energy: 200 mJ; pulse duration: 10 ms).15,24 In deeper regions of the trabecular meshwork, the collagen fibrils within the trabecular cores were only partially dissolved; in many cases, they showed a normal cross-section, but had shortened periodicity (Fig. 20). Long-spacing fibers (curly collagen), which are often observed within the trabecular cores of glaucomatous human eyes, but more rarely in monkey eyes, degenerate. The light bands of these collagen fibrils disintegrate and exhibit tree-like ramifications (Fig. 21). Within the juxtacanalicular (cribriform) region of the trabecular meshwork, no conspicuous changes of the structures are observed immediately after ALT. When using the Nd:YAG
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Fig. 16. Upper end of the subarachnoid space in connection with the sclera. HgS particles (white particles, Hg) enter the subarachnoid space through spaces in the sclera (tissue of Elschnig). A: arachnoid; SEM micrograph; negative magnification ×2500. (Reproduced from Van der Zypen et al.35 by courtesy of the publisher.)
Fig. 17. Trabecular meshwork of a monkey (Macaca speciosa) four weeks after ALT (exposure duration: 0.1 s; pulse power: 560 mW). The large trabecular beams running from the iris root directly towards Schwalbe’s line lose their regular surface and appear to be broken in various places. J: iris; C: cornea; SEM micrograph; negative magnification ×90. (Reproduced from Van der Zypen and Fankhauser29 by courtesy of the publisher.)
laser for trabeculoplasty, only a small number of collagen fibrils dissolve. In long-term observations, most morphological investigations of the trabecular meshwork after ALT demonstrate the more or less complete closure of the inner trabecular spaces in the zone of impact, due to scar formation.31-33,51 This scar formation is thought to be the result of thermal injury to the trabecular meshwork. These findings are in agreement with our investigations.6,28,29,34,35 Four weeks after
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irradiation, activated cells, with morphological features similar to those of the corneal endothelium and connected to them, migrate from Schwalbe’s line and cover the inner iridocorneal trabeculae (Fig. 22). These cells close the irradiated zone by building a hexagonal monolayer which also expands into neighbouring untreated zoned, occluding the intertrabecular spaces. In monkey eyes, cells in the operculum were also stimulated to migrate into the damage zone.35 The operculum consists of a group of cells
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Fig. 18. Trabecular meshwork of a monkey (Macaca speciosa) one week after irradiation with a free-running Nd:YAG laser (exposure duration: 10 ms; pulse energy: 200 mJ). The uveal trabeculae at the impact spot are broken up and appear to be coagulated by heat. J: iris; C: cornea; SEM micrograph; negative magnification ×160. (Reproduced from Ticho et al.33 by courtesy of the publisher.)
Fig. 19. Core of a uveal trabecula of the trabecular meshwork of a monkey (Macaca speciosa) immediately after irradiation with a free-running Nd:YAG laser (exposure duration: 10 ms; pulse energy: 200 mJ). Collagen fibrils have disintegrated into subfibrillar structures, loosing their periodicity. This corresponds to a temperature of about 90°C being reached during irradiation. ETM micrograph; negative magnification ×18,000.
near the anterior insertion of the trabeculae, and is not seen in human eyes. It may be that vasoactive intestinal protein (VIP, 28-aminoacid neuropeptide) stimulates the proliferation of migrating cells.36 Both ALT and Nd:YAG laser trabeculoplasty are accepted clinical methods for lowering IOP.4,9-11,13 It has been speculated that shrinkage of the collagen arising at the impact zone causes dilatation and elevation of the intertrabecular spaces near to the scar.3
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In contrast to the scar formation at the inner trabecular meshwork, degeneration of cells and intercellular material can be seen within the juxtacanalicular region.28,34,35 In this zone, accumulation of basal lamina-like material with heparin sulfate-type proteoglycans is found in glaucomatous human eyes.19 Degeneration of the cells starts with lysis of the cell organelles, dilatation of the cell nucleus, and accumulation of DNS material. Cell membranes remain intact for a long time (Fig. 23).
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Fig. 20. Core of a corneoscleral trabecula of the trabecular meshwork of a monkey (Macaca speciosa) immediately after irradiation with a free-running Nd:YAG laser (exposure duration: 10 ms; pulse energy: 200 mJ). Collagen fibrils have partially disintegrated into the subfibrillar structures with a loss of periodicity. This corresponds to a temperature of about 60°C being reached during irradiation. TEM micrograph; negative magnification ×18,000.
Fig. 21. Trabecular meshwork of a monkey (Macaca speciosa) seven weeks after irradiation with a free-running Nd:YAG laser (exposure duration: 10 ms; pulse energy: 200 mJ). The long-spacing (curly) collagen within the trabecular cores has degenerated. The disintegrated light bands initially show irregularities, which later exhibit tree-like ramifications; dark bands also become broken up. TEM micrograph; negative magnification ×22,000.
At the end of the degeneration process (from 13 weeks to seven months after irradiation with the Nd:YAG and argon lasers, respectively), most of the trabecular endothelial cells at the juxtacanalicular region have degenerated near the impact zone. The final organization of the cellular network has come to an end after degeneration of the basement mem-
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brane, which can withstand the degenerative process for a long period of time (Fig. 24). Apart from this degenerative process, phagocytosed cells migrate into the juxtacanalicular region of the irradiated zone. These cells phagocytize and digest collagen fibrils (Fig. 25). The collagen fragments are pressed together within a voluminous
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Fig. 22. Trabecular meshwork of a monkey eye (Macaca speciosa) four weeks after irradiation with a free-running Nd:YAG laser (exposure duration: 10 ms; pulse energy: 200 mJ). A monolayer of corneal-like endothelial cells (E) covers the inner trabecular beams near the corneal insertion in the impact zone. Pigmented cells (P: opercular cells) can be seen at the surface of the network. SEM micrograph; negative magnification ×200. (Reproduced from Van der Zypen et al.35 by courtesy of the publisher.)
Fig. 23. Trabecular meshwork of a monkey eye (Macaca speciosa) four weeks after ALT (exposure duration: 0.2 s; pulse energy: 250 mW). Lytic degeneration of the trabecular endothelial cells. Vacuolation of mitochondria (M) and conglomeration of ribosomes (R) can be seen. TEM micrograph; negative magnification ×18,000.
vacuole (Fig. 26). The accumulation of acid phosphatase, which is an important enzyme for digestion is indicative of the degradation process within the vacuoles (Fig. 27).28,29,34 As well as this process of degradation, fibroblasts, which are active in collagen synthesis, migrate into the zone of dissolution.6,37 The activity of these fibroblasts in collagen synthesis is shown by the expansion of the rough endoplasmic reticulum (Fig. 28). In summary, an active process based on the turnover of cells and intercellular substances can be seen
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from 13 weeks to seven months after irradiation (i.e., the end of the period of observation). The most prominent feature is the degenerative processes. These may be responsible for the postoperative decrease of IOP. Later on, it may be that regenerative processes come into play that cause the re-elevation of the IOP. Remodelling of the juxtacanalicular extracellular matrix may be initiated by the expression of cytokines in the juxtacanalicular region. Such cytokines have been found in organ cultures of the trabecular meshwork after ALT.22
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Fig. 24. Juxtacanalicular tissue of a monkey eye (Macaca speciosa) 13 weeks after irradiation with an Nd:YAG laser (exposure duration: 10 ms; pulse energy: 600 mJ). In some places, most of the trabecular cells have degenerated (E) and the collagen fibrils have dissolved. The folded basement membrane (Bl) approximately indicates the former organization of the network. TEM micrograph; negative magnification ×8000. (Reproduced from Van der Zypen et al.35 by courtesy of the publisher.)
Fig. 25. Juxtacanalicular tissue of a monkey eye (Macaca speciosa) 13 weeks after irradiation with an Nd:YAG laser (exposure duration: 10 ms; pulse energy: 800 mJ). Collagen-phagocytosing pigmented cells. Collagen fragments are phagocytosed and collect into larger vacuoles. TEM micrograph; negative magnification ×14,000.
Conclusions and perspectives in glaucoma surgery ALT would seem to be the most important method for lowering IOP, with the exception of perforating methods such as Elliot-type procedures. Scar formation may prevent the long-term success of ALT in lowering IOP. More recently, alternative methods have been advanced.
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Viscocanalostomy has been reported to lower IOP without creating a filtering bleb.40 With this technique, following removal the deep scleral layers, Schlemm’s canal is cannulated and is expanded with a viscoelastic material. Complications such as hypotony and hyphema, which may compromise the operative results, may be avoided.40,41 Variations of this procedure include removing the inner wall of
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Fig. 26. Juxtacanalicular tissue of the trabecular meshwork of a monkey eye (Macaca speciosa) eight weeks following ALT (pulse power: 580 mW; pulse duration: 0.2 s). Fragments of collagen fibrils are fused together within a large intracellular vacuole. The degeneration process is indicated by the intravacuolar accumulation of acid phosphatase enzyme (dark granules). TEM micrograph; negative magnification ×28,000.
Fig. 27. Juxtacanalicular tissue of the trabecular meshwork of a monkey eye (Macaca speciosa) eight weeks following ALT (exposure duration: 0.2 s; pulse power: 580 mW). Advanced (progressed) stage of collagen digestion within a collagen phagocytosing cell. Conglomeration of acid phosphatase enzyme, indicated by black granules within a large vacuole, can be seen. The collagen fragments have completely disintegrated. TEM micrograph; negative magnification ×28,000.
Schlemm’s canal and the adjacent meshwork, while leaving the inner meshwork intact and placing a collagen implant into the filtration bed in order to prevent episcleral fibrosis.42 A more likely explanation of the working mechanisms of this operative procedure is the expansion of the canal ruptures and the disruption of both the inner and outer endothelial walls of Schlemm’s canal.43 This disruption may extend into the juxtacanalicular tissue and may also rupture some of the uveal meshwork. Another surgical method for lowering IOP is the
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manipulation of Descemet’s membrane, which is not normally permeable to the aqueous humor. However, if Descemet’s membrane has become thin or is partially removed, its permeability may increase.44 Laser sclerostomy ab interno may also be a successful approach in the treatment of open-angle glaucoma.45 The failure of this and of other filtering surgery methods may be caused by the extensive wound healing process.46 However, scar formation can be suppressed by the use of antifibrotic agents such as 5-fluorouracil or mitomycin C.47-50 Mitomy-
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Fig. 28. Juxtacanalicular tissue of the trabecular meshwork of a monkey eye (Macaca speciosa) four weeks after irradiation with a free-running Nd:YAG laser (exposure duration: 10 ms; pulse energy: 200 mJ). Activated fibroblast. Collagen synthesis activity is indicated by a well-developed, dilated, rough endoplasmic reticulum containing a fine granular amorphous mass. TEM micrograph; negative magnification ×22,000.
cin C was noted to suppress the migration and proliferation of fibroblasts and macrophages for more than ten weeks after Nd:YAG laser sclerostomy ab interno in rabbits. Repolymerization of heat-damaged collagen was also unsuccessful in these experiments, and no neosynthesis of collagen fibrils was observed.50 Selective trabeculoplasty10 is a novel method using a Q-switched, frequency-doubled (KTP) laser, which it is claimed drastically reduces radiation damage to the trabeculum, while maintaining equally good results as standard ALT.10
9.
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References 13. 1. Krasnov MM: Laseropuncture of the anterior chamber angle in glaucoma. Am J Ophthalmol 75:674-678, 1973 2. Worthen DM, Wickham MG: Laser trabeculotomy in monkeys. Invest Ophthalmology Vis Sci 12:707-711, 1973 3. Wise JB, Witter SL: Argon laser therapy for open-angle glaucoma. Arch Ophthalmology 97:319-322, 1979 4. Reiss GR, Wilensky JT, Higginbotham EJ: Laser trabeculoplasty. Surv Ophthalmol 35:407-428, 1991 5. Brilakis HS, Hann CR, Johnson DH: A comparison of different embedding media on the ultrastructure of the trabecular meshwork. Curr Eye Res 22:235-244, 2001 6. Fankhauser F, Van der Zypen E, Kwasniewska S: Argon and Nd:YAG laser trabeculoplasty: the relevance of ultrastructural findings for the evaluation of therapeutic effectiveness. New Trends Ophthalmol 2:238-245, 1987 7. McMillan TA, Stewart WC, Legler UF, Powers T, Nutaitis MJ, Apple DJ: Comparison of diode and argon laser trabeculoplasty in cadaver eyes. Invest Ophthalmol Vis Sci 35:706-710, 1994 8. Hollo G: Argon and low energy, pulsed Nd:YAG laser trabeculoplasty: a prospective, comparative clinical and
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morphological study. Acta Ophthalmol Scand 74:126-131, 1996 Englert JA, Cox TA, Allingham RR, Shields MB: Argon vs diode laser trabeculoplasty. Am J Ophthalmol 124:627631, 1997 Latina MA, Sibayan SA, Shin DH, Noecker RJ, Marcellino G: Q-switched 532 nm Nd:YAG laser trabeculoplasty (selective laser trabeculoplasty): a multicenter, pilot, clinical study. Ophthalmology 105:2082-2088, 1998 Wang RF, Schumer RA, Serle JB, Podos SM: A comparison of argon laser and diode laser photocoagulation of the trabecular meshwork to produce the glaucoma monkey model. J Glaucoma 7:45-49, 1998 Kim YJ, Moon CS: One-year follow-up of laser trabeculoplasty using Q-switched frequency-doubled Nd:YAG laser of 523 nm wavelength. Ophthalmic Surgery Lasers 31:394-399, 2000 Blyth CPJ, Moriarty AP, McHugh JDA: Diode laser trabeculoplasty versus argon laser trabeculoplasty in the control primary open angle glaucoma. Lasers Med Sci 14: 105108, 1999 Dietlein TS, Jacobi PC, Mietz H, Kriegelstein GK: Morphology of the trabecular meshwork three years after erbium:YAG trabecular ablation. Ophthalmic Surg Lasers 32:483-485, 2001 Van der Zypen E, Fankhauser F, Bebie H, Marshall J: Changes in the ultrastructure of the iris after irradiation with intense light: a study of long-term effects after irradiation with argon-ion, Nd:YAG and Q-switched ruby lasers. Adv Ophthalmol 39:59-180, 1979 Van der Zypen E, Fankhauser F: Lasers in the treatment of chronic simple glaucoma. Trans Ophthalmol Soc UK 102: 147-153, 1982 McHugh D, Marshall J, Ffytche TJ, Hamilton A, Raven A: Ultrastructural changes of the human trabecular meshwork after photocoagulation with a diode laser. Invest Ophthalmol 33:2664-2667, 1992 Furuyoshi N, Furuyoshi M, Futa R, Gottanka J, LutjenDrecoll E: Ultrastructural changes in the trabecular mesh-
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S: Morphology of the trabecular meshwork within monkey (Macacca speciosa) eyes after irradiation with the freerunning Nd:YAG laser. Ophthalmology 94:171-179, 1987 Koh SW, Yeh TH, Morris SM, Leffler M, Higginbotham EJ, Brenneman DE, Yue BY: Vasoactive intestinal peptide stimulation of human trabecular meshwork cell growth. Invest Ophthalmol 38:2781-2789, 1997 Van Buskirk EM: Pathophysiology of laser trabeculoplasty. Surv Ophthalmol 33:264-272, 1989 Bill A, Walinder E: The effect of pilocarpine on the dynamics of aqueous humour in a primate (Macaca irus). Invest Ophthalmol 5:170-175, 1966 Nilsson SF: The uveal outflow routes. Eye 11:149-154, 1997 Stegmann R, Pienaar A, Miller D: Viscocanalostomy for open-angle glaucoma in black African patients. J Cataract Refract Sur 25:316-322, 1999 Sayyad FE, Helal M, El-Kohilty H, Knall M, El-Mmaghraby A: Nonpenetrating deep sclerectomy vs trabeculectomy in bilateral primary open-angle glaucoma. Ophthalmology 107:1671-1674, 2000 Mernoud A, Schnyder CC, Sickenberg M, Chiou AGY, Hediger SEA, Faggioni R: Comparison of deep sclerectomy with collagen implant and trabeculectomy in open-angle glaucoma. Ophthalmic Surg 15:734-740, 1984 Smit BA, Johnstone MA: Effect of viscocanalostomy on the histology of Schlemm’s canal in primate eyes. (Abstract 3071) Invest Ophthalmol Vis Sci 41(Suppl):5578, 2000 Splegel D, Schefthaler M, Kobuch K: Outflow facilities through Descemet’s membrane in rabbits. (Abstract 3071) Invest Ophthalmol Vis Sci 41(Suppl):5578, 2000 Fankhauser F, Dürr U, England C, Kwasniewska S, Van der Zypen E, Henchoz PD: Optical principles related to optimizing sclerostomy procedure. Ophthalmic Surg 23:752761, 1992 Iliev ME, Van der Zypen E, Fankhauser F, England C: The repair response following Nd:YAG laser sclerostomy ab interno in rabbits. Exp Eye Res 61:311-321, 1995 Yamamoto T, Varani J, Soong HK, Lichter PR: Effects of 5-fluorouracil and mitomycin C on cultured rabbit subconjunctival fibroblasts. Ophthalmology 97:1204-1210, 1990 Wand M: Minimizing conjunctival wound leaks in filtration surgery with mitomycin C. Ophthalmic Surg 24:708709, 1993 Pablo LE, Ramirez T, Alvarez R, Gonzalez I, Larrosa JM, Honrubia FM: Morphometric study of wound healing in a model of filtering surgery with mitomycin C. Eur J Ophthalmol 5:168-171, 1995 Iliev ME, Van der Zypen, E, Fankhauser F, England C: Transconjunctival application of mitomycin C in combination with laser sclerostomy ab interno: a long-term morphological study of the postoperative healing process. Exp Eye Res 64:1013-1026, 1997 Koller T, Sturmer J, Remé C, Gloor B: Membrane formation in the chamber angle after failure of argon laser trabeculoplasty: analysis and risk factors. Br J Ophthalmol 84:48-53, 2000 Van der Zypen E, Fankhauser F: Effekte eines neuartigen Lasertyps auf fixiertes Gewebe der Kammerwinkelregion des Affenauges. Klin Mbl Augenheilk 172:426, 1978 Van der Zypen E, Fankhauser F: Lasereffekte am Trabekelwerk: Ultrastrukturelle Untersuchungen am menschlichen und am Affenauge. In: Naumann GOH, Gloor B (ed) Wundheilung des Auges und ihre Komplikationen, pp 331-336. Munich: JF Bergmann Verlag 1880
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Erbium:YAG laser trabecular ablation Thomas S. Dietlein and Günter K. Krieglstein Department of Ophthalmology, University of Cologne, Cologne, Germany
Keywords: glaucoma, laser surgery, trabecular meshwork, intraocular pressure
Background Numerous morphological and functional investigations on the human trabecular meshwork have produced evidence to support the notion that the increased aqueous outflow resistance in primary open-angle glaucoma is caused by specific changes within the inner wall of Schlemm’s canal and the adjacent juxtacanalicular trabecular meshwork.1-3 By performing selective trabecular surgery, the site of surgical intervention can be limited to the original ‘locus causae’ and postoperative over-filtration with subsequent problems such as choroidal detachment and anterior chamber loss can thus be avoided. A further conceptual advantage of ab interno trabecular surgery over conventional filtration surgery4 is the untouched conjunctiva, allowing future ab externo interventions in case of failure. Theoretical calculations suggested that just a few 10-20 µm openings within the trabecular meshwork would be sufficient to re-establish a normal outflow facility in glaucomatous eyes.5 Simple incision of the trabecular meshwork was described back in the 19th century,6 but another 50 years were to pass before Otto Barkan developed the well-known technique of goniotomy, based on this approach.7 Today, goniotomy is performed with success in congenital and developmental glaucomas. Using the ab externo approach with the disadvantage of conjunctival lesions, trabeculotomy also selectively targets the trabecular meshwork by rupturing the meshwork rather than incising it.8,9 In neither procedure is the trabecular meshwork removed; it is only ruptured or incised. Histopathological studies in animals have proven that the tissue cleft after trabeculotomy tends to be reclosed
by secondary repair mechanisms.10 Ideal trabecular surgery should therefore actually remove the diseased tissue and guarantee a broad direct opening into the lumen of Schlemm’s canal. Different surgical methods to remove trabecular tissue mechanically by an ab interno approach are currently under investigation.11,12 These manual techniques require great surgical precision; the reproducibility of the mechanical effects is totally dependent on the skill of the individual surgeon. Laser surgery of the trabecular meshwork could offer the advantage of better reproducibility, owing to fixed energy parameters. Krasnov was one of the first to try to perforate the trabecular meshwork with the aid of a laser.13,14 The collateral thermal damage with these early laser systems was extensive, limiting the hope of achieving long-term effectiveness with these procedures. The use of the Nd:YAG laser via a gonioscope presented a means of selectively treating the trabecular meshwork without penetrating the eye, although this photodisruptive laser does not actually remove trabecular tissue, but only ruptures the meshwork. By perforating the trabecular tissue in several places, it was possible to lower the intraocular pressure (IOP) and to produce increased outflow facility. Despite promising results in the early postoperative phase, IOP normalization in the treated eyes and ’tissue-filling-in’ of laser-induced openings within the trabecular meshwork was noted after a few weeks.15 Clinical results were more promising in juvenile glaucoma than in adult open-angle glaucoma.16,17 Today, Nd:YAG laser goniopuncture is experiencing a revival in the postoperative management of non-penetrating glaucoma surgery. In this type of surgery, a deep scleral
Address for correspondence: Priv-Doz Dr Thomas S. Dietlein, Klinik und Poliklinik für Augenheilkunde, Universität Köln, JosephStelzmann-Strasse 9, D-50931 Köln, Germany. e-mail:
[email protected] Lasers in Ophthalmology – Basic, Diagnostic and Surgical Aspects, pp. 333–339 edited by F. Fankhauser and S. Kwasniewska © 2003 Kugler Publications, The Hague, The Netherlands
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flap is dissected and removed while the so-called trabeculo-descemetic membrane remains intact. If the IOP increases again postoperatively, this trabeculo-descemetic membrane can be perforated by Nd:YAG goniopuncture, with relatively good results.18 Erbium:YAG laser trabecular ablation Pre-clinical studies In contrast to the photodisruptive Nd:YAG laser, the Erbium:YAG (Er:YAG) laser can potentially evaporate the targeted tissue and achieve a broader opening to Schlemm’s canal. The infrared Er:YAG laser with a wavelength at 2.94 µm, corresponding to an absorption peak of water, had already been proposed for laser sclerostomy, since reduction of collateral tissue necrosis raised hope of a limited repair response after photoablative filtrating glaucoma surgery. However, being an unguarded filtration procedure, laser sclerostomy could never ultimately succeed, regardless of the wavelength used.19-21 Hill et al.22 were the first to perform photoablation of the trabecular meshwork by applying the Er:YAG laser in direct tissue contact with the trabecular meshwork of donor eyes. In a consecutive animal study, Hill et al.23 demonstrated that Er:YAG laser trabecular ablation (LTA) was capable of increasing the outflow facility and lowering the IOP for at least a number of weeks following surgical intervention. They observed a filling-in of the laser-induced craters in the trabecular meshwork with tissue when the IOP and tonographic curve had normalized again.22,23 The Er:YAG laser (MCL 29, Fa. Aesculap Meditec, Heroldsberg, Germany) that we initially used for LTA acted in the spiking mode of a 200-µsec macropulse. The laser energy was delivered through an articulated system of a zirconium fluoride fiber (350 µm diameter) and an exchangeable quartz fiber endoprobe. The quartz fiber endoprobe for LTA is longer than the conventional tip for laser sclerostomy, because the instrument is applied transcamerally in direct contact with the trabecular meshwork. In an experimental in-vitro study, we investigated the changes of outflow facility24 in the anterior segment cups of porcine cadaver eyes following LTA and in the anterior segment cups of untreated control eyes. While the mean outflow facility in the control eyes was 0.128+0.041 µl/min × mmHg, the application of 20 pulses (4 mJ pulse energy) to an 90-120° area of the trabecular meshwork yielded an increased outflow facility of 0.308 +0.093 µl/min × mmHg.25 In a similar study with human donor eyes, McHam and co-workers used sapphire tips for Er:YAG LTA.26,27 Their laser system (Candela Laser Corporation, Wayland, MA)
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worked with pulse energies of between 10 and 20 mJ, a pulse length of 250 µsec and an endoprobe core diameter of 300 µm. They also observed a significant increase in outflow facility of from 0.283+0.08 µl/min × mmHg to 0.62+0.15 µl/min × mmHg with Er:YAG LTA. Both studies underlined the functional efficacy and clinical potential of LTA. In another experimental in-vitro study, we investigated the photoablative profile of different quartz fiber endoprobes in human donor eyes.28 Using the Er:YAG laser, the pulse energy can be delivered by an end- or a side-firing tip and the core diameter of the endoprobe can also be varied. Full-thickness Er:YAG photoablation of the trabecular meshwork succeeded with each of the different fiber tips used (Figs 1-3). Opening of Schlemm’s canal could be reproducibly achieved at a specific pulse energy characteristic for each endoprobe. Laser-induced apertures in the trabecular meshwork revealed an elliptic inner surface when using a side-firing bevelled tip.29 Already in this in-vitro study, two major problem became obvious: firstly, there was the mechanical fragility of the quartz fiber tips after frequent use. Microscopic examination of aged quartz fiber endoprobes revealed mechanical damage such as splitting and cracking of the end of the tip. Whether sapphire tips are less fragile and hence more reliable in clinical practice still has to be established. Secondly, the manual pressure along the fiber axis has a critical influence on ablation depth. As already shown by Fankhauser et al., in laser sclerostomy, total energy can be massively reduced if the surgeon increases the mechanical pressure on the fiber tip.30 This ‘human’ factor – together with the fragility of the quartz fiber tips – introduces a degree of unpredictability and variability to the intended full-thickness ablation of the trabecular meshwork. According to laser physics, the larger endoprobes (> 300 µm core diameter) require higher single pulse energies for opening of Schlemm’s canal and complete removal of the trabecular meshwork, but the collateral tissue damage of around 30 µm was relatively similar for all endoprobes at the optimal energy level.31,32 In order to investigate the safety and scarring potential of this procedure, we performed Er:YAG LTA in rabbits as part of a preclinical study.32 Twenty healthy pigmented rabbits underwent surgery of the right eye, while their left eye served as a control. IOP measurements and anterior chamber inspection were performed at intervals, and the animals were sacrificed on Days 2, 10, 30, and 60 following surgery in order to document the morphological changes induced by LTA. Between ten and 15 ablation craters (2 mJ single pulse energy) were created with a 200-µm-core-diameter quartz fiber endoprobe under gonioscopic control. Although this procedure is relatively difficult in rabbits, light-microscopic morphology revealed
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Fig. 1. Scanning electron microscopy of the trabecular meshwork after Er:YAG LTA in a human donor eye. (Reproduced from Dietlein et al.28 by courtesy of the publisher.)
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Fig. 4. Light microscopy of the anterior chamber angle after Er:YAG LTA in rabbits: already ten days after surgery, massive invasion of fibroblasts into the craters has taken place; no collateral thermal damage is visible any longer.
damage was already undetectable ten days after LTA (Fig. 4), although massive secondary repair mechanisms continued for a number of weeks. After two months, dense occlusion of the laserinduced craters and even anterior synechiae in the chamber angle were seen. IOP in these rabbits slightly decreased from 9.1 mmHg preoperatively to 7.2 mmHg after two months, but was not significantly lower than that of the untreated control eyes (8.5 and 8.6 mmHg). This rabbit study appeared to confirm Hill’s experience of early tissue refilling after LTA. Clinical studies
Fig. 2. Light microscopy of Er:YAG LTA (see arrow) in a human donor eye using a 320-µm-core-diameter quartz fiber endoprobe at 6 mJ single pulse energy. The outer wall of Schlemm’s canal shows only minimal thermal damage. (Reproduced from Dietlein et al.28 by courtesy of the publisher.)
Fig. 3. Different quartz fiber tips after frequent use. The ends of the tips reveal microscopic lesions, such as splitting and cracking. (Reproduced from Dietlein et al.28 by courtesy of the publisher.)
successful photoablation of the trabecular meshwork on Day 2 with blood-filled craters in the meshwork. As in a previous animal study of laser sclerostomy in rabbits,33,34 collateral thermal tissue
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Although animal studies had indicated some limitations with LTA in vivo with regard to wound healing, we started a pilot study in patients with advanced open-angle glaucoma in order to evaluate the potential of Er:YAG LTA for glaucoma management. We were encouraged by the differences between the secondary tissue repair observed in young, healthy animals and that seen in elderly glaucoma patients. These are well-documented in trabeculectomy and laser sclerostomy studies.33,34 In one pilot study, 11 patients with open-angle glaucoma were treated by LTA. The mean preoperative maximum IOP was 36.0+8.8 mmHg. The procedure was performed using retro- or parabulbar anesthesia. After injection of high-viscosity viscoelastics into the anterior chamber via a small temporal corneal incision, the quartz endoprobe was transcamerally inserted up to the opposite trabecular meshwork (Fig. 5). Pulse energy was delivered via an end-firing quartz fiber endoprobe (core diameter 320 µm, coating diameter 385 µm). The pulse energy emitted at the fiber tip ranged between 5 and 7 mJ, and was preoperatively checked by an external joulemeter. Between ten and 30 neighbouring single laser pulses were applied under gonioscopic guidance, while exerting gentle pressure along the fiber tip, which was in direct contact with the trabecular meshwork. After
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Fig. 6. Gonioscopic view two months after Er:YAG LTA in a glaucoma patient: the ablation craters can be clearly identified within the trabecular meshwork by their lack of pigmentation.
Fig. 5. Gonioscopic view during Er:YAG LTA in a glaucoma patient: the endoprobe has been introduced into the viscoelasticfilled anterior chamber via a temporal corneal incision. The end of the tip has to be in direct contact with the opposite trabecular meshwork before laser energy will be emitted.
the procedure, the viscoelastics, blood, and debris were rinsed out by irrigation-aspiration maneuvers. The mean number of applied laser pulses in this first study was 18+5.5. The mean maximum IOP at the end of a one-year follow-up was 22.0+7.2 mmHg, representing a percentage decrease from baseline of 38.9% (p = 0.008). The average medication score per eye dropped from 2.82 to 1.46 at the end of the follow-up period, representing a reduction of 48% (p = 0.021). At the end of the follow-up period, seven of 11 eyes (64%) constantly maintained their IOP below 21 mmHg. In these successfully treated patients, the mean maximum IOP was 17.6+1.6 mmHg and the mean number of topical medications used 1.43+0.54. No serious intra- or postoperative complications occurred in the treated eyes. The laser-induced craters were gonioscopically visible, especially in the pigmented trabecular meshwork. However, visibility slightly decreased during follow-up (Figs 6-7), but without any correlation with the IOP level.35 In recent studies, Funk et al.36,37 investigated the clinical efficacy of Er:YAG LTA in combination with modern cataract surgery in patients with openangle glaucoma. After performing phacoemulsification with implantation of a foldable intraocular lens, they used an endoscope with an integrated Er:YAG laser. Owing to the larger core diameter of the endoprobe used, 15-18 single pulses of around 14 mJ each had to be applied. In 21 eyes, the mean preoperative IOP of 23.4 mmHg could be reduced to 16.4 mmHg more than one year
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Fig. 7. Gonioscopic view three years after Er:YAG LTA in a glaucoma patient: there are only a few irregularities of pigmentation in the area of laser trabecular ablation, but no clearly identifiable laser craters. IOP is still under control.
after Er:YAG LTA. Topical medication could also be significantly reduced. In terms of pressure reduction, combined Phako-LTA was as effective as combined phacoemulsification-trabeculectomy (Phako-TE) in glaucoma patients, but brought significantly fewer postoperative complications, required less re-surgery, and achieved significantly faster postoperative rehabilitation.38 Conclusively, it should be taken into account that the repair mechanisms in the relatively old glaucoma patients are very slow and incomplete in comparison to those observed in animal models. Moreover, the new scar tissue replacing the glaucomatous trabecular tissue in the treatment area could also contribute to the improved outflow facility after surgery, and newly formed vessels have indeed been found within the scar area after pressurereducing laser treatment in animals. In one of our patients, who received LTA for silicone-oil glaucoma in his blind eye, enucleation was performed three years following LTA and histological investigation of the treatment area could be performed.39 Morphological proof of laser craters without significant scarification three years following Er:YAG LTA also demonstrated the massive differences in healing processes between young healthy animals and elderly glaucoma patients. Even in treatment areas where the laser-induced craters were localized adjacent to
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second intervention was necessary to clear the blood from the anterior chamber. In another patient, transient opacity of the corneal periphery near the treatment area was described, but without disturbance of visual function. No other essential complications were observed.42 Neuhann et al. also used the excimer laser for ab interno trabecular ablation in 14 patients. After six months, the mean IOP was 15.5 mmHg compared to 26.8 mmHg before surgery.43 The surgical concept is identical to that of Er: YAG LTA. Light microscopy of the laser-induced craters after excimer ablation of the trabecular tissue in human donor eyes, reveals a morphology similar to that produced by the Er:YAG laser (Fig. 8). The collateral thermal damage seems to be a little less, but transmission variability and loss after frequent use also occur with the excimer laser and may lead to insufficient ablation of the trabecular meshwork.
Excimer laser trabecular ablation
Intracanalicular trabeculostomy
Clinical studies on trabecular ablation have also been reported for the excimer laser with a wavelength of 308 nm.41 Vogel and co-workers42 performed punctual excimer laser photoablation of the trabecular meshwork in 27 patients with openangle glaucoma and in seven patients with normal-pressure glaucoma. They achieved a mean IOP decrease of 7 mmHg by creating three to six laserinduced perforations of the meshwork. The energy fluence applied ranged between 3.5 and 5.5 J/cm². The most frequent postoperative complications in this study were temporary pressure spikes produced by retained intraocular viscoelastics in 16 patients. Reflux bleeding was only observed in ten of the treated patients during surgery. In one patient, anterior chamber bleeding was so massive that a
Kampmeier et al. proposed a new concept for photoablation of the trabecular meshwork with the Er: YAG laser, using an ab externo approach similar to conventional trabeculotomy. After dissecting a scleral flap and searching for Schlemm’s canal, a side-firing laser endoprobe is introduced into the lumen of Schlemm’s canal and the meshwork can be photovaporized. In-vitro studies have been performed and clinical studies are in progress.44 The main conceptual problem of this procedure is the ab externo approach, which causes scarification on the conjunctival side. Another point of concern is the difficulty of introducing a rigid instrument into Schlemm’s canal without damaging surrounding structures such as the drainage vessels.
Fig. 8. Light microscopy after excimer laser trabecular puncture revealing a broad opening of Schlemm’s canal with obvious thermal damage to the outer wall. The small insert shows a view of the craters on the isolated trabecular meshwork strip, as seen under the operating microscope after in-vitro excimer laser trabeculopuncture.
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Conclusions At the present time, Er:YAG LTA represents a promising ab interno approach for the treatment of open-angle glaucoma. The major conceptual advantages of this procedure are the lack of conjunctival lesions and of filtration-related problems (e.g., overfiltration, early and late bleb leakage, etc.). The main problems are the relatively expensive and complicated equipment required, difficulties in visualization, and the lack of a means of monitoring intraoperative success. Reflux bleeding during LTA may be an indicator for opening Schlemm’s canal, but it is far from reliable as a sign that sufficient ablation of the meshwork has been achieved for the opening of Schlemm’s canal. Improvements in intraocular endoscopy may help to facilitate visualization of the chamber angle, making the intervention easier than surgery under the gonioscope.45-47 The ultrasound biomicroscope could provide another viable means of controlling the intraoperative position of the fiber tip end. Even more promising is the development of ultrashort-pulsed lasers,48 including the femtosecond lasers,49 which, in the future, may allow ablation of the trabecular tissue without direct tissue contact, thus avoiding all the risks of intraocular surgery. References 1. Bill A, Maepea O, Hamanaka T: Aspekte der Kammerwasserdrainage Über den Schlemmschen Kanal. Klin Mbl Augenheilk 195:277-280, 1989 2. Grant WM: Further studies on facility of flow through the trabecular meshwork. Arch Ophthalmol 60:523-533, 1958 3. Grant WM: Experimental aqueous perfusion in enucleated human eyes. Arch Ophthalmol 69:783-801, 1963 4. Cairns JE: Trabeculectomy: preliminary report of a new method. Am J Ophthalmol 66:673-679, 1968 5. Goldschmidt CR, Ticho U: Theoretical approach to laser trabeculotomy. Med Phys 5:92-98, 1978 6. DeVincentiis C: Incisione dell angolo irideo nel glaucoma. Ann Ottalmol 22:540-542, 1893 7. Barkan O: Operation for congenital glaucoma. Am J Ophthalmol 25:552-568, 1942 8. Grehn F: The value of trabeculotomy in glaucoma surgery. Curr Opin Ophthalmol 6:52-60, 1995 9. Harms H, Dannheim R: Trabeculotomy: results and problems. Adv Ophthalmol 22:121-131, 1970 10. Ito S, Nishikawa M, Tokura T et al: Histopathological study of trabecular meshwork after trabeculotomy in monkeys. Nippon Ganka Gakkai Zasshi 98:811-819, 1994 11. Jacobi PC, Dietlein TS, Krieglstein GK: Goniocurettage for removing trabecular meshwork: clinical results of a new surgical technique in advanced chronic open-angle glaucoma. Am J Ophthalmol 127:505-510, 1999 12. Quaranta L, Hitchings RA, Quaranta CA: Ab-interno goniotrabeculotomy versus mitomycin C trabeculectomy for adult open-angle glaucoma: a 2-year randomized clinical trial. Ophthalmology 106:1357-1362, 1999 13. Krasnov MM: Laser puncture of anterior chamber angle in glaucoma. Am J Ophthalmol 75:674-678, 1973 14. Krasnov MM: Q-switched laser goniopuncture. Arch Ophthalmol 92:37-41, 1974
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15. Melamed S, Pei J, Puliafito CA, Epstein DL: Q-switched neodymium-YAG laser trabeculopuncture in monkeys. Arch Ophthalmol 103:129-133, 1985 16. Epstein DL, Melamed S, Puliafito CA, Steinert RF: Neodymium:YAG laser trabeculopuncture in open-angle glaucoma. Ophthalmology 92:931-937, 1985 17. Melamed S, Latina MA, Epstein DL: Neodymium:YAG laser trabeculopuncture in juvenile open-angle glaucoma. Ophthalmology 94:163-170, 1987 18. Mermoud A, Karlen ME, Schnyder CC, Sickenberg M, Chiou AG, Hediguer SE, Sanchez E: Nd:YAG goniopuncture after deep sclerectomy with collagen implant. Ophthalmic Surg Lasers 30:120-125, 1999 19. Berlin MS, Yoo PH, Roy JH, Ahn BA: The role of laser sclerostomy in glaucoma surgery. Curr Opin Ophthalmol 6:102-114, 1995 20. Jacobi PC, Dietlein TS, Krieglstein GK: Prospective study of ab externo Erbium:YAG laser sclerostomy in humans. Am J Ophthalmol 123:478-486, 1997 21. Kampmeier J, Klafke M, Hibst R, Wierschien S: Ab externo Sklerostomie mit dem Er:YAG Laser: Ergebnisbericht nach 2 Jahren. Klin Mbl Augenheilk 208:218-223, 1996 22. Hill RA, Baerveldt G, Ozler SA, Pickford M, Profeta GA, Berns MW: Laser trabecular ablation (LTA). Lasers Surg Med 11:341-346, 1991 23. Hill RA, Stern D, Lesiecki ML, Hsia J, Liaw LH, Baerveldt G, Berns MW: Primate experience with erbium (Er):YAG laser trabecular ablation. ARVO abstracts. Invest Ophthalmol Vis Sci 33(Suppl):1018, 1992 24. Becker B, Constant MA: Species variation in facility of aqueous outflow. Am J Ophtalmol 42:189-194, 1956 25. Jacobi PC, Dietlein TS, Krieglstein GK: Effects of Er:YAG laser trabecular ablation on outflow facility in cadaver porcine eyes. Graefe’s Arch Clin Exp Ophthalmol 234:S204208, 1996 26. Fitzgibbon JJ, Bates HE, Pryshlak AP, Phickbrick MJ: Sapphire optical fibers for the delivery of Erbium:YAG laser energy. Proc Biomed Opt Soc SPIE 2131:50-55, 1994 27. McHam ML, Eisenberg DL, Schuman JS, Wang N: Erbium: YAG laser trabecular ablation with a sapphire optical fiber. Exp Eye Res 65:151-155, 1997 28. Dietlein TS, Jacobi PC, Krieglstein GK: Erbium:YAG laser ablation on human trabecular meshwork by contact delivery endoprobes. Ophthalmic Surg Lasers 27:939-945, 1996 29. Melnik I, Krivokhizha A, Ptashnik O: Multipurpose fiber optic sensors with slope tips. SPIE 1572:118-122, 1991 30. Fankhauser F, Dürr U, England C, Kwasniewska S, Van der Zypen E, Henchoz PD: Optical principles related to optimizing sclerostomy procedures. Ophthalmic Surg 23:752-761, 1992 31. Kramer TR, Noecker RJ, Ellsworth LG, Yarborough JM: Laser trabecular ablation of human eyes with the Erbium: YAG laser: a histopathologic study. SPIE (Ophthalm Technologies IV) 2126:242-250, 1994 32. Dietlein TS, Jacobi PC, Schröder R, Krieglstein GK: Experimental Erbium:YAG laser photoablation of trabecular meshwork in rabbits: an in-vivo study. Exp Eye Res 64:701706, 1997 33. Iliev M, Van der Zypen E, Fankhauser F, England C: The repair response following Nd:YAG laser sclerostomy ab interno in rabbits. Exp Eye Res 61:311-321, 1995 34. Iliev M, Van der Zypen E, Fankhauser F: Vernarbung von Laser-erzeugten Sklerostomie-Kanälen beim Kaninchen. Klin Mbl Augenheilk 206:376-379, 1995 35. Dietlein TS, Jacobi PC, Krieglstein GK: Erbium:YAG laser trabecular ablation (LTA) in the surgical treatment of glaucoma. Lasers Surg Med 23:104-110, 1998 36. Funk J, Schlunck G: Endoscopically controlled Erbium:YAG
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laser goniotomy: initial preclinical trials. Ophthalmologe 95:33-36, 1998 Funk J, Feltgen N, Asbeck D: Augendrucksenkung durch endoskopisch kontrollierte Erb:YAG Goniotomie. Ophthalmologe (Suppl):S27, 1999 Feltgen N, Ott B, Frenz M, Funk J: Endoskopisch kontrollierte Erbium:YAG Goniotomie versus Trabekulektomie: Drucksenkung bei der Kombination mit einer Kataraktoperation. Ophthalmologe (Suppl 1):S33, 2001 Dietlein TS, Jacobi PC, Krieglstein GK: Morphology of the trabecular meshwork three years after laser trabecular ablation (LTA). Ophthalmic Surg Lasers 32:483-485, 2001 Alexander RA, Grierson I, Church WH: The effect of argon laser trabeculoplasty upon the normal human trabecular meshwork. Graefe’s Arch Clin Exp Ophthalmol 227:7277, 1989 Vogel M, Scheurer G, Neu W, Dressel M, Gerhardt H.:Die Ablation des Trabekelwerks. Klin Mbl Augenheilk 197:250253, 1990 Vogel M, Lauritzen K, Quentin CD: Punktuelle Ablation des Trabekelwerks mit dem Exzimerlaser beim primären Offenwinkelglaukom. Ophthalmologe 93:565-568, 1996 Neuhann T, Scharrer A, Haefliger E: Excimer Laser-
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Trabekelablation ab interno beim chronischen Offenwinkelglaukom. Ophthalmo-Chirurgie 13:53-58, 2001 Kampmeier J, Stock K, Hibst R, Lang GE, Steiner R, Lang GK: Intrakanalicular trabeculostomy: a new approach to glaucoma surgery. Klin Mbl Augenheilk 212:159-162, 1998 Jacobi PC, Dietlein TS, Krieglstein GK: Microendoscopic trabecular surgery in glaucoma management. Ophthalmology 106:538-544, 1999 Joos KM, Alward WLM, Folberg R: Experimental endoscopic goniotomy. A potential treatment for primary infantile glaucoma. Ophthalmology 100:1066-1070, 1993 Rivera BK, Joos KM, Shen JH, Hernandez E, Ren Q: Endoscopic goniotomy with the Erbium:YAG laser. Invest Ophthalmol Vis Sci 35:2066, 1994 Rohrschneider K, Kruse FE, Kessler R, Gölz S, Bille JF, Völcker HE: Ab-interno-Trabekulotomie mit dem Nd:YLFPikosekundenlaser. Ophthalmologe 97:748-752, 2000 Juhasz T, Kastis GA, Suarez C, Bor Z, Bron WE: Timeresolved observations of shock waves and cavitation bubbles generated by femtosecond laser pulses in cornea tissue and water. Lasers Surg Med 19:23-31, 1996
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Laser cyclodestructive procedures of the ciliary body Geoffrey P. Schwartz, Louis W. Schwartz and George L. Spaeth Glaucoma Service, Wills Eye Hospital, Philadelphia, PA, USA
Keywords: glaucoma, cyclodestruction, cyclophotocoagulation, last resort, clinical results, side-effects
Introduction
History
For many years and, still, to date, the only method found successful for preventing continuing tissue damage in patients with glaucoma is the lowering of intraocular pressure (IOP). In this chapter, we discuss a group of procedures designed to benefit patients with glaucoma by lowering IOP. This IOPlowering is related to traumatizing the ciliary body and the ciliary processes. Cyclodestructive procedures (CDCs) of the ciliary body are usually reserved for treating glaucoma surgically when all other avenues of treatment have failed or are expected to fail. Medicines are used to try to lower IOP by decreasing aqueous production or increasing aqueous outflow, but medicines are not always successful in accomplishing this. Most laser and surgical procedures for glaucoma, such as laser trabeculoplasty, filtering surgery, tube shunts, goniotomy, and trabeculotomy, attempt to decrease IOP by increasing outflow; however, the body is difficult to fool. It wants to heal wounds. Filtering procedures have a tendency to fail. CDCs rely on a different approach to lowering IOP surgically, specifically, decreasing aqueous production. When the ciliary processes or the ciliary body are destroyed, aqueous production is decreased. This decrease in flow may in itself have effects that are harmful to the eye. Additionally, CDCs may be associated with complications that are difficult to control or reverse. Consequently, this is usually considered ‘last resort surgery’. Nevertheless, CDCs have significant advantages: they are fast, simple, can be performed on an out-patient basis, often do not require opening the eye surgically, and may be appropriate when other procedures are not.
Cyclodestructive surgery was first introduced as a method for treating glaucoma by Weve1 in 1933, when he introduced the use of non-penetrating diathermy for destroying the ciliary processes. In 1936, Vogt2 suggested penetrating the sclera with diathermy, which may have had a higher rate of success, but caused significant complications and only short-term benefits. Beta irradiation therapy3 and cycloelectrolysis4 were attempted in rabbit and human eyes, but never found acceptance clinically. Cyclocryotherapy was introduced by Bietti5 in 1950. In his technique, a probe was placed 1 mm posterior to the limbus and the temperature was reduced at the tip to –80°C for 60 seconds. Applications were placed over the ciliary body for a circumference of 180-360°. Freezing of the ciliary body led to ischemic necrosis of the ciliary processes. This treatment was believed to be more successful with fewer complications than cyclodiathermy, and consequently became the cyclodestructive procedure of choice. Shields6 reported the Duke Eye Center results in 114 consecutive eyes of 102 patients treated with cryotherapy over a ten-year period. IOP reduction was considered successful in 75 eyes (66%), while 14 (12%) developed phthisis. Vision became worse in 60%, and one-fourth required more than one treatment. Thirteen of the 14 eyes developing phthisis had no more than one treatment. Neovascular glaucoma, one of the prime indications for this treatment, was associated with the worst results at 64%, phthisis occurred in 17%, and worse vision in 69%. Eyes with aphakic glaucoma alone had the best results and were controlled in 82%, and phthisis occurred in only 3%, while reduced vision was
Address for correspondence: George L. Spaeth, MD, Glaucoma Service, Wills Eye Hospital, 840 Walnut Street, Philadelphia, PA 19107, USA. e-mail:
[email protected] Lasers in Ophthalmology – Basic, Diagnostic and Surgical Aspects, pp. 341–351 edited by F. Fankhauser and S. Kwasniewska © 2003 Kugler Publications, The Hague, The Netherlands
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noted in 52%. Because cyclocryotherapy has a high incidence of significant complications and a limited success rate, the search continued for a better cyclodestructive procedure. In 1961, xenon arc was introduced by Weekers et al.7 to transsclerally ablate the ciliary body, but was apparently no better than cyclodiathermy. In 1964, Purnell8 introduced transscleral ultrasound to destroy the ciliary body. Coleman9,10 refined this technique in rabbits and humans, but difficulty with the technique, unpredictability, and troublesome side-effects have limited its acceptance by clinicians. Transscleral cyclophotocoagulation (CPC) was successfully introduced by Beckman11 in 1972, with the use of a ruby laser in 241 eyes. In 1973, Beckman and Sugar12 reported the use of a neodymium laser to ablate the ciliary body. Transpupillary argon laser CPC was introduced about the same time.13 Since the introduction of lasers to treat the ciliary body in 1971, several other lasers have been utilized including the neodymium:yttrium-aluminum-garnet (Nd:YAG), diode, and krypton. This chapter will discuss various laser types and techniques presently being used to ablate the ciliary body. Indications, contraindications, complications, histology, and results will also be described. Indications Ablating the ciliary body is a destructive procedure associated with serious complications, and, as such, should usually only be used after all other modalities of lowering IOP have been employed. However, there are times when a patient is willing to have a laser treatment in a clinic or office setting but refuses to undergo the ‘knife’ under anesthesia in an operating room. At times, the vision is so poor that it is not worth the risk of an intraocular procedure when a ‘simple’ outpatient laser may be the better choice. These procedures have been suggested for controlling pain in blind, painful eyes. However, this is a highly controversial indication. Pain in association with glaucoma is either associated with a rapid rise in pressure or with severe inflammation in the eye. In itself, chronic elevation of IOP, is rarely a cause for pain. Patients in whom there is a rapid rise in IOP often have a type of glaucoma, such as acute primary angle-closure glaucoma, or glaucoma in association with trauma, in which CPC is not an appropriate procedure, because other, safer therapeutic methods are available and are usually effective. In a second groups of patients, those with inflamed, painful eyes, the appropriate treatment is usually directed at the pain, not at the level of IOP. The use of topical atropine and topical corticosteroids frequently causes such eyes to become comfortable. It is important to recall that the primary purpose of CDCs is to lower IOP, not to relieve
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Table 1. Indications for laser cyclodestructive procedures 1. 2. 3. 4. 5. 6. 7. 8. 9. 10. 11. 12.
Glaucoma after penetrating keratoplasty Pseudophakic open-angle glaucoma Aphakic open-angle glaucoma Congenital glaucoma, especially after failure of other surgical procedures Advanced angle-closure glaucoma, primary and secondary End-stage open-angle glaucoma Inflammatory open-angle glaucoma Traumatic glaucoma Silicone oil glaucoma Neovascular glaucoma Glaucoma in association with the aqueous misdirection syndrome Blind, painful eyes*
*Some authors believe that unilateral blindness is a contraindication for cyclodiathermy because of the possibility of inducing sympathetic ophthalmia
pain. It is also important to remember that CDCs have been associated with the development of sympathetic ophthalmia. Patients in whom CPC is to be employed need, of course, to be informed of pertinent complications, especially sympathetic ophthalmia, as that information frequently influences the choices they make. In those rare instances in which it is believed that it is appropriate to proceed with a CPC in a blind, painful eye of a patient who has a seeing fellow eye, it is essential to be sure that there is no intraocular tumor. Table 1 lists the types of glaucoma that have been treated with varying degrees of success by this method. The alternative treatments that are usually considered in this group of disease entities are cyclocryotherapy, filtering surgery with antimetabolite, or tube shunt. In each instance, the relative merits of each procedure must be considered with respect to the individual eye, such as vision potential, health of the conjunctiva and cornea, and status of the anterior chamber. Transscleral treatment does not require clear media, while transpupillary laser and some types of endolaser do. Contraindications There are few absolute contraindications to performing CPC. The main reason to seek an alternative method of treatment is a phakic patient with good vision. Some clinicians will not perform CPC if the patient has good vision or the potential for good vision, while others only eliminate those individuals who are phakic. Marked uveitis is both a relative contraindication and an indication for CPC. All patients develop increased inflammation following CPC, but uveitic glaucoma is one of the secondary glaucomas that can be treated successfully. Because trabeculectomy, even with antimetabolites, has a decreased success rate in patients with marked uveitis, CPC may be an attractive alternative therapy in these cases. Poor cooperation is a relative contraindication. If the patient is unable to
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Laser cyclodestructive procedures of the ciliary body cooperate by holding still, or cannot sit at the slitlamp, heavy sedation or general anesthesia may be required. As mentioned in the last section, total lack of vision in one eye is considered by some authors to be a contraindication for CDC because of the potential for initiating sympathetic ophthalmia. Complications It is very important to remember that eyes selected for CDCs, including CPC, are often seriously unhealthy eyes prior to CDC; such sick eyes not infrequently proceed to blindness despite treatment, but not due to treatment. Thus, trying to ascertain the actual complication rate of CDC and CPC itself is impossible. For example, eyes with severe neovascular glaucoma are often mortally wounded by the neovascular glaucoma itself, and despite every attempt, become blind. In such cases, it is not usually the attempts that cause the blindness, but the underlying illness itself. Complications are listed in Table 2. Uveitis-iritis occurs to some extent in all patients undergoing a CDC. Depending on the underlying disease process, this occurs to a varying degree. Following CPC, all patients have inflammation, which may last for years. The long-term use of corticosteroids is not required in many of these cases, although occasionally the inflammation may be sufficiently severe to necessitate their continued use. Pain occurs in all cases and is troublesome in about 50% of patients; this pain may last for several hours to several weeks. It is impossible to predict from the type of glaucoma being treated which patients will have this significant pain. Using a contact diode laser seems to cause less severe pain.14 It is usual for patients to develop some degree of cataract post-CPC. Hyphema and/or vitreous hemorrhage may occur, especially in patients with rubeosis. Corneal edema is an unusual secondary complication unless severe uveitis develops or the graft fails when treating glaucoma associated with penetrating keratoplasty. More often, once the IOP is lowered, corneal edema presents before the procedure disappears. Thinning of the sclera after contact CPC does not seem to cause long-term problems.15,16 Retinal detachment,17 sympathetic ophthalmia,18 and aqueous misdirection syndrome.19,20 have been reported, but are extremely rare complications. Anterior segment ischemia or necrosis may occur if the long ciliary arteries are coagulated. Consequently, care should be taken to avoid the 3 and 9 o’clock positions (Fig. 1). Loss of vision greater than two lines is a commonly reported complication, but these laser treatments are usually performed in patients with severe problems in addition to the glaucoma, and it is often difficult to differentiate progression of the underlying disease process (i.e., diabetic retinopathy, etc.) from effects of the laser treatment. Decreased accommodation may
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343 Table 2. Complications (in order of decreasing frequency) 1. 2. 3. 4. 5. 6. 7. 8. 9. 10. 11. 12. 13. 14. 15. 16. 17.
Acute uveitis Pain Increase in intraocular pressure (transient) Chronic uveitis Cataract (in phakic patients) Chronic hypotony Phthisis Vitritis Macular edema Vitreous hemorrhage Thinning of the sclera Hyphema Anterior segment necrosis or ischemia Retinal detachment Corneal edema Aqueous misdirection syndrome Sympathetic ophthalmia
Fig. 1. Laser burns near the limbus, avoiding the 3 o’clock position.
occur in young patients with atrophy of the ciliary muscle. Chronic hypotony and eventually phthisis are among the most feared complications, and occur in about 8-10%. Maus and Katz22 reported severe hypotony, flat anterior chambers, and serous choroidal detachment in three patients with previous filtering surgery. Having neovascular glaucoma, being black, and hypotony, are associated with a greater risk of visual loss.23 Technique Several different instruments have been used to perform laser CPC. CPC can be performed with a slit-lamp delivery system (non-contact), using an Nd:YAG,24-35 argon,6,36,37 or diode,38,39 light source, either by a transscleral or transpupillary method. Contact probe Nd:YAG,40-45 diode,39,46-57 and krypton21,58,59 lasers have been used successfully transsclerally to ablate the ciliary body. An endoscopic probe has been developed and used to coagulate the ciliary processes with a diode laser light source.60-63 This is performed through a limbal or pars plana approach. Beckman initially performed non-contact transscleral CPC using a ruby11 and
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Fig. 2. Shields lens for non-contact Nd:YAG CPC.
then a neodymium12 laser. However, these lasers were only available in his laboratory. In 1984, LASAG introduced the first commercially available Nd:YAG laser to perform this procedure (LASAG Nd:YAG microrupter II, Thun, Switzerland). It was developed under the direction of Fankhauser at the University Eye Clinic, Bern, Switzerland. Retrobulbar anesthesia is required with all CDCs. The procedures can cause excruciating pain, and the anesthesia must be effective. Some find a lid speculum helpful. The contact lens developed by Shields64 to keep the lids open (Ocular Instruments, Inc., Bellevue, WA) can also be used (Fig. 2). This lens also blocks laser light from entering the pupil, has marks at 1-mm intervals to help judge distance from the limbus, and can be used to blanch an inflamed conjunctiva in order to decrease superficial charring of the conjunctiva. However, many surgeons merely hold the lids apart with their fingers or applicators. Eight to ten burns are placed in each quadrant 14 mm posterior to the limbus26 for a circumferen-
tial distance of 180-360°. The distance posterior to the limbus is determined by the angle of the eye and the position of the ciliary body. For example, in an eye in a slight downward gaze, the applications are placed more posteriorly in the superior portion of the globe, and more anteriorly at the inferior aspect of the globe. Transillumination of the globe is helpful for identifying the exact position of the ciliary body. The 3 and 9 o’clock meridians are avoided in order not to coagulate the long posterior ciliary arteries. Energy levels of between 4 to 8 J are often utilized. The laser beam is aimed parallel to the visual axis and focused 3.6 mm internally to the conjunctival surface. The aiming beam is actually focused on the surface of the conjunctiva, but the laser is set such that, when fired, the Nd:YAG is focused in the ciliary body, causing a greater degree of destruction (Fig. 3). In general, the greater energy levels used, the more the inflammation. Consequently, some surgeons prefer to use less energy and to repeat the procedure more often. If the targeted pressure is not achieved by one month, the procedure may be repeated. This can be done as often as necessary (Figs. 4 and 5). In non-contact semiconductor diode laser transscleral CPC, a Microlase laser (Keeler, Inc., Broomall, PA) has been utilized. This unit has two infrared laser diodes with wave lengths of 780-830 nm. The laser burns are applied 1 mm posterior to the limbus and defocused 1 mm toward the ciliary body. The spot size is 100 µm with a power of 1.2 W at a pulse duration of 0.7-0.99 seconds. The laser light parallels the visual axis, and 40 burns are placed for 360°.38,39,65 Transpupillary argon laser is also delivered by means of a biomicroscope. It has the advantage of
Fig. 3. Non-contact laser focused on the conjunctiva and de-focused so that a laser beam of energy is absorbed by the ciliary body. [Reprinted by permission from Schwartz LW, Moster MR, Neodymium:YAG laser transscleral cyclodiathermy, Ophthalmic Laser Therapy, Vol. 1, No. 3, Spring 1986:137 (Figure 2). Publisher Mary Ann Liebert, Inc.]
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Laser cyclodestructive procedures of the ciliary body
Fig. 4. Ciliary processes as viewed through a large sector iridectomy before non-contact Nd:YAG CPC.
Fig. 5. Same patient as in Figure 4 several months after noncontact Nd:YAG CPC. Note atrophy of the ciliary processes.
visualizing the ciliary processes as they are being treated, with more precise application of the laser burns. However, it has the disadvantage of needing a large sector iridectomy in order to visualize enough ciliary processes to be successful. The laser is aimed through a Goldmann-type gonioprism with or without scleral depression. Enough burns are utilized to blanch visible ciliary processes with a 50-100 µm spot size for 0.1 seconds at 700-1000 mW. Settings are adjusted to obtain the desired blanch. Contact transscleral CPC was first reported using an Nd:YAG probe to ablate the ciliary body by Federman’s group in 1987.40 The contact probe was
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345 developed by Surgical Laser Technologies, Inc., Malvern, PA. It uses a 600-mM quartz fiberoptic contact probe. The tip is 2.2 mm in diameter and constructed of synthetic sapphire. The ideal setting seems to be 7 W for 0.7 seconds. Thirty-two to 40 applications are given for 360°, avoiding the 3 and 9 o’clock positions. The anterior edge of the probe is 0.5-1 mm posterior to the limbus. It is performed with the patient in the supine position after retrobulbar anesthesia has been given. The contact diode laser is probably the most popular method of transscleral CPC at this time. The laser is relatively small and portable (G-Probe: Iris Medical Instruments, Inc., Mountain View, CA). The patient is supine, retrobulbar anesthesia is given, and a lid speculum is placed. Ideal settings consist of 30-40 applications of 1.5-2.0 W of energy applied for 1.5-2 seconds over 360°, avoiding the 3 and 9 o’clock positions. The G-Probe is designed to deliver the laser energy 1.2 mm from the limbus, when the anterior edge of the probe is placed at the limbus. The fiberoptic tip protrudes 0.7 mm beyond the contact surface, thus indenting the sclera. If an audible pop is heard, the energy level is reduced by 0.25 W until no further pops are heard, since audible pops are associated with more inflammation and hyphema postoperatively.54 A krypton laser (Lasertek 41 AKTrKr – Lasermatic OY, Helsinki) has been utilized in Finland to perform transscleral CPC. The technique is similar to the other contact lasers. The probe is held perpendicularly 1.2-3 mm from the limbus, centered over the ciliary body which is identified by transillumination. It firmly compresses the sclera. Four to five hundred mW is delivered for ten seconds. Ten applications are given per quadrant for 90-360°. Postoperatively, in all these treatments, topical corticosteroids as well as sub-Tenon’s steroids are given to lessen the inflammation that occurs in all patients. Atropine drops may or may not be given. Ice packs may be prescribed for pain, as well as oral analgesics. When performing endoscopic cyclophotocoagulation (ECP), the laser endoscope has an image guide, light guide, and laser guide in a 20- or 18gauge endoprobe. With the 20-gauge probe, the field of view is 70°, with depth of focus of from 0.5-15 mm. With the 18-gauge endoprobe, the field of view is 110°, with a depth of focus of from 130 mm. The probe is connected to a video camera, light source, video monitor, and video recorder. The laser guide is attached to a semi-conductor diode laser at an 810 nM wavelength. The surgeon controls the surgery by viewing the video monitor. The ciliary processes are accessed from either a limbal or a pars plana approach. This is performed in the operating room under local retrobulbar anesthesia. Power settings of 500-900 mW are used, with continuous surgeon-controlled duration. Laser applications of from 0.5-2 seconds are used to produce an end point of whitening and shrinkage of each cili-
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ary process. Laser power or duration, or both, are decreased when popping is heard or bubbles form. An area of from 180-360° is treated, depending upon the desired effect. With the limbal approach, an incision is made approximately 2.5 mm in length, using a keratome, after the pupil has been dilated. Viscoelastic is then introduced between the iris and crystalline lens or pseudophakus, to access the ciliary process. A maximum of 180° can be treated through one incision. A second incision can be made 180° from the first. The viscoelastic is washed out after the procedure using a Simcoe cannula, and #10-0 nylon suture is used to close the wounds. Cataract extraction with an intraocular lens implant may be combined with this procedure and is performed first, following which the ECP can be performed through the same incision. When performing ECP through a pars plana incision, the patient must be aphakic or pseudophakic. The anterior vitreous is removed, and the laser endoscope inserted through the pars plana 3.5-4.0 mm from the limbus. The ciliary processes are easily viewed and photocoagulated. Two incisions are necessary if more than 180° are to be treated, and the sclerostomies are closed with #7-0 vicryl suture. Subconjunctival antibiotics and steroids are given at the end of the procedure. Topical antibiotics, steroids, and cycloplegics are given for as long as necessary. Histology Numerous histopathological studies utilizing both Nd:YAG and diode lasers have been performed on both animals and human autopsy eyes.24,64,66-71 Here, we present histological changes noted in the ciliary body of living eyes, which were then enucleated at various times post-treatment. Scleral transmission steadily increases with longer wavelengths.72 The destructive effect on the ciliary body depends on the light absorbency of the melanin in the ciliary pigment epithelium. The absorption coefficient of melanin is higher at 810 nm. It only takes half the energy with the diode (810 nm) than with the Nd:YAG (1064 nm) laser to produce similar lesions. However, because Nd:YAG penetrates the sclera more easily, at similar energy levels the lesions are equal. Ferry et al.73 conducted a study using a non-contact Nd:YAG laser in eyes that were enucleated one day, 20 days, and three months post-treatment. They concluded that: (1) energy levels of 4.4-5.6 J were effective; (2) laser focus 1.0-1.5 mm behind the limbus caused lesions of the pars plicata; (3) toward the periphery of the individual treatment sites, the stroma and ciliary muscle continued to exhibit severe degeneration, as did the epithelium lining the valleys between the crests of the ciliary processes; (4) bleb-like separations of the ciliary epithelium were a prominent early feature; (5)
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pigmented epithelium was affected more than nonpigmented epithelium; (6) destruction of the ciliary epithelium was permanent; (7) deeply-pigmented persons have more melanocytes in the ciliary body, muscle, and stroma, than do more lightly-pigmented individuals, thus making more-pigmented individuals more responsive to treatment; (8) the ciliary muscle was always severely damaged; (9) no significant damage was noted in the sclera; and (10) few inflammatory cells were noted, except for in the episclera. Damage noted was due to coagulative or hemorrhagic necrosis. Blasini et al.74 found similar results one to three days after enucleation following non-contact Nd: YAG laser CPC. Most notable was the disruption of the ciliary epithelium with separation of the epithelium layers from the underlying stroma. Fibrin and a few leukocytes were noted between these separated structures. Both contact the diode and contact Nd:YAG lasers cause similar lesions at similar energy levels. At lower energy levels, both produce thermal damage in the ciliary body and coagulative necrosis of the ciliary non-pigmented and pigmented epithelium. As the energy is increased, the damage increases with widespread disorganization of the ciliary processes and more marked thermal coagulation of the ciliary body, with vacuolization and collagen fiber swelling. No explosive-like lesions were produced with either laser. In addition, no significant scleral damage was noted.75,76 Ten months after contact krypton CPC, in enucleated eyes, the posterior processes had totally disappeared and the adjacent inner connective tissue layers and ciliary muscles were markedly atrophic. Clump-like cells surrounded dispersed pigment in the subepithelial stroma. The adjacent zonules and overlying sclera were normal.21 Results Transpupillary argon laser The largest study of transpupillary argon laser ablation of the ciliary body was reported by Shields.36 Of the 27 patients treated, six (or 22%) had a successful outcome. Failure may partially be related to the number of ciliary processes, which can be visualized and treated, and partially to the number of ciliary processes that can be treated, since, even with scleral indentation, only a small part of the ciliary process can be ablated. In addition, a white reaction may not indicate tissue destruction, and the energy level needs to be increased until a brown burn is produced with pigment dispersion and/or gas bubble formation.6
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Laser cyclodestructive procedures of the ciliary body Non-contact Nd:YAG CPC Table 3 presents the results of several studies using non-contact Nd:YAG CPC. Most studies have a success rate of around 70%, although the range is from 45-87%. Decreased vision greater than two Snellen lines appears to be the most significant complication. Hypotony ranged from 0-15%. It should be noted that, in most studies, repeat treatments were needed in a significant number of patients. Contact Nd:YAG CPC The largest study of contact transscleral Nd:YAG CPC in humans was reported by Schuman et al.44 This group followed 116 eyes for a minimum of one year (range, 12-36 months). Utilizing the Surgical Laser Technologies Nd:YAG laser, up to 40 applications of 7-9 W for 0.7 seconds each were delivered for 360°, avoiding the 3 and 9 o’clock meridians. Seventy-three percent of eyes required one treatment, but up to six treatments were performed in one eye. Success, defined as an IOP between 3 and 22 mmHg, was 65%, while 8% of eyes had an IOP below 3 mmHg. Forty-one percent lost two or more lines of vision, and 16% progressed to no light perception. Non-contact transscleral diode laser CPC The non-contact diode laser has been little used. Hennis and Stewart38 reported their results using this mode of treatment for advanced glaucoma. A 100-µm spot size, for 990 msec at 1.2 W of power was given for 40-45 applications, 360° 1 mm from the limbus in 14 eyes. Only one treatment was given. Ten of the 14 patients (71%) were considered successes, based on an IOP of less than 21 mmHg or total pain relief. There was no visual loss, hypotony, or phthisis in the six-month followup. Contact transscleral diode laser CPC Table 4 lists the results of several studies utilizing the most commonly used diode laser (Iris Medical SLX with the G-Probe). Success rates varied from 50-100% with follow-ups ranging from one to 40 months. Hypotony, defined as an IOP under 5 mmHg, ranged from 0-8%; most studies being in the 2-4% range. Loss of vision of two lines or more occurred in 10-38%. All these studies allowed repeat treatments if the initial treatment did not produce the desired effect. Endoscopic diode laser CPC Uram60 described his experience with endoscopic laser CPC (ECP) in 1995. In 150 patients with neovascular glaucoma, followed for six to 52 months, after 180° treatment, there was a 67.6%
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347 mean decrease in IOP to 15.6 mmHg. Eight percent required a second treatment. No phthisis was reported, and treatment for pain control was reported to be 100% successful. In aphakic and pseudophakic glaucoma, a similar good response was noted in 108 patients with a mean decrease in IOP of 56.1%. No cystoid macular edema (CME) was seen at three months, however, 8.3% had progressive field loss at one year. During a follow-up of six to 50 months, no phthisis or endophthalmitis was reported. Eighteen percent required a second procedure. In another group Uram studied 56 eyes with phakic, uncontrolled open-angle glaucoma utilizing ECP.60 The mean decrease in IOP was 46.3%. None of the patients developed lens dislocation, CME, phthisis, or endophthalmitis in the six to 33 months of follow-up. At one year, 8.9% had developed visual field loss, and 14.3% had developed cataracts by the last postoperative visit. These results have not been confirmed by other investigators. In 1997, Chen et al.62 reported their results with ECP in 68 patients with various forms of advanced glaucoma which had failed other forms of therapy. Treatment was performed between 180 and 360°, and a second laser treatment was required in 7%. Patients were followed for from six to 25.8 months, with a mean of 12.9 months. Success, defined by an IOP of less than or equal to 21 mmHg, occurred in 90%. Four eyes (6%) lost two or more lines of vision, while no hypotony or phthisis was noted. Twelve of the 21 phakic eyes had combined cataract extraction with ECP. Of the nine remaining phakic eyes, one (11%) developed a cataract. Krypton laser CPC The contact krypton laser has been used in Finland to perform CPC in advanced glaucoma. The results with this laser parallel those achieved with contact and non-contact Nd:YAG and contact diode lasers. Immonen58 reported on 62 eyes, with a success rate (8-24 mmHg) after six months of 82% in eyes treated to preserve vision and of 50% in eyes to relieve pain (72.8% overall). Hypotony (IOP of less than 8 mmHg) was noted in 3.6%. A decline in visual acuity of more than two Snellen lines was seen in 9.6%. This same group found the krypton laser to be effective in treating post-traumatic glaucoma (56% success; 8-21 mmHg),59 and in children and young adults (64% success; 8-21 mmHg).78 Conclusions The use of CDCs to treat the refractory glaucomas has a long history, and is steadily being refined. A few authors have suggested CPC as a primary treatment for primary open-angle glaucoma, especially in underdeveloped countries.79 When comparing the results of cyclocryotherapy with those of laser CPC, there is less pain, inflammation, and loss of vision in the laser group, but the success rates are similar.
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128 160
Baez 33 Dickens35
40 13-44
32 32 13-40 10-40 10-36 20 40 30-50 30
No. of applications
5-7 5.8-8
6.5 0.5-2.75 3.4-8.6 4-6 av. 6.2 mean 6.9 mean 6.8-1.2 1.5-6.5 4-8
Energy per application (joules)
No. of patients
210 48 20 27
26 41
47 31 12 58 37
Group
Bloom48 Brancato43 Brancato46 Kosoko47
Bock 49 Yap-Veloso52
Threlkeld53 Rebolleda54 Rebolleda54 Spencer55 Pucci57
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3.6-4.8 2.8-5.5
2.25 3.9 3.9 3.5-4
Energy per application (joules)
variable for 360° 3-5 variable for 270° 3-4 variable for 270° 3-4 12-14 (270°) 4 8-15 (90-270°) 3.2-4
19-36 8-40
20-40 16-20 16-20 17-19
No. of applications
Table 4. Success of contact diode cyclophotocoagulation in various studies
18 29 35 60 36 45 21 28 458
No. of patients
Beckman12 Schwartz25 Wright28 Brooks27 Suzuki29 Hardten30 Hardten30 Cohen77 Shields32
Group
1-4 1-3 1-2 1-5 1-4
1-6 1-4
1-5 1-4 1-4 1-2
No. of treatments
1-8 1-8
1-2 1-5 1-3 1 or more 1 mean 2 1-3 1-5 1-6
No. of treatments
Table 3. Success of non-contact Nd:YAG cyclophotocoagulation in various studies
66% 70.8% 100% 72% 1 year 52% 2 years 50% 64%
≤22 >2≤21 no pain >20% reduction or ≤22 ≤21 >20% reduction or ≤22 7-21 ≤5-≤21 no pain
