Introduction

Historically, a number of light sources have been utilized for retinal phototherapy, including the sun, various flash lamps, and lasers. The sun is capable of producing a retinal burn either accidentally (e.g., in the case of solar eclipse retinopathy) or on purpose, as demonstrated first by Meyer-Schwickerath.¹ However, its dependence on weather conditions, its constant motion in the sky, and relatively large angular size (0.52°) make it an impractical method for intraocular therapy.

Arc lamps are very bright sources of light used in many therapeutic and imaging applications. In such lamps, a high-voltage discharge initially ionizes gas between an anode and cathode, creating an electric arc – a high-current-density discharge in the lamp. 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 and tendency to produce intense retinal burns, it was replaced in clinical practice by laser-based systems in the early 1970s.²

Lasers became the preferred light source for retinal photocoagulation due their narrow spectrum, wide selection of wavelengths, excellent collimation (directionality), high brightness, and variable pulse duration. The directionality of the laser makes it easy to manipulate the beam optically before its introduction into the eye, and to focus it 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 limiting the thermal diffusion to small distances, 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, lens capsule opacification, 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 pigment epithelium (RPE) treatment and sublethal thermal therapy.

In the following sections we will describe the underlying principles of laser–tissue interactions and the types of lasers available and appropriate for various vitreoretinal applications.

Optical properties of the eye

The relaxed eye has an approximate optical power of 60 D (i.e., its focal length is 16.7 mm in air), with the corneal power being about 40 D, or two-thirds of the total power.³ 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.⁴ The refractive index of the cornea is n ≈ 1.3765 ± 0.0005.⁴ 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, which has 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 tough, 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 ≈ 1.335.⁴

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 (400–800 nm) these chromophores include: (1) melanin located in the retinal and iris pigmented epithelia, choroid, uvea, and trabecular meshwork; (2) hemoglobin, located in the red blood cells; (3) macular xanthophyl, which is located in the plexiform layers of the retina, especially near and in the macula; (4) rhodopsin and cone photopigments which are located in the photoreceptors; and (5) lipofuscin, located primarily in the RPE layer. These pigments are of major importance in absorption of visible light in the retina from both physiologic and pathologic standpoints. The absorption spectra of these pigments, as well as water and proteins, are illustrated in Figure 39.1. In the mid-infrared part of the spectrum (3–15 µm) the major absorber is water, while in ultraviolet (below 250 nm) protein absorption is dominant.

Fig. 39.1 (A) Absorption coefficients for major chromophores in a spectral range of 0.2–10 µm. (B) Absorption coefficients of the major ocular chromophores in the visible part of the spectrum: 400–900 nm. Spectral locations of some of the popular laser lines in this range are indicated above the plots.

Basics of lasers

The term “laser” stands 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 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. As a result, lasing can occur only when material is not in thermal equilibrium. The nonequilibrium state 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 lasing material; and (3) some method of retaining a fraction of the light emitted by the lasing material to stimulate further emission. Figure 39.2 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 that has already been 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. In this way, a laser produces a beam of light between the two mirrors in which all of the quanta move in phase with one another. This property of light is called coherence.

Fig. 39.2 Laser typically consists of the energy source (pump), the lasing medium, and the optical cavity with a partially transparent front mirror.

Coherence is related to synchronization of light in time, or along the laser beam. The duration of the synchronized emission from the laser 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 the spectrum, the shorter the temporal coherence. A laser may produce one or several discrete spectral lines in the infrared, visible, or ultraviolet domains, in contrast to conventional light sources (incandescent or arc lamps) which typically produce noncoherent polychromatic (broadband 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 (Fig. 39.2). 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 an argon laser emitting a 1-mm-wide beam at 515 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 (Fig. 39.3). With a f = 25 mm lens the same argon laser beam can be focused to a spot of 16 µm in diameter, having a focal depth of 820 µm. It must be emphasized, though, that exact definition of the spot size depends on the beam profile, which varies in various configurations of laser cavities. For therapeutic laser photocoagulation such tight focusing is usually not required, and laser spots typically vary in diameter between 50 and 500 µm in various applications.

Fig. 39.3 Laser beam of diameter D focused with a lens of focal length f produces a waist with diameter d and a focal depth F (see equations in the text).

As an alternative to a free-propagating beam, laser light can be transported via optical fibers. An optical fiber, schematically shown in Figure 39.4, typically consists of a core, cladding, and jacket. Light is trapped within the core due to total internal reflection at the interface of 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 the numerical aperture (NA) 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. 39.5), and to intraocular surgical probes (Fig. 39.6).

Fig. 39.4 Optical fiber typically consists of a core, cladding, and a jacket (buffer coating). Light launched into the fiber with its acceptance cone Θ_ac is trapped within a core due to the total internal reflection at the core–cladding interface.

Fig. 39.5 Laser photocoagulation on a slit-lamp system. 1, Optical fiber and electronic cable connecting laser with a slit-lamp system; 2, optical coupler projecting the beam exiting from the fiber on to the retina; 3, contact lens.

Fig. 39.6 Intraocular hand pieces and optical fiber for vitreoretinal dissection with the Er:YAG laser.

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^5,6 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, 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.⁶ 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. The use of such lenses has many advantages, in particular for providing 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. A number of contact lenses have been developed for this purpose, and the most common types are listed in Table 39.1. 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 5°, 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 within ±5° of perpendicular to the laser beam (Fig. 39.5). 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.

Table 39.1 Ocular contact lenses and their magnifications in a human eye

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 39.1. It is important to keep in mind that magnification of the retinal image demagnifies the beam size on the retina by the same amount: the higher the magnification of the retina, the smaller the laser spot on the retina.

Interactions of light with tissue

In linear interactions the irradiance (or power fluence) I(z) (W/cm²) of the beam propagating inside tissue decreases exponentially with depth (Beer’s law) due to absorption and scattering of light: I(z) = (1 – k_s)I₀exp(–µz), where I₀ 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 Figure 39.1, 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.^7,8

Photochemical interactions

Photochemical interactions are based upon nonthermal light-induced chemical reactions. The best-known 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/cm²) 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 targeted cell structures.

The concept of treating 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.^9,10 Closure of subretinal neovascularization using PDT with rose Bengal was demonstrated a few years later.¹¹ Following upon the successful initial results with phthalocyanine, liposomal benzoporphyrin derivative, which selectively attaches to the endothelium of new blood vessels was developed (verteporfin, or Visudyne).^12,13

Clinical indication: photodynamic therapy for subfoveal choroidal neovascularization

PDT for subfoveal choroidal neovascularization (CNV) is performed as a two-step procedure. First, a specialized spectrally adapted photosensitizer, such as verteporfin, is injected intravenously into the patient. 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 physicochemical basis for selectivity of PDT against neovascularization compared with the normal choroidal and retinal vasculature. The presence of increased numbers of new vessels in tumors has also led to the use of PDT as a treatment for several different ocular and nonocular tumors.

Upon absorption of a photon at the proper wavelength the porphyrin molecule undergoes an electronic transition from the ground state into the singlet excited state (¹P*). The singlet excited porphyrin can decay back to the ground state with release of energy in the form of fluorescence – enabling identification of tumor tissue (Fig. 39.7). The singlet can also be converted into the triplet excited state (³P*) 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. Energy transfer therefore takes place, producing toxic singlet oxygen (¹O₂) from ground-state dioxygen (³O₂). Singlet oxygen is very reactive and thus it has a very short diffusion pathlength – less than 0.02 µm – so all its interactions are highly localized. The photochemical processes and subsequent reactions are complex and different for different sensitizers, and are also subject to the microenvironment. Singlet oxygen and other highly reactive species produced by photosensitizers damage endothelial cells in close proximity to the adsorbed drug, resulting in photothrombosis and temporary closure of the neovascularization.

Fig. 39.7 Simplified diagram (with vibrational levels omitted) of the photoexcitation process in porphyrin and energy transfer to oxygen. Light causes excitation of porphyrin to the singlet excited state (¹P*). It can decay back to the ground state with release of fluorescence or convert into excited state (³P*) which is able to transfer energy to a triplet ground state of oxygen. Such energy transfer results in highly toxic singlet oxygen (¹O₂).

Many photosensitizing dyes have been tested for PDT, including rose Bengal, hematoporphyrin derivatives, lutetium, texafrim, Photofrin, and various benzoporphyrins. Only verteporfin (Visudyne) has been approved for ophthalmic use in the 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. 39.8). This is because of the lower sensitivity of retina to far-red light and its superior penetration into the choroid.¹⁴

Fig. 39.8 Absorption spectrum of verteporfin. Arrow indicates the 689-nm peak typically used for photodynamic therapy.

In PDT treatment with verteporfin activation by laser is typically performed 15–20 minutes 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, irradiating a spot chosen to exceed the dimension of the neovascularization minimally, with light intensity of 600 mW/cm², for 83 seconds, resulting in a total radiant exposure of 50 J/cm².^15,16

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