retinal laser therapy biophysical basis

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INTRODUCTION

Historically, a number o light sources have been utilized orretinal phototherapy, including the sun, various ash lamps, and

lasers. The sun is capable o producing a retinal burn either

accidentally (e.g., in the case o solar eclipse retinopathy) or onpurpose, as demonstrated frst by Meyer-Schwickerath.1

However, its dependence on weather conditions, its constantmotion in the sky, and relatively large angular size (0.52) make

it an impractical method or intraocular therapy.Arc lamps are very bright sources o light used in many thera-

peutic and imaging applications. In such lamps, a high-voltagedischarge initially ionizes gas between an anode and cathode,

creating an electric arc a high-current-density discharge in thelamp. The ions emit light at specifc wavelengths, and the spec-

trum o the plasma emission depends on the types o atoms

involved, their temperatures, and gas pressure. Thus the spec-trum o an arc lamp may have a distinct signature. The xenon arc

lamp was the frst to become widely used or retinal photocoagu-lation because o its strong visible and near-inrared emission,

convenience, and low price. However, because o its large sizeand tendency to produce intense retinal burns, it was replaced in

clinical practice by laser-based systems in the early 1970s.

2

Lasers became the preerred light source or retinal photoco-

agulation due their narrow spectrum, wide selection o wave-lengths, excellent collimation (directionality), high brightness,

and variable pulse duration. The directionality o the laser makes

it easy to manipulate the beam optically beore its introductioninto the eye, and to ocus it into very small spots. Its monochro-

maticity makes it possible to choose a wavelength or selectiveabsorption in specifc tissues o the eye. Adjustable pulse dura-

tion allows limiting the thermal diusion to small distances, thusproducing very precise and selective interactions with minimal

collateral damage.The most widespread medical applications o lasers in medi-

cine have been in ophthalmology. Since the introduction o the

ruby laser more than three decades ago, ophthalmic laser appli-cations have experienced rapid growth with the use o argon,

krypton, argon-pumped dye, Nd:YAG, diode, Er:YAG, excimer,and Ti:sapphire lasers. Lasers have been applied to a wide

variety o slit-lamp-based retinal therapies, as well as to vitreo-retinal surgery, glaucoma, lens capsule opacifcation, and rerac-

tive surgery. These applications are based on dierentmechanisms o lasertissue interactions including photothermal,

photodisruptive, and photochemical interactions. The mostcommon vitreoretinal application is retinal photocoagulation.

Additionally, a number o novel therapies have recently been

Retinal Laser Therapy: Biophysical Basisand ApplicationsDaniel Palanker, Mark S. Blumenkranz

Section 4 Translational Basic Science

introduced and are under active evaluation, including select

retinal pigment epithelium (RPE) treatment and subleththermal therapy.

In the ollowing sections we will describe the underlying pr

ciples o lasertissue interactions and the types o lasers avable and appropriate or various vitreoretinal applications.

Optical properties of the eyeThe relaxed eye has an approximate optical power o 60 D (i

its ocal length is 16.7 mm in air), with the corneal power beiabout 40 D, or two-thirds o the total power.3 Due to ordearrangement o collagen fbrils in the cornea it is highly transp

ent, with transmission above 95% in the spectral range 400900 nm.4 The reractive index o the cornea is n 1.376

0.0005.4 The amount o light reaching the retina is regulated

the pupil size, which can vary between 1.5 and 8 mm. The anrior chamber o the eye, which is located between the cornea a

lens capsule, is flled with a clear liquid the aqueous humwhich has a reractive index n 1.3335. The crystalline lens

the eye, located behind the iris, is composed o specialized crtallin proteins with reractive index o n= 1.401.42. The lens

about 4 mm in thickness and 10 mm in diameter and is enclos

in a tough, thin (515 m), transparent collagenous capsule.the relaxed eye the lens has a power o about 20 D, while in tully accommodated state it can temporarily increase to 33

The vitreous humor, a transparent jelly-like substance flling t

large cavity posterior to the lens, and anterior to the retina, ha reractive index n 1.335.4

Light entering the eye can be reected, scattered, transmittor absorbed. Reected or scattered light contains inormati

that can be used or noninvasive diagnostic purposes. Tabsorption characteristics o ocular tissues are determined by t

chromophores resident within the tissue. In the visible partthe spectrum (400800 nm) these chromophores include:

melanin located in the retinal and iris pigmented epithe

choroid, uvea, and trabecular meshwork; (2) hemoglobin, loca

in the red blood cells; (3) macular xanthophyl, which is locatin the plexiorm layers o the retina, especially near and in macula; (4) rhodopsin and cone photopigments which a

located in the photoreceptors; and (5) lipouscin, located primily in the RPE layer. These pigments are o major importance

absorption o visible light in the retina rom both physioloand pathologic standpoints. The absorption spectra o these p

ments, as well as water and proteins, are illustrated in Figu39.1. In the mid-inrared part o the spectrum (315 m)

major absorber is water, while in ultraviolet (below 250 n

protein absorption is dominant.


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BASICS OF LASERS

The term laser stands or light amplifcation by stimulatedemission o radiation. A beam o light is composed o individual

packets o energy that are called quanta or photons. Each othese photons has a particular energy and direction o travel. The

energy o a quantum o light is proportional to its requency, i.e.,

it is reciprocal o its wavelength. In the presence o a properlyprepared laser material, it is possible or a quantum o light to

trigger the release o other quanta with the same wavelength and

direction o travel. This phenomenon is called stimulated emis-sion, and it is an essential element in lasing. In thermal equilib-rium the energy levels o atoms and molecules are populated

according to the Boltzmann distribution, in which upper levelsare always less populated than the lower levels. Stimulated

emission requires an inverted population o energy levels, suchthat 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 lasingmaterial by an excitation source or pump.

Generally, a laser is composed o three basic components:(1) a material that can store energy to be released by stimulated

emission; (2) a means o replenishing the energy stored in the

lasing material; and (3) some method o retaining a raction othe light emitted by the lasing material to stimulate urther emis-sion. Figure 39.2 schematically illustrates a general confguration

o a laser. An energy source is used to introduce energy into thelasing 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 circu-lates between the two mirrors on either end o the laser cavity,

with a raction o the light escaping through one mirror to ormthe laser beam. The trapped light stimulates emission o new

quanta o light rom the laser material with the same wavelengthand direction as the original quanta. In this way, a laser produces

a beam o light between the two mirrors in which all o thequanta move in phase with one another. This property o light

is called coherence.Coherence is related to synchronization o light in time, or

along the laser beam. The duration o the synchronized emission

rom the laser multiplied by the speed o light is called the coher-ence length o the laser emission. This is the distance along

Fig. 39.1 (A) Absorption coefcients or major chromophores in aspectral range o 0.210 m. (B) Absorption coefcients o the majorocular chromophores in the visible part o the spectrum: 400900 nm.Spectral locations o some o the popular laser lines in this range areindicated above the plots.

10000

diode

805 nm

Krypton

647 nm

Krypton

568 nm

Argon

514 nm

Argon

488 nm

Wavelength, nm

1000

Hb

Hb02

Melanin

Xantrophyll

Water

Proteins

Hb

Hb02

Melanin

1E-4

1E-3

0.01

0.1

1

10

100

1000

10000

1400 500 600 700 800 900

Wavelength, nm

10

100

1000

Excimer

193 nm

Nd: YAG-2nd

532 nm

Nd: YAG

1064 nm

Er: YAG

2.94 m

CO2

10.6 m

Absorptioncoefficient,a

(cm

1)

Absorptioncoefficient,a

(cm

1)

500200

A

B

Fig. 39.2 Laser typically consists o theenergy source (pump), the lasing medium,and the optical cavity with a partiallytransparent ront mirror.

Mirror

100%

Mirror

90%

Pumping energy

Output beamLasting medium


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Fig. 39.3 Laser beam o diameter Docuswith a lens o ocal length fproduces a wawith diameter dand a ocal depth F(seeequations in the text).

D

F

f

d

Fig. 39.4 Optical fber typically consists o core, cladding, and a jacket (buer coatingLight launched into the fber with itsacceptance cone ac is trapped within a codue to the total internal reection at thecorecladding interace.

Jacket (buffer coating)

Cladding

CoreAcceptance

cone ac

which the photons are coherent, or moving in step. To remain

in phase with one another, these quanta must have approxi-mately the same wavelength. Thus temporal coherence is related

to the monochromaticity (or spectral width) o the light emittedrom the laser: the broader the spectrum, the shorter the temp-

oral coherence. A laser may produce one or several discrete

spectral lines in the inrared, visible, or ultraviolet domains, incontrast to conventional light sources (incandescent or arc

lamps) which typically produce noncoherent polychromatic

(broadband or white) light.Collimation (directionality) o the emitted beam is governedby the mirror confguration o the laser cavity. In its simplest

orm, a cavity consists o two mirrors arranged such that lightbounces back and orth, each time passing through the gain

medium. One o the two mirrors, the output coupler, is par-tially transparent, allowing the output beam to exit through

it (Fig. 39.2). The reection coefcient o the output coupler

determines how many times photons are reected back tocirculate inside the cavity beore exiting it. For example, with

a reection coefcient o 0.99, the photon will bounce, onaverage, 99 times beore exiting the cavity. The structure o

the laser cavity determines directionality (collimation) o thelaser beam, which determines its ability to be ocused 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 ashlamps, by electric discharges in the laser medium, by chemical

reactions, by an electron beam, by direct conversion o electric

current into photons in semiconductors, or by light rom otherlasers. Laser pulse durations can vary rom emtoseconds to

infnity. Pulsing techniques used or dierent ranges o pulsedurations include electronic shutters (down to 1 ms), pulsed

ash lamps (typically down to a ew s), Q-switching (down toa ew ns), or mode-locking (down to s).

Laser beam delivery to tissueLaser beams are typically very well collimated. Diracti

causes light waves to spread transversely as they propagate, a

it is thereore impossible to have a perectly collimated beaThe diraction-limited divergence angle o a gaussian beam w

diameter D and wavelength is =

4

D. As an example,

an argon laser emitting a 1-mm-wide beam at 515 nm wav

length, the divergence (hal-angle ) is about 0.66 mrad, i.e., tbeam spreads by 1.3 mm over a distance o 1 meter.

Using a lens or a concave mirror with ocal length f, a la

beam can be ocused to a spot with a diameter df

D=

4

. T

depth o the ocal region is Ff

D=

82

2

(Fig. 39.3). With

f= 25 mm lens the same argon laser beam can be ocused

a spot o 16 m in diameter, having a ocal depth o 820 It must be emphasized, though, that exact defnition o

spot size depends on the beam profle, which varies in vario

confgurations o laser cavities. For therapeutic laser photoagulation such tight ocusing is usually not required, and la

spots typically vary in diameter between 50 and 500 m

various applications.As an alternative to a ree-propagating beam, laser light c

be transported via optical fbers. An optical fber, schematica

shown in Figure 39.4, typically consists o a core, cladding, ajacket. Light is trapped within the core due to total inter

reection at the interace o core with cladding. To satisy contions o total internal reection, the incidence angle o light

the corecladding interace should not exceed the critical an

o total internal reection: sincr=ncore/nclad, where ncore and nare reractive indices o the core and cladding, respectively.

satisy these criteria or total internal reection, the light shou


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be launched within a so-called acceptance cone, and the sine o

its hal-angle ac is defned as the numerical aperture (NA) othe fber: NA n nac core clad= = sin

2 2 . Typically the NA is within

a range o 0.10.2. Optical fbers are oten used or delivery o

laser light to slit-lamp-based systems (Fig. 39.5), and to intraocu-lar surgical probes (Fig. 39.6).

ABERRATIONSWith nonperect ocusing optics, the ocal spot size o the laserbeam is limited not only by diraction, but also by aberrations.

Measurements o optical aberrations in the human eye demon-strate5,6 that or pupillary dilation o up to 3 mm in diameter,

an average emmetropic human eye is optically well corrected,

and the ocal spot is close to the diraction limit. However, orpupils greater than 3 mm in diameter, central aberrations

increase, resulting in increases in the ocal spot size. Peripheralfeld aberrations lead to rapidly increasing blur o the image

with angle o visual feld, strongly limiting the ocusing

Fig. 39.6 Intraocular hand pieces and optical fber or vitreoretinaldissection with the Er:YAG laser.

Table 39.1 Ocular contact lenses and their magnifcationsin a human eye

LensImagemagnifcation

Laser beammagnifcation

Ocular Mainster Standard 0.95 1.05

Ocular Fundus Laser 0.93 1.08

Ocular Mainster Wide Field 0.67 1.50

Ocular Mainster Ultra Mag 0.53 1.90

Ocular Mainster 165 0.51 1.96

Ocular Three Mirror Universal 0.93 1.08

Volk G-3 Gonio Fundus 1.06 0.94

Volk Area Centralis 1.06 0.94

Volk Trans Equator 0.69 1.44

Volk SuperQuad 160 0.5 2.00

Volk QuadrAspheric 0.51 1.97

Volk High Resolution WideField

0.5 2.00

Rodenstock Panundoscope 0.67 1.50

Fig. 39.5 Laser photocoagulation on a slit-lamp system. 1, Optical fberand electronic cable connecting laser with a slit-lamp system; 2, opticalcoupler projecting the beam exiting rom the fber on to the retina;3, contact lens.

capability o laser in the periphery o the retina.6 In retinal

photocoagulation, a at contact lens is typically used to reducethe optical power o the ront surace o the cornea. I the lens

is used properly it aids greatly in controlling peripheral aber-rations during photocoagulation. Other methods include using

an aspheric lens to control optical aberrations in the peripheryduring photocoagulation. The use o such lenses has many

advantages, in particular or providing wide-feld viewing,

although aberrations are difcult to correct over the totality ofelds o interest and additional reections may be introduced

by the lens suraces.

CONTACT LENSES

Currently, retinal laser photocoagulation relies heavily on the

use o contact lenses. A number o contact lenses have beendeveloped or this purpose, and the most common types are

listed in Table 39.1. The universal (Goldmann) three-mirrorcontact lens provides a at ront surace that nearly cancels

the positive reractive power o the ront surace o the cornea.Mirrors at 5, 67, and 73 aid in visualization and photoco-agulation o the periphery and anterior segment. To obtain the

most reproducible results in photocoagulation the operatorshould hold the contact lens so that the at surace is within

5 o perpendicular to the laser beam (Fig. 39.5). The use omirrors in contact lenses helps the operator keep the laser beam

properly aligned to the lens while photocoagulating over alarge feld.

Another useul photocoagulation lens is the inverted image

lens system, typifed by the Rodenstock, Quadraspheric, and


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Mainster photocoagulation lenses. These lenses contain a lens

element in contact with the corneal surace and another posi-tive lens element at a fxed distance rom the cornea. These

systems magniy the spot size on the retina, while increasingthe feld o view, requiring the operator to adjust the power

accordingly. Magnifcation actors o the most common contact

lenses are listed in Table 39.1. It is important to keep in mindthat magnifcation o the retinal image demagnifes the beam

size on the retina by the same amount: the higher the

magnifcation o the retina, the smaller the laser spot onthe retina.

INTERACTIONS OF LIGHT WITH TISSUE

In linear interactions the irradiance (or power uence) I(z) (W/cm2) o the beam propagating inside tissue decreases exponen-

tially with depth (Beers law) due to absorption and scatteringo light: I(z) = (1 ks)I0exp(z), where I0 is the light intensity at

the surace o tissue (z= 0), ks is the specular reection coefcient

at the tissue surace, =a+s is the attenuation coefcient,combined o absorption and scattering components. The pene-

tration depth () o the beam into tissue is defned as a depth atwhich light intensity is reduced by a actor o e: = 1/. Reec-

tion o light rom ocular tissue at normal incidence typically doesnot exceed 2%. As shown in Figure 39.1, absorption o light by

various chromophores in the eye strongly varies with wave-length. Scattering o light is also a very strong unction o the

wavelength: scattering on subwavelength inhomogeneities oreractive index in ocular media (e.g., collagen fbrils) is recipro-

cal to the ourth power o the wavelength: s ~ 4 (Rayleigh

scattering). For example, the scattering coefcient or light at1064 nm is 16 times lower than that at 532 nm. Scattering rom

structures larger than the wavelength (e.g., cellular organelles)is described by Mie theory and has more complex wavelength

dependence and spatial distribution. Scaling o the Mie scatter-ing coefcient with wavelength in various tissues can be approx-

imated as s ~ b

, with b 0.52.7,8

Photochemical interactionsPhotochemical interactions are based upon nonthermallight-induced chemical reactions. The best-known natural pho-

tochemical reactions are photosynthesis in plants and photo-

transduction in photoreceptors. Therapeutic photochemicalinteractions used in photodynamic therapy (PDT) take place at

very low power densities (typically


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Photothermal interactionsTemperature is the governing parameter in all thermal laser

tissue interactions. Depending on the duration and peak valueo the temperature, dierent tissue eects, including necrosis,

coagulation, vaporization, carbonization, and melting, mayoccur.

Heat generation in tissue is determined by the laser parame-ters and optical tissue properties irradiance, exposure time,

and the absorption coefcient, which is a unction o the laser

wavelength. Heat transport is characterized by heat conductivityand heat capacity. Heat eects depend on the type o tissue and

temperature history (values and durations).Absorption o light in tissue leads to heating. I heat diusion

is not taken into account, then at a constant beam intensity the

temperature rise is linear with time: T z tI z t

c

a,( )

( ) =

, where

is tissue density, and c is its heat capacity (c = 4.2 J/g/K,

= 1 g/cm3 or water). To assess whether heat diusion playsa signifcant role during the laser pulse, one should compare

pulse duration with a characteristic time it takes or heat tospread by the distance equal to the zone o initial heat deposi-

tion in tissue. For the heated zone (laser penetration depth) o

length L, the heat diusion time is: = L2

/4, where isthermal diusivity (= 1.4103 cm2/s or water). For example,

or L= 1 m in water the characteristic heat diusion time ist= 1.7 s, while or L= 1 mm the diusion time t= 1.7 s. I

the laser pulse duration is comparable or longer than the char-acteristic diusion time across the light absorption zone, then

proper estimation o the peak temperature in tissue shouldtake heat diusion into account.

Sublethal thermotherapyThere is a growing body o clinical evidence that diabetic macular

edema can be successully treated by pulsed near-inrared diodelaser (810 nm) without producing visible lesions.1721 Using

300-ms bursts o submillisecond micropulses, laser energy is

applied with no visible lesions and no uorescein leakage, as

treatment o classic and occult CNV. Verteporfn has a very

broad absorption spectrum, but only the ar-red peak at688691 nm is typically utilized in clinical practice (Fig. 39.8).

This is because o the lower sensitivity o retina to ar-red lightand its superior penetration into the choroid.14

In PDT treatment with verteporfn activation by laser is typi-

cally perormed 1520 minutes ater the intravenous injection othe dye. A beam o 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 o the neovascularization mini-mally, with light intensity o 600 mW/cm2, or 83 seconds,resulting in a total radiant exposure o 50 J/cm2.15,16 Closure o

abnormal (leaking) blood vessels occurs or approximately 612weeks in most patients. Reperusion is common and multiple

treatments are oten required. Figures 39.9A (pretreatment) and39.9B (1 week posttreatment) showing uorescein angiography

images demonstrate closure o a suboveal CNV membrane ater

PDT.

Fig. 39.9 (A) Fluorescein angiogram o a patient with predominantly occult suboveal choroidal neovascular membrane in let eye. (B) Same eye1 week ollowing photodynamic therapy with verteporfn and intravitreal triamcinolone injection. Note absence o hyperuorescence in the area oprevious neovascularization and subtle darkening o choroid corresponding to area o photodynamic closure o the membrane.

BA

Fig. 39.8 Absorption spectrum o verteporfn. Arrow indicates the

689-nm peak typically used or photodynamic therapy.

Wavelength, nm

300 400 500 600 700 8000

10

20

30

40

50

60

70

80

90

Molarextinctioncoefficient[nM

-1cm-1]


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observed acutely and in subsequent clinical exams. However,

application o higher power can lead to signifcant heat accumu-lation and result in damage to RPE and photoreceptors. 17,22

Avoidance o acute laser-induced retinal damage permits apply-ing this treatment conuently on the target retina with a large

number o small (125 m), densely placed laser spots.19,23 The

absence o damage allows this therapy to be repeated as neces-sary, and makes this technique potentially appealing in applica-

tions proximal to the ovea. However, due to the lack o reliable

dosimetry and observable measures or such subvisibletreatment, this technique is difcult to optimize, leading topotentially broad variability in the outcomes.

NecrosisTemperature rises induce conormational changes in various

proteins, which denature at characteristic rates specifc toprotein species. These thermal processes, which may eventually

lead to cell necrosis, depend on both the temperature and dura-tion o the hyperthermia. Thermal cellular damage in a milli-

second regime can be approximated using the Arrheniusmodel.24,25 It assumes a rate o decline in the concentration o a

critical molecular component or cellular metabolism D(t) with

temperature T(t):

dD t D t AE

R T tdt( ) ( ) exp

( )=

*(1)

where E* and A are the activation energy and rate constant

setting parameters to the process, and R is the gas constant,8.3 J/(K mol). Tissue damage, i.e., decrease in critical molecular

component D(), relative to its initial value D0 over the pulselength is encapsulated in the Arrhenius integral :

( ) ln( )

exp( )

=

=

D

DA

E

R T tdt

0 0

*(2)

The model assumes that irreversible tissue damage takes placewhen concentration o the critical molecular component drops

below some threshold value. Conventionally, this threshold cor-responds to a reduction in concentration by a actor o e, or an

Arrhenius integral o unity. Thus when = 1, the nondenatured

raction o proteins is 1/e 37%, or in other words, 63% o pro-teins have been damaged.

Measurements o the RPE damage at various irradiation con-ditions yields the ollowing average values25: E* = 340 kJ/mol,

A= 1.6 1055. It is important to keep in mind, though, that accu-rate estimation o cell survival under thermal stress is much

more complex than just assessment o the denaturation rate oone type o protein or another. For example: (1) there are mul-

tiple types o proteins in cells and they denature at dierent

rates; (2) dierent proteins have dierent importance or cellularsurvival; (3) cellular repair mechanisms cannot be ignored at

long exposures. Thereore the single values o the reaction rateA and activation energy E most likely represent characteristics

o the weak link in cellular metabolism, most susceptible tothermal damage.

Figure 39.10 shows an example o the temperature rise intissue or a hypothetical square pulse o heating, which is su-

fcient or cell death, according to the Arrhenius model param-eters listed above. Cells exposed to temperatures above the

threshold curve are coagulated and the tissue becomes necrotic.

Fig. 39.10 Solid line depicts Arrhenius approximation o the cellulardamage threshold as a unction o duration o a hypothetical squarepulse o heating. Dashed line indicates deviation o the damagethreshold rom the Arrhenius model at long exposures.

Sublethal thermal stress

Irreversible damage

T

emperature,

C

40

45

50

55

60

65

70

75

80

1.E-04 1.E-03 1.E-02 1.E-01 1.E-00

Time, s

1.E+0

For example, at 1 second long exposure T 50C, while at 10 it requires 67C or a lethal thermal damage. This curve

approximate, and the exact values depend on the shape o tactual pulse o temperature, type o cells, and tissues involve

The Arrhenius model ails to predict correct threshold tempetures at exposures longer than approximately 1 second, sin

cellular repair mechanisms should be taken into account at loexposures.

Transpupillary thermotherapyThe possibility o localized tissue coagulation and necrosis or

the basis o the tumor treatment technique called laser-inducinterstitial thermotherapy (LITT) or transpupillary thermoth

apy (TTT). It has been applied to tumors in retina, brain, prtate, and liver, and is considered a orm o minimally invas

surgery. The concept o LITT is to apply a laser beam to tiss

in such a way that the target tissue is heated or a prolongperiod o time (about 1 minute) to temperatures above t

threshold o necrosis (on the order o 60C). To achieve depenetration into tissue continuous wave lasers in the ne

inrared region (8001064 nm) are typically employed. TTT the treatment o intraocular tumors26 typically requires exposu

times o 1 minute and irradiances varying rom 5 to 12 W/cmThe use o TTT has been tested in the treatment o CNV

AMD.15,27,28 Proponents o this approach have hypothesized

selective eect on the heating o actively dividing cells in newormed blood vessels due to their higher susceptibility to therm

injury than nondividing cells in normal tissue. The estimatretinal temperature elevation with TTT at standard settin

(810 nm, 800 mW, 60 seconds, 3.0 mm spot size) is appromately 10C.29,30 The mechanism o treatment o CNV by T

may occur through vascular thrombosis, apoptosis, or thermal inhibition o angiogenesis.31

PhotocoagulationRetinal photocoagulation typically involves the application

laser pulses with durations ranging rom 10 to 200 ms, and trsient hyperthermia by tens o degrees above body temperatu


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Fig. 39.11 Histology o the rabbit retina at1 day (let column) and 4 months (rightcolumn) ater photocoagulation. Retinal spotsize is 330 m, power 175 mW. (A) Intenseretinal burn produced with 100 ms exposure.

Yellow bar shows the lateral extent o thelesion. Note ull-thickness retinal injury,including the inner retinal layers. (B) Lightburn produced with 15 ms exposure.Photoreceptors are coagulated, while innerretina is well preserved. (C) Barely visiblelesion produced with 7 ms pulse. Rightcolumn shows corresponding retinal scarringat 4 months. Note complete closure o thedamage zone by shiting photoreceptors inthe barely visible lesion.

A

B

C

1 day100 m

4 months

Various lasers have been used in the past (ruby (694 nm), argon

(488, 514 nm), krypton (647 nm)). Currently the most commonlasers in photocoagulation are requency-doubled Nd:YAG

(532 nm), and yellow semiconductor laser (577 nm). The laserenergy is absorbed primarily by melanin in the RPE and choroid,

and by hemoglobin in blood. At a 532-nm wavelength approxi-

mately hal o the laser energy incident on the retina is absorbedin the RPE, and the rest in the choroid.25 The heat generated di-

uses rom the RPE and choroid into the retina and causes coagu-

lation o the photoreceptors and, sometimes, o the inner retina.During 100-ms applications, the heat diuses distances o up to200 m, thus smoothing the edge and extending the coagu-

lated zone beyond the boundaries o the laser spot, termedthermal blooming. Heat diusion using shorter pulses and

with smaller spot sizes can be limited to the photoreceptor layer,thereby avoiding the inner retinal damage.

The let panel in Figure 39.11A demonstrates the acute eects

o an intense burn in a rabbit retina produced by 100 ms laserapplications, including ull-thickness injury and early necrotic

eatures 24 hours ater treatment. The let panel in Figure 39.11Bdemonstrates a light lesion produced by 15 ms pulses. Damaged

photoreceptors are pyknotic, but the inner nuclear layer andganglion cell layer are very well preserved.

Figure 39.12 illustrates the eect o laser power and pulseduration on the size o the coagulated zone in pigmented

rabbits.32Table 39.2 lists the ratio o the lesion width to the retinalbeam size or lesions o various clinical grades in human patients,

as measured by optical coherence tomography (OCT) within 1

hour o treatment.33 As can be seen, lesion size increases relativeto the beam width with more intense lesions and longer pulses.

The threshold power required or the creation o retinal lesions

increases with shorter pulses, since higher temperatures arerequired or coagulation during shorter exposures. An example

o the threshold powers or lesions o various grades is plottedin Figure 39.13 as a unction o pulse duration or a 132 m

retinal laser spot in the rabbit. A relatively modest power

increase is required to produce comparable lesion grades goingrom 100 to 10 ms, whereas a much steeper increase is seen or

durations under 10 ms. For pulse durations o 20, 50, and 100 ms,

all the grades (mild, moderate, intense, very intense, and rupture)could be created with appropriate choice o power settings. Atpulse durations below 10 ms, it became increasingly difcult to

create intense lesions reproducibly without inadvertently rup-turing the retina. At 2 ms or less it was not possible to create a

moderate lesion reproducibly without rupturing the retina. At1 ms, there was little or no dierence between the power required

to create a mild retinal lesion or produce a rupture.

The ratio o the threshold power required to produce a ruptureto that required to produce a mild lesion is defned as the thera-

peutic window, and represents one means o quantiying therelative saety (dynamic range) o retinal photocoagulation. The

larger this ratio, the greater the margin o saety to create avisible lesion without inadvertently inducing a retinal rupture.

Figure 39.14 depicts the width o this therapeutic window as aunction o pulse duration or two dierent laser spot sizes. For

a 132 m retinal laser beam size, as pulse duration decreasesrom 100 to 20 ms, the width o the therapeutic window declines

rom 3.9 to 3.0. When pulse duration is urther decreased to

10 ms, the therapeutic window decreases urther to 2.5, and itapproaches unity at a pulse durations o 1 ms. For a 330 m


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Fig. 39.12 (A) Retinal lesions in the rabbit eye with variable power and duration o exposure. Retinal beam size is 132 m. (B) Lesion diameteas a unction o laser power and duration o exposure. Laser beam size on the retina is 132 m (indicated by dashed line and arrow).

250 mW

200 mW

150 mW

100 mW

50 mW

100908070605040

Pulse duration, ms

3020100

500

400

300

200

100

0

Burndiameter,m

BPulse duration, ms

10 20 50 100

Power

A

Table 39.2 Ratio o the lesion width to retinal beam size or various pulse durations and clinical grades, as measured by optical coherencetomography in human patients within 1 hour o application. Coagulation was perormed with Area Centralis lens (laser beam magnifcation0.94)

Beam size Lesion clinical grade

In air On retina Moderate Light Barely visible

100 ms 20 ms 100 ms 20 ms 100 ms 20 ms

100 m 94 m 3.81 0.98 2.50 0.30 2.08 0.24

200 m 188 m 2.08 0.22 1.49 0.09 1.24 0.08 0.93 0.08

400 m 376 m 1.39 0.08 1.15 0.07 1.19 0.11 0.99 0.09 0.99 0.08 0.74 0.12

Fig. 39.13 Threshold power o retinal photocoagulation in rabbit eye,as a unction o pulse duration. Laser beam size on the retina is132 m. Clinical grades indicated by the colors: light, moderate,intense, very intense, and rupture.

Power,mW

Light

Moderate

Intense

Very intense

Rupture

Pulse duration, ms

1001010

50

100

150

200

250

300

350

400

Fig. 39.14 Sae therapeutic window o retinal photocoagulation (ratioo the threshold o rupture to that o light coagulation) increases withpulse duration, and with a beam size on the retina (shown or 132 a330 m).

5

4

3

2

1

1 10 10

132 um

330 m

Pulse duration, ms

Safe

Unsafe


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Patterns included square arrays with up to 5 5 spots, arcs

with the number o concentric rows varying rom 1 to 3, andcircular patterns or photocoagulation o small holes and other

lesions in the retinal periphery. Patterns or macular photoco-agulation included rings and arcs with an adjustable central

exclusion zone o up to 2 mm in diameter to allow or laser

application, reducing the risk o inadvertent damage to theoveal avascular zone.

To deliver the whole pattern within the eye fxation time and

avoid beam movement due to the ocular muscles, each exposurewas required to be shorter than in conventional photocoagula-tion: 1020 ms instead o 100200 ms, traditionally applied with

single spot exposures. Reduced heat diusion into choroidduring shorter exposures also resulted in patients experiencing

less pain.3941 Short pulse lesions appear smaller and lighter thanconventional burns produced with the same beam size, and

thereore a larger number o them are required to treat the same

total area.33

An automatic laser delivery, guided by diagnostic imaging

and stabilized using eye tracking, has been recently introducedin a Navilase system (OD-OS). This system includes retinal

image acquisition, annotation o the images to create a detailedtreatment plan, and then automated delivery o the laser to the

retina according to the treatment plan.

Clinical indications: treatment of diabetic retinopathy

Photocoagulation has proven sae and eective in the treatmento prolierative diabetic retinopathy. In this disorder the retina

becomes ischemic and releases a variety o chemical messengers,most importantly, vascular endothelial growth actor (VEGF),

that stimulate the growth o new blood vessels and also mark-edly increase retinal vascular permeability. The abnormal new

vessels, associated fbrous tissue, and macular edema are major

causes o the sight-threatening complications in diabetic eyedisease. By destroying a portion o the peripheral retina with

laser, it has been hypothesized that retinal metabolic demandsand available nutrients are better balanced and the stimulus or

growth o the new blood vessels is decreased. This treatment hasbeen termed panretinal photocoagulation (Fig. 39.15) and signif-

cantly reduces the risk o vision loss due to neovascularization.

Fig. 39.15 Fundus photograph o a patient 1 week ollowingapplication o panretinal photocoagulation with argon laser.

retinal laser spot size the therapeutic window declines rom 5.4

to 3.7 to 3.1 when pulse durations decrease rom 100 to 20 to10 ms, respectively. With both spot sizes, the therapeutic window

decreases to unity as pulse durations decrease to 1 ms. At thispoint there is eectively no sae range o retinal photocoagula-

tion: mild lesion and rupture are equally likely to occur at the

same power.The width o the sae therapeutic window should sufce to

accommodate or variations in undus pigmentation, which typi-

cally do not exceed a actor o 2. To provide a sae therapeuticwindow larger than 2.5, pulse durations should equal or preer-ably exceed 10 ms or a beam o 330 m, and 20 ms or the

132 m spot size.It is important to keep in mind that coagulation o blood

vessels requires more energy than other tissue due to cooling bythe blood ow. For example, i a spot size o 200 m with expo-

sure time o 200 ms is applied to occlude a blood vessel with

ow velocity o 5 mm/s, the laser energy is eectively distrib-uted over the column fve times longer than the diameter o the

laser spot. Thus the eective energy remaining at the photoco-agulation site is fve times lower than it would be in stationary

tissue.

Healing of retinal lesionsStudies in rabbits demonstrate that in photocoagulation lesionsthe RPE layer is restored within a week, though its pigmentation

may remain abnormal either hyper- or hypopigmented.34 Inintense and moderate lesions gliotic scar flling the coagulated

retinal layers stabilizes ater 1 month, and the wound contracts

to approximately 40% o the original lesion diameter, as shownin the right column in Figure 39.11A and B. However, in very

light lesions (barely visible clinical grade), photoreceptors con-tinue to shit into the damage zone and completely refll it by 4

months, as shown in Figure 39.11C, in the right panel. As aresult, scarring and scotomata typically associated with conven-

tional photocoagulation may be minimized or even completely

avoided.

34

A similar phenomenon o restorative retinal plasticityhas recently been observed in rats35 and primates.36 It is impor-tant to keep in mind, though, that in order to maintain clinical

efcacy o photocoagulation with smaller and lighter lesions, a

larger number o them should be applied, to keep the same totalcoagulated area.33

Pattern-scanning laser photocoagulationThe frst attempts to make photocoagulation a completely auto-

mated procedure involved rather complex equipment, includingimage recognition sotware and eye tracking.37 The complexity

o such systems prevented their commercial introduction andacceptance in clinical practice.

A semiautomatic pattern-scanning photocoagulator (PASCAL,

Topcon Medical Laser Systems) was introduced by OptiMedicain 2005.38 It delivered patterns o laser spots, ranging rom a

single spot to 56 spots applied in a rapid sequence with a singledepression o a oot pedal. The control o laser parameters was

perormed by means o a touch screen graphic user interace,acilitating selection o the dierent patterns o photocoagula-

tion. The laser was activated by pressing a oot pedal, which waskept depressed until the entire pattern was completed, although

it is possible or the physician to release the oot pedal and stopthe laser at will, prior to completion o the pattern, i clinically

indicated.


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Absorption o these wavelengths in macular pigments m

cause heating and destruction o the nerve fber layer, resultiin loss o vision. As shown inFigure 39.1B, in the macular regi

wavelengths longer than 500 nm should be chosen, such as tgreen argon (514 nm) or the requency-doubled YAG (532 n

or semiconductor yellow (577 nm) laser. Melanin provides go

absorption at most photocoagulation wavelengths. Wavelengselection is thereore less important when melanin is the prima

absorber. To minimize scattering loss in cataract or in vitreo

opacities the longer wavelengths (yellow, 577 nm, or r640680 nm) are efcacious. I scattering by the ocular tissuesnot signifcant, the argon green or doubled YAG continues

serve well.When hemoglobin is the primary absorber (Fig. 39.1B), as

the treatment o vascularized tumors, a wavelength shorter th600 nm is preerable. Treatment o CNV may be eective usi

red light through indirect heat transer rom the surroundi

melanin. In general, when photocoagulating structures conta large quantity o hemoglobin, wavelengths between 520 a

580 nm are best suited. Ideally, or coagulation o blood vessthe photon penetration depth should be similar to the ves

diameter, thus providing uniorm heating o the blood veswithout superfcial damage and peroration.

Tunable lasers may provide the exibility to select a walength o choice or required photothermal procedure. Howev

tunable lasers are more costly, require more maintenance, aare now less commonly employed clinically than previously.

PhotodisruptionWhen tissue temperature exceeds the vaporization threshovapor bubbles are produced, which may lead to rupture o t

tissue within a zone comparable to the bubble size. This proco explosive vaporization is typically employed or tissue diss

tion. The actual temperature required or vaporization varbetween 100 and 305C depending on pulse duration and

presence o the bubble nucleation sites.44 For efcient heatingtissue the pulse energy should be delivered ast enough to avo

heat diusion rom the laser absorption zone during the pul

a condition called thermal confnement. In other words, laser pulse duration should be shorter than the heat relaxati

time, or the time o heat diusion rom the zone o laser absotion, L:


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to the tissue surace may produce a water jet which is capable o

damaging tissue at distances exceeding the bubble radius by aactor o 4.46,47

Explosive vaporization can be produced by absorption o laserradiation in water or in tissue. Strong absorption in water can

be achieved at mid-inrared wavelengths. For example, penetra-

tion o the Er:YAG laser (= 2.94 mm) in water is about 1 m.Shallow penetration o these lasers necessitates the fber-based

delivery o this light into a liquid medium. A thin layer o water

in ront o the intraocular probe is overheated with the laserpulse, and the resulting vaporization leads to rupture o tissuein proximity to the probe. The best implementation o this prin-

ciple has been achieved with Er:YAG-based dissection o epireti-nal membranes48,49 (Fig. 39.18). Since a burst o closely spaced

pulses, rather than a single pulse, was applied in that device, theactual vapor bubble had an elliptical shape and extended several

hundreds o microns rom the probe.50

Alternatively, overheating o liquid can be achieved with alaser strongly absorbed in the tissue constituents. For example,

a fber-delivered ArF excimer laser ( = 193 nm), which isstrongly absorbed by proteins (penetration depth 0.2 m in

generated which propagate with supersonic velocities and may

result in signifcant damage to tissue, such as disruption andragmentation.44

Vaporization o water in an overheated volume results in theormation o a short-living vapor bubble (a so-called cavitation

bubble) which expands, cools down, and collapses during the

time determined by its radius at maximal expansion. The lietimeo the spherical cavity with radius R0 in ree nonviscous liquid

is described by the Rayleigh equation:

= 0 91 0

0

. RP

, where

is liquid density and P0 is ambient pressure.45 For example, in

water at atmospheric pressure, a cavity o 0.1 mm in radius col-

lapses in about 10 s. (Growth and collapse o a spherical bubbletake approximately the same time.) Symmetric collapse o a

spherical cavity may lead to overheating o the liquid and orma-tion o the secondary bubble. Due to the short lietime these

bubbles are not visible to the eye during surgery, but they can

be easily visualized using ast ash photography. As shown inFigure 39.17, at the liquidtissue interace the bubble might be

deormed, and the secondary bubble may not be created. Impor-tantly, a collapsing cavitation bubble at the fber probe or next

Fig. 39.17 Dynamics o the cavitation bubble at the gelatinsaline interace observed with ast ash photography. Bubble is created by a pulse oArF excimer laser delivered via the tapered optical fber. Numbers in each rame indicate a delay in microseconds between the laser pulse(10 ns) and a microsecond-long ash.


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clinical studies have shown the efcacy o SRT in diabetic mac

lopathy, central serous choroidopathy,62,63 and suboveal uater rhegmatogenous retinal detachment.64,65 Despite its clini

promise, this technique has not been commercialized. One o tdifculties with SRT is the lack o visible change in the retin

appearance, making it difcult to assess adequate laser do

metry as the applications are being placed. An acousto-optisystem is currently under development that may help to ass

the threshold energy density required or cavitation in RPE.66

alternative approach to SRT is the rapid scanning o a laser be

providing microsecond exposures, sufciently short or selecttreatment o the RPE.67,68

FUTURE DEVELOPMENTS

Monitoring retinal temperatureRetinal thermal therapies with temperature rise below t

threshold o immediately visible tissue response, such as sulethal hyperthermia, do not have as high a degree o predi

ability as conventional thermal photocoagulation methods. Fsuccessul sublethal treatment the temperature rise in tissue

on the order o 10C.69 However, due to the strong variationundus pigmentation the light absorption varies rom patient

patient, and even between dierent areas in the same eye. The

ore the same irradiation settings may lead to very diereresults in dierent patients, and so direct measurement o reti

temperature during such treatments is highly desirable. Simlarly, it would be desirable to monitor retinal temperature in

treated spot during photocoagulation to provide uniorm ocomes in areas with dierent pigmentation.

A noninvasive method o determination o retinal temperatuhas been developed, which is based on the detection o acous

waves generated in RPE by short laser pulses.70 An acoustransducer or the detection o the pressure waves is built int

contact lens attached to the treated eye during the procedu

The pressure waves are generated due to thermoelastic expsion o melanosomes upon absorption o the short (submicros

ond) laser pulses. The key issue in this approach is that thermoelastic expansion coefcient o water var

tissue), has been applied or dissection o epiretinal and subreti-nal membranes.51,52 In this case the laser light overheats the tissue

and leads to vaporization o its water content with subsequent

tissue rupturing.53,54 Despite the early promise o both o thesedevices (Er:YAG and ArF excimer lasers) in clinical tests, both

have ailed to achieve widespread acceptance in medical practicedue in part to cost, the rigidity o the fbers, and the lack o

coagulation capability.Another approach to dissection o transparent tissue utilizes

ionization o the material and ormation o plasma rom a high-intensity laser beam. At extremely high irradiances (108-1011 W/

cm2), that can be achieved in a short-pulsed (nss) tightlyocused laser beam, transparent material can be ionized and ions

absorbing the laser light reach very high temperatures.55 This

mechanism, called dielectric breakdown, allows or a very local-ized deposition o energy in the middle o a transparent liquid

or solid at the ocal point o the laser beam. This process iswidely used in ragmentation o the opacifed posterior lens

capsule (secondary cataract) with nanosecond Nd:YAG lasers.At shorter pulse durations (1 ps100 s) and lower energies, this

process is applied to intrastromal ablation ormation o acorneal ap or reractive surgery.56,57 This approach has also

been tested in the dissection o epiretinal membranes using the

tightly ocused beam directed rom outside the eye.58 Despite theact that very low energy is required or this process (several

microjoules with ps lasers), its applicability in the posterior poleis limited due to the difculty in axial dierentiation between

the epiretinal membranes and the retina located very closebehind them. In addition, strong optical aberrations in the

periphery o the posterior pole preclude tight ocusing o thelaser beam in these areas.

Selective retina therapy (SRT)Light is strongly absorbed in the melanosomes in the RPE (a

8000 cm-1).59 The application o short (submicrosecond) laserpulses allows or confnement o the thermal and mechanical

eects o this absorption within the RPE layer, thus sparing the

photoreceptors and the inner retina60,61 (Fig. 39.19). It has beendemonstrated that the application o repetitive pulses o micro-

second and submicrosecond duration results in selective damageto RPE, presumably due to the ormation o small cavitation

bubbles around melanosomes.59 Subsequent RPE prolierationand migration restore continuity o the RPE layer. Several small

Fig. 39.18 Epiretinal membrane dissected by the Er:YAG laser. Notethe absence o retinal damage despite its close proximity.

Fig. 39.19 Rabbit retina 24 hours ater selective retina therapy withinrared semiconductor laser. Note that damage is almost exclusivelyconfned to the retinal pigment epithelial (RPE) layer with preservatioo the outer nuclear layer and even the outer segments o overlyingphotoreceptors. There is a small localized eusion between the RPEand photoreceptors.


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neovascularization-verteporfn in photodynamic therapy report 2. Am JOphthalmol 2001;131:54160.

17. Luttrull JK, Musch DC, Mainster MA. Subthreshold diode micropulse photo-coagulation or the treatment o clinically signifcant diabetic macular oedema.Br J Ophthalmol 2005;89:7480.

18. Figueira J, Khan J, Nunes S, et al. Prospective randomised controlled trialcomparing sub-threshold micropulse diode laser photocoagulation andconventional green laser or clinically signifcant diabetic macular oedema.Br J Ophthalmol 2009;93:13414.

19. Luttrull JK, Musch DC, Mainster MA. Subthreshold diode micropulse photo-coagulation or the treatment o clinically signifcant diabetic macular oedema.Br J Ophthalmol 2005;89:7480.

20. Luttrull JK, Musch DC, Spink CA. Subthreshold diode micropulse panretinalphotocoagulation or prolierative diabetic retinopathy. Eye (Lond) 2008;22:60712.

21. Ohkoshi K, Yamaguchi T. Subthreshold micropulse diode laser photocoagula-tion or diabetic macular edema in Japanese patients. Am J Ophthalmol 2010;149:1339.

22. Desmettre TJ, Mordon SR, Buzawa DM, et al. Micropulse and continuouswave diode retinal photocoagulation: visible and subvisible lesion parameters.Br J Ophthalmol 2006;90:70912.

23. Luttrull JK, Sramek C, Palanker D, et al. Long-term saety, high-resolutionimaging, and tissue temperature modeling o sub-visible diode micropulsephotocoagulation or retinovascular macular edema. Retina 2011;epub aheado print.

24. Niemz M. Lasertissue interactions. Fundamentals and applications. Berlin:Springer; 2002.

25. Sramek C, Paulus Y, Nomoto H, et al. Dynamics o retinal photocoagulationand rupture. J Biomed Optics 2009;14:03400713.

26. Oosterhuis JA, Journeedekorver HG, Kakebeekekemme HM. Transpupillarythermotherapy in choroidal melanomas. Arch Ophthalmol 1995;113:693693.

27. Reichel E, Berrocal AM, Ip M, et al. Transpupillary thermotherapy o occultsuboveal choroidal neovascularization in patients with age-related macular

degeneration. Ophthalmology 1999;106:190814.28. Newsom RSB, McAlister JC, Saeed M, et al. Transpupillary thermotherapy

(TTT) or the treatment o choroidal neovascularisation. Br J Ophthalmol2001;85:1738.

29. Svaasand LO. Laser-induced hyperthermia physics considerations andlimitations. Laser Surg Med 1988;8:182182.

30. Svaasand LO, Gomer CJ, Profo AE. Laser-induced hyperthermia o oculartumors. Appl Optics 1989;28:22807.

31. Mainster MA, Reichel E. Transpupillary thermotherapy or age-related maculardegeneration: Long-pulse photocoagulation, apoptosis, and heat shock pro-teins. Ophthalmic Surg Las 2000;31:35973.

32. Jain A, Blumenkranz MS, Paulus Y, et al. Eect o pulse duration on size andcharacter o the lesion in retinal photocoagulation. Arch Ophthalmol2008;126:7885.

33. Palanker D, Lavinsky D, Blumenkranz MS, et al. The impact o pulse durationand burn grade on size o retinal photocoagulation lesion: implications orpattern density. Retina 2011;31:16649.

34. Paulus YM, Jain A, Gariano RF, et al. Healing o retinal photocoagulationlesions. Invest Ophthalmol Vis Sci 2008;49:55405.

35. Belokopytov M, Belkin M, Dubinsky G, et al. Development and recovery olaser-induced retinal lesion in rats. Retina 2010;30:66270.36. Merigan WH, Strazzeri J, DiLoreto DA, Jr, et al. Visual recovery ater outer

retinal damage in the macaque. Invest Ophthalmol Vis Sci 2011;52:3202.37. Wright CHG, Barrett SF, Ferguson RD, et al. Initial in vivo results o a hybrid

retinal photocoagulation system. J Biomed Opt 2000;5:5661.38. Blumenkranz MS, Yellachich D, Andersen DE, et al. Semiautomated patterned

scanning laser or retinal photocoagulation. Retina 2006;26:3706.39. Nagpal M, Marlecha S, Nagpal K. Comparison o laser photocoagulation or

diabetic retinopathy using 532-nm standard laser versus multispot pattern scanlaser. Retina 2010;30:4528.

40. Muqit MM, Marcellino GR, Gray JC, et al. Pain responses o Pascal 20 ms multi-spot and 100 ms single-spot panretinal photocoagulation: Manchester PascalStudy, MAPASS report 2. Br J Ophthalmol 2010;94:14938.

41. Muqit MM, Gray JC, Marcellino GR, et al. In vivo lasertissue interactions andhealing responses rom 20- vs 100-millisecond pulse Pascal photocoagulationburns. Arch Ophthalmol 2010;128:44855.

42. DRS Study Group. Photocoagulation treatment or prolierative diabeticretinopathy. clinical application o DRS fndings, DRS report number 8.Ophthalmology 1981;88:583600.

43. ETDRS Study Group. Early Photocoagulation or Diabetic Retinopathy. ETDRSreport number 9. Ophthalmology 1991;98:76685.

44. Vogel A, Venugopalan V. Mechanisms o pulsed laser ablation o biologicaltissues. Chem Rev 2003;103:20792079.

45. Young FR. Cavitation. Maidenhead: McGraw-Hill; 1989. p. 136.46. Palanker D, Vankov A, Miller J. Eect o the probe geometry on dynamics o

cavitation. SPIE, LaserTissue Interaction XIII 2002;4617:1127.47. Palanker D, Vankov A, Miller J, et al. Prevention o tissue damage by water jet

during cavitation. J Appl Phys 2003;94:265461.48. DAmico DJ, Blumenkranz MS, Lavin MJ, et al. Multicenter clinical experience

using an erbium:YAG laser or vitreoretinal surgery. Ophthalmology1996;103:157585.

49. DAmico DJ, Brazitikos PD, Marcellino GR, et al. Initial clinical experience withan erbium:YAG laser or vitreoretinal surgery. Am J Ophthalmol 1996;121:41425.

with temperature, by about 1% per 1C.70 This eect allows or

monitoring the changes in temperature o the RPE cells by moni-toring the changes in amplitude o acoustic waves generated by

the laser pulses o constant energy. The probing laser pulses areapplied simultaneously with application o a therapeutic laser

to detect temperature rise in tissue during the exposure. It has

been demonstrated that precision o this method is on the ordero 1C. Clinical testing o the system is currently in progress.

Optical monitoring of tissue changes

in real timeAn optical approach to real-time eedback during retinal photo-

coagulation has recently been demonstrated.71 It is based on OCTmonitoring o the heated tissue expansion and changes in retinal

scattering during coagulation. The system operates with ms tem-poral resolution, which should be ast enough or real-time

monitoring o retinal photocoagulation.Another technique or detection o tissue condition during

slow thermal therapy is based on spectroscopy o white lightscattered rom the tissue.72 Cellular response to thermal stress

involves expression o various proteins, as well as changes in

their aggregation and concentration. All these eects result inchanges o the reractive indices and/or the sizes and shapes o

the cellular organelles, which can be detected using light-scattering spectroscopy. Particle sizes down to 100 nm in diam-

eter can be detected using light within the spectral range o3501000 nm.72 Since the inormation is obtained optically and

without any staining this technique operates in real time and isnoninvasive. It has been observed that scattering coefcients o

some organelles change very strongly (up to 70%) and rapidly(20 seconds) in the heated cells.73

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4. Thompson KP, Ren QS, Parel J-M. Therapeutic and diagnostic application olasers in ophthalmology. In: Waynant RW, editor. Lasers in medicine.Boca Raton FL: CRC Press; 2002, chapter 8.

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8. Troy TL, Thennadil SN. Optical properties o human skin in the near inraredwavelength range o 1000 to 2200 nm. J Biomed Opt 2001;6:16776.

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16. Arnold J, Kilmartin D, Olson J, et al. Verteporfn therapy o suboveal choroidalneovascularization in age-related macular degeneration: Two-year results o arandomized clinical trial including lesions with occult with no classic choroidal


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63. Elsner H, Porksen E, Klatt C, et al. Selective retina therapy in patients wcentral serous chorioretinopathy. Graees Arch Clin Exp Ophthal2006;244:163845.

64. Koinzer S, Elsner H, Klatt C, et al. Selective retina therapy (SRT) o chrosuboveal uid ater surgery o rhegmatogenous retinal detachment: three creports. Graees Arch Clin Exp Ophthalmol 2008;246:13738.

65. Brinkmann R, Roider J, Birngruber R. Selective retina therapy (SRT): a revon methods, techniques, preclinical and frst clinical results. Bull Soc BeOphtalmol 2006;5169.

66. Brinkmann R, Schuele G, Joachimmeyer E, et al. Determination o absoundus temperatures during retinal laser photocoagulation and selective Rtreatment. Invest Ophthalmol Vis Sci 2001;42:S696S696.

67. Framme C, Alt C, Schnell S, et al. Selective targeting o the retinal pigmepithelium in rabbit eyes with a scanning laser beam. Invest Ophthalmol Sci 2007;48:178292.

68. Paulus YM, Jain A, Nomoto H, et al. Selective retinal therapy with microsecexposures using a continuous line scanning laser. Retina 2011. 31:3808.

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retinal laser therapy biophysical basis

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