Abstract
Lasers enable exquisitely precise spatial and temporal control of the light delivery to tissue, which can enact a wide variety of interaction mechanisms, including photochemical, photothermal, and photomechanical. A broad range of wavelengths and pulse durations provide excellent means for very selective interactions with the target layer of cells in the retina. Long exposures at low power can activate photosensitizers in photodynamic therapy, while avoiding tissue heating. Millisecond exposures are used for coagulation of photoreceptors and sealing of microaneurysms. Below photocoagulation threshold, millisecond exposures can produce sublethal hyperthermia, enabling nondamaging therapy of the macula. Microsecond and nanosecond pulses can selectively destroy RPE by exploding melanosomes, without damaging photoreceptors. Computer-guided pattern scanning allows rapid and reliable delivery of visible or subvisible treatments to designated locations. Progress with understanding the tissue response to laser treatment helps perfect the dosimetry and improve clinical efficacy while minimizing or avoiding the detrimental side effects.
Key words
laser, photocoagulation, hyperthermia, selective retina therapy, photodynamic therapy, nondamaging retinal laser therapy
Introduction
Historically, a number of light sources have been utilized for retinal phototherapy including the sun, various flash lamps, and lasers. The sun light is capable of producing a retinal burn either accidentally (e.g., in the case of solar eclipse gazing) or purposefully, as demonstrated first by Meyer-Schwickerath. However, its dependence on weather conditions, its constant motion in the sky, and relatively large angular size (0.52 o ) make its use impractical for retinal therapy.
Arc lamps are very bright sources of light used in many therapeutic and imaging applications. In such lamps, a high-voltage discharge ionizes gas between an anode and cathode creating an electric arc – a high current density discharge. The ions emit light at specific wavelengths, and the spectrum of the plasma emission depends on the types of atoms involved, their temperatures and gas pressure. Thus the spectrum of an arc lamp may have a distinct signature. The xenon arc lamp was the first to become widely used for retinal photocoagulation because of its strong visible and near-infrared emission, convenience, and low price. However, because of its large size, poorly collimated beam, and tendency to produce intense retinal burns, it was replaced in clinical practice by laser-based systems in the early 1970s.
Lasers became the preferred light source for retinal photocoagulation due to their narrow spectrum, wide selection of wavelengths, excellent collimation (directionality), high brightness, and variable pulse duration. The highly collimated laser beam is easy to guide before its introduction into the eye, and to focus into very small spots. Its monochromaticity makes it possible to choose a wavelength for selective absorption in specific tissues of the eye. Adjustable pulse duration allows controlling the extent of thermal diffusion, thus producing very precise and selective interactions with minimal collateral damage.
The most widespread medical applications of lasers in medicine have been in ophthalmology. Since the introduction of the ruby laser more than three decades ago, ophthalmic laser applications have experienced rapid growth with the use of argon, krypton, argon-pumped dye, Nd : YAG, diode, Er : YAG, excimer, and Ti : Sapphire lasers. Lasers have been applied to a wide variety of slit lamp-based retinal therapies, as well as to vitreoretinal surgery, glaucoma therapy, posterior capsulotomy, and refractive surgery. These applications are based on different mechanisms of laser–tissue interactions, including photothermal, photodisruptive, and photochemical interactions. The most common vitreoretinal application is retinal photocoagulation. Additionally, a number of novel therapies have recently been introduced and are under active evaluation, including selective retinal pigmented epithelium (RPE) treatment (SRT) and nondamaging retinal therapy (NRT).
In the following sections we will describe the underlying principles of laser–tissue interactions and the types of lasers available and appropriate for various retinal applications.
Optical Properties of the Eye
The relaxed eye has an approximate optical power of 60 diopters (D) (i.e., its focal length is 16.7 mm), with the corneal power being about 40 D, or two-thirds of the total power. Due to orderly arrangement of collagen fibrils in the cornea, it is highly transparent – with transmission above 95% in the spectral range of 400–900 nm. The refractive index of the cornea is n≈1.3765± 0.0005. The amount of light reaching the retina is regulated by the pupil size, which can vary between 1.5 and 8 mm. The anterior chamber of the eye, which is located between the cornea and lens capsule, is filled with a clear liquid – the aqueous humor having a refractive index n ≈1.3335. The crystalline lens of the eye, located behind the iris is composed of specialized crystallin proteins with refractive index of n =1.40–1.42. The lens is about 4 mm in thickness and 10 mm in diameter and is enclosed in a elastic, thin (5–15 µm), transparent collagenous capsule. In the relaxed eye the lens has a power of about 20 D, while in the fully accommodated state it can temporarily increase to 33 D. The vitreous humor – a transparent jelly-like substance filling the large cavity posterior to the lens and anterior to the retina has a refractive index n ≈1.335.
Light entering the eye can be reflected, scattered, transmitted or absorbed. Reflected or scattered light contains information that can be used for noninvasive diagnostic purposes. The absorption characteristics of ocular tissues are determined by the chromophores resident within the tissue. In the visible part of the spectrum (400–750 nm) these chromophores include: (a) melanin located in the retinal and iris pigmented epithelia, choroid, uvea, trabecular meshwork; (b) hemoglobin, located in the red blood cells; (c) macular xanthophyll, which is located in the plexiform layers of the retina, especially in the macula; (d) rhodopsin and cone photopigments which are located in the photoreceptors; and (e) lipofuscin located primarily in the RPE layer. These pigments are of major importance in absorption of visible light in the retina from both a physiologic and pathologic standpoints. The absorption spectra of these pigments, as well as water and proteins are illustrated in Fig. 41.1 . In the mid-infrared part of the spectrum (3–15 µm) the major absorber is water, while protein absorption is dominant in ultraviolet range (below 250 nm).
Basics of Lasers
The term “laser” is an acronym for light amplification by stimulated emission of radiation. A beam of light is composed of individual packets of energy that are called quanta or photons. Each of these photons has a particular energy and direction of travel. The energy of a quantum of light is proportional to its frequency, i.e., it is a reciprocal of its wavelength. In the presence of a properly prepared laser material, it is possible for a quantum of light to trigger the release of other quanta with the same wavelength and direction of travel. This phenomenon is called stimulated emission, and it is an essential element in lasing. In thermal equilibrium the energy levels of atoms and molecules are populated according to the Boltzmann distribution, in which upper levels are always less populated than the lower levels. Stimulated emission requires an inverted population of energy levels, such that the upper levels are more populated than the lower ones. Therefore, lasing can occur only when material is not in thermal equilibrium. The nonequilibrium state of the population inversion is created in the lasing material by an excitation source or “pump.”
Generally, a laser is composed of three basic components: (1) a material that can store energy to be released by stimulated emission; (2) a means of replenishing the energy stored in the laser material; and (3) some method of retaining a fraction of the light emitted by the lasing material to stimulate further emission. Fig. 41.2 schematically illustrates a general configuration of a laser. An energy source is used to introduce energy into the lasing material. This energy is stored as atomic or molecular excitation waiting to be released by stimulated emission. Laser light emitted by the lasing material circulates between the two mirrors on either end of the laser cavity, with a fraction of the light escaping through one mirror to form the laser beam. The trapped light stimulates emission of new quanta of light from the laser material with the same wavelength and direction as the original quanta. This way laser produces a beam of light in which all of the quanta move “in phase” with one another. This property of light is called coherence. Duration of the synchronized emission multiplied by the speed of light is called the coherence length of the laser emission. This is the distance along which the photons are coherent, or moving “in step.” To remain in phase with one another, these quanta must have approximately the same wavelength. Thus temporal coherence is related to the monochromaticity (or spectral width) of the light emitted from the laser: the broader is the spectrum the shorter is the temporal coherence. A laser may produce one or several discrete spectral lines in either the infrared, visible or ultraviolet domains, in contrast to conventional light sources (incandescent or arc lamps) which typically produce polychromatic (broad-band or white) light.
Collimation (directionality) of the emitted beam is governed by the mirror configuration of the laser cavity. In its simplest form, a cavity consists of two mirrors arranged such that light bounces back and forth, each time passing through the gain medium. One of the two mirrors, the output coupler, is partially transparent, allowing the output beam to exit through it (see Fig. 41.2 ). The reflection coefficient of the output coupler determines how many times photons are reflected back to circulate inside the cavity before exiting it. For example, with a reflection coefficient of 0.99, the photon will bounce, on average, 99 times before exiting the cavity. The structure of the laser cavity determines directionality (collimation) of the laser beam, which determines its ability to be focused into a small spot.
The lasing medium can be a gas, liquid, or solid. Lasers can be pumped by continuous discharge lamps and by pulsed flash lamps, by electric discharges in the laser medium, by chemical reactions, by an electron beam, by direct conversion of electric current into photons in semiconductors, or by light from other lasers. Laser pulse durations can vary from femtoseconds to infinity. Pulsing techniques used for different ranges of pulse durations include electronic shutters (down to 1 ms), pulsed flash lamps (typically down to a few µs), Q-switching (down to a few ns), or mode-locking (down to fs).
Laser Beam Delivery to Tissue
Laser beams are typically very well collimated. Diffraction causes light waves to spread transversely as they propagate, and it is therefore impossible to have a perfectly collimated beam. The diffraction-limited divergence angle of a Gaussian beam with diameter D and wavelength λ is 
. As an example, for a laser emitting a 1 mm wide beam at 532 nm wavelength, the divergence (half-angle Θ) is about 0.66 mrad, i.e., the beam spreads by 1.3 mm over a distance of 1 meter.
Using a lens or a concave mirror with focal length f , a laser beam can be focused to a spot with a diameter 
. The depth of the focal region is: 
(see Fig. 41.3 ). With a lens having focal distance f = 25 mm, the same 532 nm laser beam can be focused to a spot of 16 µm in diameter, having a focal depth of about 850 µm. It must be emphasized though that exact definition of the spot size depends on the beam profile, which varies with configurations of the laser cavities. For retinal photocoagulation such tight focusing is usually not required, and laser spots typically vary from 50 to 500 µm in diameter in various therapeutic applications.
As an alternative to a free-propagating beam, laser light can be transported via optical fibers. An optical fiber, schematically shown in Fig. 41.4 , typically consists of a core, cladding, and jacket. Light is trapped within the core due to total internal reflection at the interface of the core with cladding. To satisfy conditions of total internal reflection, the incidence angle of light at the core/cladding interface should not exceed the critical angle of total internal reflection: sinΘ cr = n core / n clad , where n core and n clad are refractive indices of the core and cladding, respectively. To satisfy these criteria for total internal reflection, the light should be launched within a so-called acceptance cone, and the sine of its half-angle Θ ac is defined as theNumerical Aperture of the fiber: 
. Typically the NA is within a range of 0.1–0.2. Optical fibers are often used for delivery of laser light to slit lamp-based systems ( Fig. 41.5 ), and to the intraocular surgical probes ( Fig. 41.6 ).
Aberrations
With nonperfect focusing optics, the focal spot size of the laser beam is limited not only by diffraction, but also by aberrations. Measurements of optical aberrations in the human eye demonstrate that for pupillary dilation of up to 3 mm in diameter, an average emmetropic human eye is optically well corrected, and the focal spot is close to the diffraction limit. However, for pupils greater than 3 mm in diameter, the central aberrations increase, resulting in increases in the focal spot size. Peripheral field aberrations lead to rapidly increasing blur of the image with angle of visual field, strongly limiting the focusing capability of laser in the periphery of the retina. In retinal photocoagulation, a flat contact lens is typically used to reduce the optical power of the front surface of the cornea. If the lens is used properly, it aids greatly in controlling peripheral aberrations during photocoagulation. Other methods include using an aspheric lens to control optical aberrations in the periphery during photocoagulation. Such lenses typically provide wide-field viewing, although aberrations are difficult to correct over the totality of fields of interest and additional reflections may be introduced by the lens surfaces.
Contact Lenses
Currently, retinal laser photocoagulation relies heavily on the use of contact lenses developed for this purpose, and the most common types are listed in Table 41.1 . The universal (Goldmann) three-mirror contact lens provides a flat front surface that nearly cancels the positive refractive power of the front surface of the cornea. Mirrors at 59°, 67°, and 73° aid in visualization and photocoagulation of the periphery and anterior segment. To obtain the most reproducible results in photocoagulation the operator should hold the contact lens so that the flat surface is nearly normal (within ±5 degrees) to the laser beam ( Fig. 41.5 ). The use of mirrors in contact lenses helps the operator keep the laser beam properly aligned to the lens while photocoagulating over a large field.
| Lens | Image Magnification | Laser Beam Magnification |
|---|---|---|
| Ocular Mainster Std | 0.95 | 1.05 |
| Ocular Fundus Laser | 0.93 | 1.08 |
| Ocular Karichoff Laser | 0.93 | 1.08 |
| Ocular 3 Mirror Univ. | 0.93 | 1.08 |
| Ocular Mainster Wide | 0.67 | 1.50 |
| Ocular Mainster Ultra | 0.53 | 1.90 |
| Ocular Mainster 165 | 0.51 | 1.96 |
| Rodenstock Panfundoscope | 0.67 | 1.50 |
| Volk G-3 Gonio | 1.06 | 0.94 |
| Volk Area Centralis | 1.06 | 0.94 |
| Volk TransEquator | 0.69 | 1.44 |
| Volk SuperQuad 160 | 0.5 | 2.00 |
| Volk QuadrAspheric | 0.51 | 1.97 |
| Volk HRWF | 0.5 | 2.00 |
| Goldmann 3 mirror | 1.00 | 1.00 |
Another useful photocoagulation lens is the inverted image lens system, typified by the Rodenstock, Quadraspheric, and Mainster photocoagulation lenses. These lenses contain a lens element in contact with the corneal surface and another positive lens element at a fixed distance from the cornea. These systems magnify the spot size on the retina, while increasing the field of view, requiring the operator to adjust the power accordingly. Magnification factors of the most common contact lenses are listed in Table 41.1 . It is important to keep in mind that magnification of the retinal image demagnifies the beam size on the retina by the same amount, i.e., the higher magnification of the retina, the smaller is the laser spot on the retina.
Interactions of Light With Tissue
In linear interactions the irradiance (or power fluence) I(z) [W/cm 2 ] of the beam propagating inside the tissue decreases exponentially with depth (Beer–Lambert law) due to absorption and scattering of light: I (z) = (1- k s ) I 0 exp(-µz), where I 0 is the light intensity at the surface of tissue ( z =0), k s is the specular reflection coefficient at the tissue surface, µ=µ a +µ s is the attenuation coefficient combined of absorption and scattering components. The penetration depth (δ) of the beam into tissue is defined as a depth at which light intensity is reduced by a factor of e : δ=1/µ. Reflection of light from ocular tissue at normal incidence typically does not exceed 2%. As shown in Fig. 41.1 , absorption of light by various chromophores in the eye strongly varies with wavelength. Scattering of light is also a very strong function of the wavelength: scattering on subwavelength inhomogeneities of refractive index in ocular media (e.g., collagen fibrils) is reciprocal to the fourth power of the wavelength: µ s ~ λ -4 (Rayleigh scattering). For example, the scattering coefficient for light at 1064 nm is 16 times lower than that at 532 nm. Scattering from structures larger than the wavelength (e.g., cellular organelles) is described by Mie theory and has more complex wavelength dependence and spatial distribution. Scaling of the Mie scattering coefficient with wavelength in various tissues can be approximated as µ s ~λ -b , with b ≈ 0.5–2.
Photochemical Interactions
Photochemical interactions are based on nonthermal light-induced chemical reactions. The common natural photochemical reactions are photosynthesis in plants and phototransduction in photoreceptors. Therapeutic photochemical interactions used in photodynamic therapy (PDT) take place at very low power densities (typically <1 W/cm 2 ) and long exposures – lasting tens of seconds. For PDT, special chromophores, called photosensitizers, are injected intravenously and allowed to accumulate in the target tissues. After excitation by laser radiation, the photosensitizer produces highly cytotoxic reactants – free radicals or singlet oxygen, which, in turn, produce irreversible oxidation of the nearby cell structures.
The concept to treat exudative age-related macular degeneration (AMD) by selectively targeting vascular endothelial cells using a specific photosensitizer-carrier complex (phthalocyanine, CASPc) activated by near-infrared laser was adapted from tumor therapy. Closure of subretinal neovascularization using PDT with Rose Bengal was demonstrated a few years later. Following upon the successful initial results with phthalocyanine, liposomal benzoporphyrin derivative (BPD) was developed, which selectively attaches to the endothelium of new blood vessels (verteporfin, or Visudyne TM ).
Clinical Indication: Photodynamic Therapy for Subfoveal Choroidal Neovascularization
Photodynamic therapy for subfoveal choroidal neovascularization (CNV) is performed as a two-step procedure. First, a specialized spectrally adapted photosensitizer, such as verteporfin is injected intravenously. Within minutes it is distributed via the bloodstream to tissues, including the retina and choroid. It selectively accumulates in neovascular tissue which is rich in low-density lipoprotein receptors, while it is rapidly cleared from the surrounding normal tissues. This differential accumulation and clearance is the basis for selectivity of PDT against neovascularization, compared with the normal choroidal and retinal vasculature. Increased presence of new blood vessels in tumors is also the reason for the use of PDT as a treatment for several ocular and nonocular tumors.
Upon absorption of a photon at the proper wavelength the porphyrin molecule undergoes a transition from the ground state into the singlet excited state ( 1 P*). The singlet excited porphyrin can decay back to the ground state with release of energy in the form of fluorescence, which enables optical identification of tumor tissue ( Fig. 41.7 ). The singlet can also be converted into the triplet excited state ( 3 P*) which is able to transfer energy to another triplet. One of the very few molecules with a triplet ground state is dioxygen, which is found in most cells. Such energy transfer produces toxic singlet oxygen ( 1 O 2 ) from ground state dioxygen ( 3 O 2 ). Singlet oxygen is very reactive and therefore it has a very short diffusion path length – less than 20 nm, so all its interactions are highly localized. The photochemical processes and subsequent reactions are complex and vary for different sensitizers, which are also subject to the microenvironment. Singlet oxygen and other highly reactive species damage endothelial cells in close proximity to the adsorbed drug resulting in thrombosis and temporary closure of the neovascularization.
Many photosensitizing dyes have been tested for PDT, including Rose Bengal, hematoporphyrin derivatives, lutetium texaphyrin Photofrin, and various benzoporphyrins. Only verteporfin (Visudyne TM ) has been approved for ophthalmic use in treatment of classic and occult CNV. Verteporfin has a very broad absorption spectrum, but only the far-red peak at 688–691 nm is typically utilized in clinical practice ( Fig. 41.8 ). This is because of the lower sensitivity of retina to far-red light and its superior penetration into the choroid.
In PDT treatment with verteporfin, the activation by laser is typically performed 15–20 min after the intravenous injection of the dye. A beam of red laser light (689 nm diode laser) is applied to the retina via a slit lamp delivery system. The irradiated area on the retina is typically chosen to minimally exceed the dimension of the neovascularization, with light intensity of 600 mW/cm 2 , for 83 seconds, resulting in a total radiant exposure of 50 J/cm 2 . Closure of the abnormal (leaking) blood vessels occurs within approximately 6–12 weeks in most patients. Reperfusion is common and multiple treatments are often required. Fluorescein angiography images shown in Fig. 41.9A (pretreatment) and 41.9B (1 week post treatment) demonstrate closure of a subfoveal CNV membrane after PDT.
Photothermal Interactions
Absorption of light in tissue leads to heating. Temperature is the governing parameter in all thermal laser–tissue interactions. Depending on duration and value of the temperature, different tissue effects may occur, including necrosis, coagulation, vaporization, and carbonization.
Heat generation in tissue is determined by the laser parameters and optical tissue properties – irradiance, exposure time, and the absorption coefficient, which depends on the laser wavelength. Heat transport is characterized by heat conductivity and heat capacity.
If heat diffusion is neglected, then at a constant beam power the temperature rise is linear with time: 
, where ρ is tissue density, and c is its heat capacity ( c = 4.2 J/g/K, ρ=1 g/cm 3 for water). To assess whether heat diffusion plays a significant role during the laser pulse, one should compare pulse duration with a characteristic time it takes for heat to spread by the distance equal to the zone of initial heat deposition in tissue. For the heated zone (laser penetration depth) of length L , the heat diffusion time is: τ= L 2 /4α, where α is thermal diffusivity (α=1.4 10 -3 cm 2 /s for water). For example, for L =1 µm in water the characteristic heat diffusion time is t =1.7 µs, while for L =1 mm, the diffusion time t =1.7 s. If the laser pulse duration is comparable or longer than the characteristic diffusion time across the light absorption zone, then proper estimation of the peak temperature in tissue should take heat diffusion into account.
Necrosis
Temperature rise induces conformational changes in various proteins, which denature at characteristic rates specific to protein species. These thermal processes, which may eventually lead to cell necrosis, depend on both the temperature and duration of the hyperthermia. Thermal damage to cells in a millisecond range of pulse durations can be approximated using the Arrhenius model. It assumes a rate of decline in the concentration of a critical molecular component D(t) with temperature T(t) :
d D ( t ) = − D ( t ) ⋅ A ⋅ exp ( − E * R ⋅ T ( t ) ) d t
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