calibrated real-time control of lesion size based on reflectance images

Calibrated real-time control of lesion size basedon reflectance images

Maya R. Jerath, Danielle Kaisig, H. Grady Rylander 111, and Ashley J. Welch

Lesion size induced by laser photocoagulation is controlled in real time based on a two-dimensionalreflectance image recorded by a CCD array during lesion formation. A feedback system usingcomponents of the reflectance image achieves uniform lesions by compensating for light absorptionvariability in biological media. Lesions are formed in a phantom by an argon laser to simulate retinalphotocoagulation. The tissue model consists of a thin absorptive layer covered by a clear albuminprotein layer. Results show a low variance in the sizes of the lesions (diameter or depth) produced indifferent irradiation conditions and the ability to produce lesions of a predefined size in varyingillumination conditions.

Key words: Laser, photocoagulation, feedback, reflectance.

Introduction

Laser-induced retinal lesions find many therapeuticapplications in ophthalmology.'-3 The lesions are aresult of thermal damage induced in tissue by theabsorbed laser light. Photocoagulation of the retinaby an argon laser is used to treat diseases rangingfrom retinal detachment to diabetic retinopathy.Understandably this thermal damage has therapeuticvalue only if the lesions are carefully controlled.

A lesion is a region of thermally damaged tissue.The extent of photocoagulation is dependent on theheat generated in the tissue by the conversion of laserlight energy to thermal energy when the light isabsorbed in tissue. The amount of light absorbed atany given wavelength is a function of the tissue'swavelength-dependent optical characteristics (suchas absorption and scattering coefficients) and of thelaser exposure parameters (such as power, spot size,and exposure time). Even when the laser parame-ters at the cornea are known, neither tissue opticalproperties nor the amount of laser radiation reachingthe retina are known a priori. The primary ab-sorber of argon laser light in the fundus is thepigment epithelium, and it has been documented byGabel et al. 4 that local absorption in the retina canvary by a factor of 2. Thus absorption varies notonly from fundus to fundus but even within the same

The authors are with the Biomedical Engineering Program, TheUniversity of Texas at Austin, Austin, Texas 78712.

Received 26 December 1991.0003-6935/93/071200-10$05.00/0.© 1993 Optical Society of America.

fundus. It has been shown through computer simu-lations that when any one of the exposure or tissueparameters is changed, even if all the rest of theparameters are kept constant, lesions vary greatly insize.5 Thus it is impossible to predict accuratelylesion depth or diameter.

The boundary of a lesion is a biological marker ofthe extent of significant heat flow and temperaturerise. This thermally damaged region grows bothradially and axially from the location of the generatedheat source, i.e., the pigment epithelium. Radialgrowth (perpendicular to the beam axis) gives thelesion a certain diameter, and axial growth (along thebeam axis) gives it a certain depth (see Fig. 1). Thesize of the lesion, both radially and axially, is ofcritical concern. Lesions are therapeutic only if theydestroy targeted tissue and introduce no complica-tions of their own. Since absolute size cannot bepredicted ideally, the lesion should be monitored as itforms. When it reaches the preset size, in a keydimension such as diameter or depth, the laser wouldbe shut off. Unfortunately this real-time controlcannot be carried out by the retinal surgeon. Typi-cal time durations required to form ophthalmologi-cally significant lesions (ranging in size from 100 to500 [lm) are of the order of 100 ms with an argonlaser power setting of 100 mW and a laser spot size of100 Am. Human response time is not fast enough tofollow the lesion formation process. Thus there is aneed for computer control.

Lesion size must be monitored in real time by someindirect method since lesion size cannot be measureddirectly in vivo. Correlating reflectance from the

1200 APPLIED OPTICS / Vol. 32, No. 7 / 1 March 1993

- Z

Fig. 1. Cross section of the fundus. RE, neural retina; PE,pigment epithelium; CH, choroid, th, layer thickness; d, distancebetween the layers; SC, sclera; r, radial direction; -z, atial direc-tion. (Figure reprinted from Ref. 6; 1984 IEEE used withpermission.)

white lesion to its size is noninvasive and has shownthe most promise as a monitoring method.6 7

The idea of using reflectance to improve in someway the accuracy of the photocoagulation process isquite old. Geeraets et al. 8 have shown that reflec-tance is an indicator of the absorption of the fundus,and this information may provide a means of predeter-mining the dosage for retinal photocoagulation.9Also reflectance provides information on the dynamicdevelopment of the laser-induced lesions. The use ofreflectance to monitor lesion development was inves-tigated first by Birngruber et al.10 The work wasextended when Weinberg et al.6 measured a portion oftotal reflection from a lesion by using a single photo-multiplier tube. They attempted to correlate thisreflectance to the area of damage; however, totalreflectance did not correlate well enough with themicroscopically measured area to allow for accuratecomputer control.

Markow 1 and Yang7 tried to relate the lesiondiameter to reflectance information obtained in non-retinal tissue with a two-dimensional array detector.They obtained a high correlation between the diame-ter of the lesion in the reflectance image and theactual lesion diameter. Markow formed lesions inbeef liver with a fixed laser power and spot size byusing feedback from the two-dimensional image toproduce a fixed lesion size by controlling exposuretime. The average lesion diameters were within 8%of desired widths with a standard deviation of 13%.Yang et al.12 13 demonstrated that two-dimensionalreflectance feedback control was feasible in a simu-lated eye phantom consisting of naturally clear albu-min and an absorbing layer. During a single experi-ment with constant illumination and a fixed laserpower setting, lesion diameters of the order of 2 mmwith a 7% standard deviation were formed to within

6% of the desired size. Lesion depth control was notsuccessful. The depths of lesions produced by agiven control criterion varied by 15% and deviated byas much as 38% from expected depths. Thus, whilethey were able to produce fairly uniform lesions, theyhad no way of relating a specified depth to the actualphysical lesion depth. Also, no allowances were madefor variations in illumination that affect directly themeasured gray levels of the CCD array. Reflectanceis extremely dependent on illumination. To producelesions of a predefined size, it is necessary to specify agray level that must be related to illumination.

While a lesion is developing, a two-dimensionalreflectance image of it has two distinct features: (1)the lesion grows monotonically larger in size (corre-sponding to a diameter increase), and (2) the lesiongrows whiter (i.e., it reflects more light). The latterchange is represented by a numerical increase in thegray level of pixels that image the lesion, where blackis represented by a gray level of zero and white by agray level of 255. The diameter can be controlled bymonitoring the reflectance diameter. The diameterof the lesion refers to its growth in the plane perpen-dicular to the laser beam, which is the plane the CCDarray images. It is a fair assumption then that thereis a direct relationship between the diameter deter-mined from the two-dimensional reflectance imageand the actual lesion diameter. The accuracy of themeasurement from the image is dependent on theselection of a threshold gray level to separate thelesion reflectance from the background reflectance.A reflectance image parameter related to lesion depthis not as easy to identify. Information about a thirddimension is sought from a two-dimensional image.Thus an indirect indicator is required.

How can the gray level be related to the depth? Asa lesion thickens a thicker scattering medium isformed, and thus more light is reflected and the lesionappears whiter. However, as the lesion grows dur-ing laser irradiation, there is a certain depth beyondwhich the reflectance does not change (Fig. 2).Thus, as the thickness of a scattering slab increases,reflection approaches a maximum value. This valuethat corresponds to the reflectance for a semi-infinitescattering medium is termed R,, and it is determinedby the absorption and scattering properties of themedium and the optical boundary conditions. 14

When reflection reaches a value in the vicinity of R.,large increases in the thickness of the scatteringmedium produce small increases in reflection. Thuscontrolling the depth based on the magnitude ofreflectance is possible as long as reflectance is muchless than R., and the nature of a relationship betweenthe gray level and depth has been identified. Withthe lesion depth defined as the thickness of the lesionat the center, the central reflectance is an indirectindicator of lesion depth.

We have implemented lesion control by using reflec-tance from a two-dimensional array detector, choos-ing to monitor the reflectance diameter as an indica-tor of lesion diameter and the magnitude of central

1 March 1993 / Vol. 32, No. 7 / APPLIED OPTICS 1201

250

I-'CU

e5.

200 -

150 -a power= 179 mW

* power=110 mW100

~500

0.000 0.528 1.056 1.584 2.112Time (s)

Fig. 2. Central intensity versus time for two lesions formed with different powers (179 and 110 mW) and an exposure time of 2 s. Thecentral intensity for the lesion formed at a power of 179 mW begins to level off after 1 s. The curve seems to approach an R.o gray level of200. The central intensity curve for the lesion formed at 110 mW has a smaller slope and does not begin to level off; i.e., it does not reachthe R. gray level during the exposure time.

reflectance as an indicator of lesion depth. Our workincludes the creation of lesions of a predefined size,either radially or axially, with a calibrated illumina-tion source. This allows for the reproducibility ofresults despite variations not only in the exposureand absorption conditions but also in illuminationintensity. Previous research has not addressed theproblem of illumination variability. This type offeedback control could find application in a computer-

controlled laser delivery system for retinal photocoag-ulation.

Methods

The experimental system for the delivery of argonlight to a target and the measurement of reflectanceimages is shown in Fig. 3. A 5-W-rated CoherentArgon laser, functioning on all argon lines (the pri-mary lines are 488 and 514.5 nm), is fed to a fiber and

I Argon laser

LI

U12x512 Q - -camera

L4

A

-1 - -1- - N INL2

50-50. beam splitter

I BS

L5 I

slotted spinningwheel

(3600 rpm)

Fig. 3. Experimental set up.


I

_ L3

t I

I

target

serves as the coagulating light source. The laserbeam is collimated with two lenses (Li and L2) anddirected to the target with two mirrors (Ml and M2),a focusing lens (L3), and a beam splitter (BS).

A microscope light provides a diffuse, uniformwhite-light illumination source and is simply fed tothe target from the side. The beam splitter mergesthe optical axis of the camera and the axis of the laserbeam. The lenses in the camera leg (L4 and L5)magnify the target image and focus it onto the CCDarray detector of the camera. A five-slotted spinningwheel alternately shutters the laser and the camera toprevent argon laser reflectance from reaching thecamera. The irradiation and CCD interrogation dutycycles are each 33%.

A photodiode and a counter measure the frequencyof the irradiation pulses and hence of the wheel slots.The speed of the spinning wheel is adjusted to 3600rpm (which is a harmonic of the frame acquisitionrate) to ensure that the images recorded by thecamera are of a constant intensity. The power of theportion of the laser light that is directed away fromthe target by the beam splitter is monitored. Thisyields the average power of the chopped laser beamfor every exposure, since the ratio of the powerreaching the target to the power of the split beam isknown.

A shutter in the path of the laser beam is controlledby the computer and can be opened and closed within7 ms with a feedback signal. The signal reaching theshutter is monitored on a storage oscilloscope andyields the exposure time.

Experimental Phantom

A simple phantom is shown in Fig. 4. A Pyrex Petridish is coated with a thin, absorbing layer of high-temperature black paint. The thickness of this layervaries from one location to another but is 40 tim.A 10-mm-thick layer of naturally clear egg white(albumin) overlays the paint.

The absorbing layer simulates the pigment epithe-lium of the eye. This layer does not absorb 100% ofthe light, and its varying thickness provides thevariability in its total absorption. Argon laser lightis absorbed by the layer and heat is generated. Heatis conducted from the absorbing layer to the albumin,and the albumin whitens as it coagulates. The largeincrease in scattering as albumin coagulates is similarto the effect of coagulation on the scattering of retinaltissue. Both whiten as the tissue coagulates, and weassume that this whitening is related to the volumeAv that denatures, whether it be in the model me-

fresh eggwite :i-.4 4E Ed. d...-i..... g .C.-:' .. ..... ..... : 4 .. : SE-EE. ,i. . .. ri..

X i .. _ ~~~~~~~~~~~401Lrn. high temperature black paint 4

Fig. 4. Cross section of the model experimental medium.

dium or in retinal tissue. Thus general principles onreflectance control hold for both media.

Irradiation Parameters

Irradiation parameters were chosen as a compromisebetween the values that are typical of retinal photoco-agulation and the capabilities of control with a stan-dard frame rate camera. Thus the laser spot size,800 jim (1/e2 diameter), was slightly larger than theretinal laser diameters. Power ranging from 163 to204 mW was typical of that used for photocoagulation.Although the coagulation of both albumin and retinais a temperature-time-dependent rate process, thetwo have different rate coefficients, whose reportedvalues for the two are shown in Table 1. The lowerrate of accumulation of damage at a fixed tempera-ture for albumin relative to the retina increased thetime required to achieve a specified lesion size relativeto corresponding irradiation time for retinal photoco-agulation.

Image ProcessingAn Imaging Technology Inc. 151 image processorcontrolled by a Sun 3/260 computer (see Fig. 3) grabsthe reflectance images of the lesion recorded by theCCD camera (512 x 512 pixels, 30 frames/s) anddigitizes the pixels of each frame to 256 gray levels.The location of a lesion in a frame (x0, yo) will dependon the alignment of the system and is identifiedbefore a feedback experiment. The image processorcontains a frame buffer that can store a limitednumber of frames (four 512 x 512 frames) and per-form simple operations in real time. For example,the gray-level values of pixels can be determined,pixel values can be summed, and pixel values can becompared. The amount of processing that can bedone in real time is limited by the time intervalbetween the frame acquisitions. In general spatialfiltering is not possible at standard frame ratesbecause multiple passes of frames through a bufferare required. Given these limitations real-time pro-cessing of the frames for feedback control has beenkept simple.

In the diameter feedback control algorithm fourlocations are monitored for the threshold gray levelthat defines the boundary of a lesion. The thresholdused must be defined a priori. In the current imple-mentation the threshold is set to a gray level of 5%more than the average gray level of the pixels in aframe before a lesion begins to form, i.e., 1.05x theaverage background reflectance of the medium. If 4)is the selected reflectance diameter in pixels, the fourlocations monitored are (x0, yo + 4)/2), (xo, yo - 4)/2),(x0 + 4)/2, yo), and (x0 - 4)/2, yo). That is, the radiusis measured along only four directions, emanating

Table 1. Rate Coefficients Reported for Retinal Tissue17 and Albumins

Medium A (s'1) E (cal/M)

Retina 1.3 x 1099 1.5 X 105Albumin 3.8 x 1057 9.2 x 104


north, south, east, and west from the center. Thesum of the gray levels at those four locations iscompared with 4 times the threshold.

In the depth feedback control algorithm, the cen-tral reflectance is defined as the average gray level ofthe central five pixels. The central reflectance ismonitored for a threshold gray level that must also bedefined a priori. The sum of the gray levels of thepixel at the center of the lesion and its four immediateneighbors, i.e., (xo, yo), (xo + 1, yo), (x0 - 1, YO),(x0, yo + 1), and (xo, yo - 1), is compared to 5 timesthe selected threshold.

The maximum delay time for feedback control is 40ms. Consider that in our forming lesion the controlcondition is met immediately after a frame is grabbedand processed, e.g., the central reflectance exceedsthe threshold just after a frame is clocked out.Given our frame rate of 30 frames/s, 33 ms willelapse before the next frame is examined by theprocessor. At this point the processor will deter-mine that the laser needs to be shut off since thecontrol condition is met. The shutter response timeis bounded from above by 7 ms. Therefore a maxi-mum of 40 ms (33 + 7) could elapse between thecontrol condition being met and action being taken.

Measurement of Physical Lesion Size

Immediately after an experiment, the lesion diameterwas measured with a stereo microscope. The micro-scopic diameter has been shown to correlate to thediameter determined histologically.6 The lesions aremagnified and measured with a calibrated grid. Theresolution of the grid is ± 20 m. Lesion boundariesare visually determined in making this measurement.This task is relatively simple when the whiteness of alesion is in stark contrast to its background, as is thecase with lesions in our medium.

The lesion depth is more difficult to measure. Thelaser-induced lesions are too small to be sliced with aknife, turned on their side, and measured with amicroscope without the use of special chemical fixa-tives and a microtome. Histology is unreliable formeasuring depth since the fixative shrinks the me-dium. In addition, only a slice through the exactcenter of the lesion yields the true value of themaximal depth of the lesion We can eliminate manyof these problems by measuring lesion depth with aZeiss LSCM10 (confocal) microscope. The micro-scope allows accurate optical slicing of a specimen.Using a He-Ne light source, it images different zplanes showing only those portions of the sample thatare in focus. Thus it is possible to focus first on thebottom of the lesion and then step through the lesionto the top in adjustable Az steps. The total distancemoved to pass through the lesion is recorded, and thisvalue corresponds to the depth of the lesion. Theresolution of the measurement is 5 im.

Feedback Control Experiments

With the above experimental tools and simulatedmedium we have attempted, based on parameters of

the reflectance image, to control lesion size as thelesion is forming. Figure 5 summarizes our experi-mental approach.

Preliminary experiments were uncalibrated, i.e.,simply geared toward verifying that our chosen param-eters can control the lesion size at a fixed level ofillumination. Selecting arbitrary values for our con-trol parameters for diameter and depth, several le-sions were formed at differing power levels with theexposure time controlled by the feedback system.That is, we defined a random reflectance diameter fordiameter control and a random central reflectancethreshold for depth control.

CalibrationCalibrating the lesion diameter involved determiningthe relationship between the lesion diameter and thereflectance diameter. This was influenced by thesystem's magnification of the lesion, by the definitionof the reflectance diameter (i.e., the threshold graylevel at the boundary), and by the accuracy in measur-ing the physical lesion size. In addition, the thresh-old gray level depends on illumination. We deter-mined the relationship empirically by forming severallesions of various sizes as a function of power andexposure time. The physical lesion diameter wasmeasured and compared with the diameter measuredfrom the CCD reflectance images. The images weretaken in varying illumination conditions and pro-cessed by using a threshold gray level that is 1.05times the background gray level to define the lesionboundary. This definition of the threshold as arelative reflectance should render it independent ofillumination.

reflectancelesion diameter

I< 'Icentral reflectance

0I-

illumination

E e

-c egg whitec (slightly scattering medium)

W high temperature black paint

I I

tI __ (ref]

reflected light

/

/

lesionlecting hemisphere)

Fig. 5. Experimental approach.


I l

I

Calibrating the lesion depth involved determiningthe relationship between central reflectance and le-sion thickness. Since the central reflectance is anaverage gray level, its value depends on the level ofillumination. The calibration of lesion depth wasalso empirical. Lesions of varying sizes were formed,and their physical depths were measured. Thesedepths were compared with the central reflectancesmeasured from the lesion reflectance images taken invarying illumination conditions. The illuminationlevels were identified by the gray-level reflectance offof a uniform gray reference surface. From thesedata a relationship between the central reflectance(gray level), the illumination intensity, and the lesiondepth was generated.

Calibrated Feedback ControlUsing calibration information, we attempted to formconsistent lesions of a predefined size in varyingillumination, absorption, and laser-power-deliveryconditions. The desired diameters were 450 and 700jim, and the desired depths were 60 and 85 im.Depth dimensions were chosen so that they resem-bled those desired in retinal photocoagulation. Wemade the diameter sizes larger to lengthen the re-sponse time with the laser spot diameter given so thatstandard video rates provided reasonable sampling ofthe transient reflectance response. This created acontrol situation within the limits of our system.

In the calibrated diameter control experiment, theaverage background reflectance value gl was mea-

180

160

::.0)

iUI.-

._

-1

U_0)

..

a0_

140

120

100

sured before a lesion was formed, and the gray-levelthreshold of the lesion boundary was calculated as X =1.05 x gl. When the average gray level of themonitored pixels reached r, irradiation was stoppedby closing the shutter. For each of the two diame-ters attempted, 20 lesions were created: five each attwo different powers for each of two illuminationlevels. The physical diameters of the resulting le-sions were measured.

For the calibrated control of depth the centralreflectance threshold X for each of the desired depthswas determined from the relationship between thegray level and the depth (Fig. 6). When the averagegray level of the monitored pixels reached T, irradia-tion was stopped by closing the shutter. Once againfor each desired depth 20 lesions were formed: fiveeach at two different powers for each of two differentillumination-intensity levels. The physical depths ofthe resulting lesions were measured.

Theoretical Analysis

A theoretical analysis of the lesion-formation processwas carried out to examine the effect of the maximumdelay time for feedback control on the resulting lesionsize. The denaturation that occurs as the mediumcools after the laser has been turned off was alsodetermined. A model for the prediction of retinalinjury because of the laser irradiation developed forthe U. S. Air Force School of Aerospace Medicine byTakata et al. 15,16 was used. The model implementeda finite difference solution to the heat conduction

0 20 40 60 80

E ill int= 132

* ill int = 149

* ill int= 159

1 00

depth (im)

Fig. 6. Calibration curves for central intensity versus the maximum depth for the average illumination gray levels of 132, 149, and 159 interms of the reflectance off a reference surface: ill int, illumination intensity.


incident energy

40 lm

non-absorbing layer withthermal properties of water

absorbing layers

non-absorbing layer withthermal properties of Pyrex

Fig. 7. Geometry of the Takata model as used in our simulation.

equation to compute the temperature rise in multiplelayers in response to laser irradiation. The Hen-riques damage integral,' 7 which is based on theArrhenius reaction rate equation, was calculatedfrom the modeled temperature-time history. Themodel predicted the extent of thermal damage byidentifying locations in the r - z grid where the resultof the damage integral fl was > 1.

The rate reaction parameters for the albumin usedwere based on those calculated by Yang et al.18Their values were used as a starting point, and thevalues were then fine tuned to yield simulated lesionsizes that are close to the experimentally observedones for a particular power-exposure time combina-tion. The final values chosen were E = 9.2 x 104cal/M (which is exactly the same as Yang et al. 's) andA = 3.8 x 1058 s1 (which is 1 order of magnitudelarger). The laser incident powers used in the modelwere the same as the average power used in thefeedback experiments. The exposure times used werebased on experimental measurements. The geome-try and parameter values used in runs to simulate thethermal damage in our experimental phantom areshown in Fig. 7 and Tables 2 and 3, respectively.

Results and Discussion

The results of preliminary experiments reaffirmedYang et al. 's results on lesion size control based on thechosen parameters.' 3 Within each group of lesionsformed with one power setting, the feedback-con-trolled lesion size varied by < 5%, while the exposuretime required to create them varied by at least 10%.This variability in the exposure time indicated differ-ing absorption conditions, as was expected because ofthe uneven thickness of the paint layer. The size oflesions formed from one power setting to another

Table 3. Power and Exposure Time Combinations Used in TakataModel Simulation Runs

Power Exposure TimeRun (mW) (ins)

1 221 5252 221 5653 202 4554 202 4755 202 495

varied by only 5%, despite a 100% increase in expo-sure time required for the lower power setting.

Calibration

The calibration experiments yielded the quantitativerelationship between the reflectance image parame-ters and the physical lesion size that is necessary toform lesions of a predefined size. The reflectancediameter was on the average 0.48 times the physicaldiameter. This was a constant that held true consis-tently for all the lesions and, as expected, in allillumination conditions. For a given illuminationthe central reflectance increased with lesion depthfairly linearly before leveling off to R.. Also theslope of the linear portion of the curve was found to beproportional to illumination. The calibration curvesobtained are shown in Fig. 6.

Calibrated Feedback Control

Experimental results for diameter control with varia-tions in power and illumination intensity (whichcorresponds to differing background gray levels) areshown in Table 4. Note that for each of the twodiameter sets the actual lesion sizes obtained in eachset are close to the desired values. Also, the actuallesion diameters have low standard deviations despitevariations in power and illumination.

Experimental results for depth control in varyingillumination conditions are shown in Table 5. Forthe larger desired depth (85 m), not only are theachieved depths in all sets close to the desired valuebut also, within each set, the standard deviations arelow. For the smaller desired depth (60 m), how-ever, while the lesions of each set are close to eachother in size and have low standard deviations, theachieved sizes are 10-15% larger than the desireddepth.

The thermal gradient in tissue induced by thelocalized absorption of laser light is dynamic. Ini-

Table 2. Parameter Values Used for Takata Model Simulations

Parameters

Layer in Original Corresponding Layer Thickness P-a K Cp X p

Takata Model in Simulations (cm) (cm -) (W/cm/0 C) (J/cm 3/C)

Retina Clear egg white 1 6.28 x 10-3 4.187Pigment epithelium Black paint 2 x 10-3 500 6.28 x 10-3 4.187Choroid Black paint 2 x 10-3 500 6.28 x 10-3 4.187

Sclera Pyrex Petri dish 0.9 1.4 x 10-2 1.96


1-1

gotI llm�2 cm

Table 4. Results for Calibrated Feedback Control of the LesionDiametera

AverageDesired Average Lesion

Diameter Control Power Diameter(PM)b Conditions (mW) (P-m) or

700 gl = 125 185 665 20= 131= 342 163 660 12.2

Over the set 663 16.8

gl = 118 192 685 12.2= 124= 342 178 675 27.4

Over the set 680 21.8

450 gl = 125 163 470 10T = 131

- = 220 178 455 18.7Over the set 463 16.8

gl = 129 178 445 18.7T = 135

= 220 170 455 18.7Over the set 450 19.4

aAverage lesions diameters were determined over five lesionswith a standard deviation of or. gl is the background gray level, Tthe threshold gray level of the lesion boundary, and + the corre-sponding desired reflectance diameter in pixels. The laser spotl/e 2 diameter was 800 im.

tially the gradient is large, which leads to a rapidincrease in temperature. As the heat is conductedthrough the tissue, the gradient decreases and henceheat conduction slows.6"9 This process has been

10

8

6

4

04

2

0

Table 5. Results for the Calibrated Feedback Control of Lesion Deptha

AverageDesired Average LesionDepth Control Power Depth(P-) Conditions (mW) (P-m) a

60 Ill int = 132 193 69 1.2T = 137 181 66 2.0Over the set 67.5 2.2

Illint = 149 181 69 2.0T = 152 202 73.5 1.2Over the set 71.3 2.8

85 Ill int = 132 163 85 1.6T = 149 190 87 1.9Over the set 86 2.0

Ill int = 149 204 89 2.0T = 170 181 82.5 2.7Over the set 85.8 4.0

aAverage lesion depths were determined over five lesions with astandard deviation of a: Ill int is the illumination intensitymeasured as the reflectance off a reference surface. is thecorresponding gray-level limit of the central reflectance. Thelaser spot 1l/e2 diameter was 800 PLm.

theoretically demonstrated as corresponding to aninitial rapid growth of lesion size followed by a slowergrowth. 5

This could explain our inconsistent results whenwe used a standard CCD array (30 frames/s), i.e., ourability to control lesion depths of 85 jim but not lesiondepths of 60 pm. The effect of maximum delay timefor control is more critical for smaller lesions than for

mW ms

202 455

\ - -202 495\ \ ~~~221 525

\ '\ --- 221 565

- -- 202 475

\ \

' \

- \

-K L\

30 40 50 60 70 80 90 1 00depth (m)

Fig. 8. Results of the simulation runs of the Takata model.represents fl = 1, the damage boundary.

Each line corresponds to a power-exposure time pair. The horizontal line


larger ones. This is seen in the results of the Takatamodel simulation runs.

The results are shown in Fig. 8. A 221-mW,525-ms exposure produces an 87-jim-deep lesion.When the irradiation time is extended by the delaytime, an 89-jim lesion is formed, which is a differenceof only 2 jim. However, a 202-mW, 455-ms exposureproduces a 59-jim-deep lesion, and a 202-mW, 495-msexposure produces a 68-jim lesion, which is a differ-ence of 9 jim. Thus the control time is more criticalfor the smaller lesion (i.e., earlier in the lesionformation process when thermal gradients are larger)than for the larger lesion whose growth has presum-ably stabilized more with the decreased rate of heatconduction. If the control time is reduced to 20 ms,the control of smaller lesions is more successful. Forexample, a 202-mW, 475-ms exposure produces a64-jim lesion.

The computations of damage for albumin illustratethat only 1 jim of the 87-jim lesion growth and 2 jimof the 60-jim lesion growth occur after the laser isshut off. Thus, at least in the model medium, thesustained elevated temperature field after the laser isshut off adds an insignificant amount to the lesionsize.

Sources of Error

We have several sources of error; discussed below arethe more important ones:

(1) There is a limitation imposed by our frameacquisition rate. At a standard frame rate of 30frames/s, 33 ms could elapse before the controlsystem recognizes that a threshold has been reached.Thus the maximum delay time is fairly significant.

(2) There are errors introduced by the minimumresolution of the physical lesion size measurementtechniques, which are ± 5 m for the lesion depth and+20 jim for the lesion diameter. Also, defining thephysical lesion boundaries through the microscopes issomewhat subjective. These errors were minimizedwhen the same person made repeated measurements.

(3) The illumination source is not completelystable. Any change in the illumination intensityaffects the gray-level values in the reflectance image.This in turn affects what the control system calibra-tion, since it is making decisions based solely on thegray-level values.

Conclusions

The ability to produce consistent laser-induced le-sions of a predefined size, either radially or axially,independent of the illumination intensity, with acalibrated feedback system has been demonstrated.Calibration is necessary for extending this work toapplications such as retinal photocoagulation wherethe illumination intensity may vary throughout thetarget area. For a system realistically to producelesions of a predefined size by using reflectancecontrol, it must compensate for illumination differ-ences.

The modeling of the temperature and the extent ofdamage provide a realistic picture of photocoagulation.Computed results provide critical information for theselection of the control sampling rate. A maximumdelay time of 20 ms (e.g., a frame rate of 60 frames/s)would be necessary to control 60 jim or larger lesiondepths within ± 5 jim in our albumin model. Model-ing should be an integral part of a reflectance controlsystem for medical applications such as retinal photo-coagulation.

This work was supported in part by the TexasCoordinating Board and in part by the U.S. Office ofNaval Research under grant N00014-91-J-1564.

Ashley J. Welch is the Marion E. Forsman Centen-nial Professor of Electrical and Computer Engineer-ing and Biomedical Engineering.

References1. V. P. Gabel, "Lasers in ophthalmology," in Lasers in Biology

and Medicine, F. Hillenkamp, ed. (Plenum, New York, 1979),pp.383-400.

2. M. L. Wolbarsht, "Ophthalmic uses of lasers," in Lasers inBiology and Medicine, F. Hillenkamp, ed. (Plenum, New York,1979), pp. 442-446.

3. H. C. Zweng, "Lasers in ophthalmology," in Laser Applica-tions in Medicine and Biology I, M. L. Wolbarsht, ed (Plenum,New York, 1971), pp. 239-254.

4. V. P. Gabel, R. Birngruber, and F. Hillenkamp, Die Lichtabsorp-tion am Augenhintergrund (Gesellschaft fur Strahlen-undUmweltforschung mbH, Munchen, 1976).

5. M. R. Jerath, "A software package for the analysis of laserinduced retinal lesions," M. S. thesis (University of Texas atAustin, Austin, Tex., 1989).

6. W. Weinberg, R. Birngruber, and B. Lorenz, "The change inlight reflection of the retina during therapeutic laser-photocoagulation," IEEE J. Quantum Electron. QE-20, 1481-1489 (1984).

7. Y. Yang, "Automatic control of the extent of laser inducedcoagulation," Ph.D. dissertation (University of Texas at Aus-tin, Austin, Tex., 1990).

8. W. J. Geeraets, R. C. Williams, M. Ghosh, W. T. Ham, D.Guerry, F. Schmidt, and R. Ruffin, "Light reflectance from theocular fundus," Arch. Ophthalmol 69, 612-617 (1963).

9. 0. Pomerantzeff, G. J. Wang, M. Pankratov, and J. Schneider,"A method to predetermine the correct photocoagulationdosage," Arch. Ophthalmol. 101, 949-953 (1983).

10. R. Birngruber, V P. Gabel, and F. Hillenkamp, "Fundusreflectometry: a step towards optimization and retina photocoagulation," Mod. Probl. Ophthalmol 18, 383-390 (1979).

11. M. S. Markow, "The experimental investigation of a roboticlaser system used for ophthalmic surgery," Ph.D. dissertation(University of Texas at Austin, Austin, Tex., 1987).

12. Y. Yang, M. S. Markow, H. G. Rylander III, W. S. Weinberg,and A. J. Welch, "Reflectance as an indirect measurement ofthe extent of laser-induced coagulation," IEEE Trans. Biomed.Eng. 37, 466-473 (1990).

13. Y. Yang, M. S. Markow, H. G. Rylander III, and A. J. Welch,"Automatic control of lesion size in a simulated model of theeye," IEEE J. Quantum Electron. 26, 2232-2239 (1990).

14. P. Kubelka, "New contributions to the optics of intensely lightscattering materials, Part I," J. Opt. Soc. Am. 38, 448-457(1948).

15. A. N. Takata, L. Goldfinch, J. K. Hinds, P. Kuan, N. Thomopu-lis, and A. Weigand, "Thermal model of laser induced eyedamage," IITRI Tech. Rep. 74-6324 (ITT Research Institute,Chicago, Ill., 1974).


16. A. J. Welch and G. D. Polhamus, "Measurement of thermalinjury in the retina of the rhesus monkey," IEEE Trans.Biomed. Eng. 31, 633-644 (1984).

17. F. C. Henriques, "Studies of thermal injury," Arch. Pathol.43, 489-502 (1947).

18. Y. Yang, A. J. Welch, and H. G. Rylander III, "Rate processparameters of albumen," Lasers Surg. Med. 11, 188-190(1991).

19. H. S. Carslaw and J. C. Yaeger, Conduction of Heat in Solids(Clarendon, Oxford, 1959).


calibrated real-time control of lesion size based on reflectance images

Documents