light dosimetry: effects of dehydration and thermal damage on the optical properties of the human...

Light dosimetry: effects of dehydration andthermal damage on the optical properties of thehuman aorta

Inci F. Qilesiz and Ashley J. Welch

The influences of dehydration and thermal damage on in vitro optical properties of human aorta werestudied. The absorption coefficient increased by 20-50%, especially in the visible range when at least40% of total tissue weight was lost as a result of dehydration. The reduced scattering coefficientincreased by 10-45% in the visible and 30% to over 150% in the near IR after the tissue samples wereheated in a constant temperature water bath at 100 0C for 300 ± 10 s. This study implies thatdehydration and protein coagulation during photothermal treatment of tissue are important factorsaltering optical properties of tissue.

Key words: Dosimetry, hydration, optical properties, protein coagulation, spectrophotometer.

Introduction and BackgroundThe choice of lasers and exposure parameters ingeneral has been based on clinical results rather thanon an understanding of (1) the mechanisms involvedin laser treatment, (2) the optical and thermal proper-ties of tissue, and (3) a rational theoretical approachto maximize therapeutic benefits. Previous studieson laser-tissue interaction have shown that the ap-pearance of tissues changes considerably duringphotothermal processes.1-3 A number of authorshave observed that the clinical levels of laser irradia-tions vaporize water in the near-surface tissue caus-ing sharp increases in temperature. 6 Increasingthe tissue temperature above 50'C may also causedenaturation of molecules, dehydration, coagulation,and eventually ablation.7

Previous investigations showed that tissue thick-ness and volume shrink as a result of dehydrationduring thermal insult to the human aorta. Signifi-cant changes in transmittance and reflectance startedbetween 450 and 70'C where transmission decreasedand reflectance increased. At higher temperatures(80-100'C) dehydration was the dominating factorincreasing transmittance and decreasing reflectance.Beyond 110'C reflectance increased, whereas trans-

I. F. ilesiz and A. J. Welch are with the Biomedical EngineeringProgram, University of Texas at Austin, Austin, Texas 78712-1084.

Received 6 December 1991.0003-6935/93/040477-11$05.00/0.i 1993 Optical Society of America.

mittance decreased. 8 More recently, studies on ther-mally induced optical property changes in myocar-dium emphasized irreversible changes between 60°and 750C, the temperature range in which extracellu-lar protein and collagen denaturation are known tooccur. Derbyshire et al. 9 and Splinter et al. 10 specu-lated that protein macromolecules that are responsi-ble for light scattering had little effect on absorption,because they observed a twofold to threefold increasein the reduced scattering coefficient as a result ofthermal insult, whereas the absorption coefficientremained relatively constant.

Hence the deposition of thermal energy in tissue isnot only a function of laser irradiation parameters,but it also depends on the physical properties of tissueincluding the optical properties. Thus the photother-mal response of tissue depends on temperature, watercontent, and tissue condition. An understanding ofthe influence of these parameters on optical proper-ties will enhance our ability to predict the photother-mal response of tissue to laser irradiation.

Most of the recent advances describing laser-tissueinteractions are based on transport theory. Tissueis considered to be a medium of dense distributions ofrandom scatterers." Although analytic models havebeen developed to solve Maxwell equations for scatter-ing particles, transport theory is preferred in studiesof tissue optics involving optically thick biologicalmedia. Yet implicit assumptions must be made iftransport theory is used to describe laser-tissueinteraction: (1) The biological medium under inves-tigation is homogeneous. (2) Each particle in the

1 February 1993 / Vol. 32, No. 4 / APPLIED OPTICS 477

biological medium is sufficiently isolated that itsscattering pattern is independent of all other particles.(3) Scattering by all particles may be described by asingle phase function. (4) Polarization effects areneglected. (5) Radiation is transverse with only smallchanges in electrical permittivity at boundaries. Adescription of transport theory applied to tissue canbe found in Ishimaru's book."'

The optical properties of tissue are obtained byconverting measurements of observable quantitiesinto parameters that characterize light propagationin tissue. The fundamental optical properties oftissue related to the radiative transfer theory'2 arethe absorption coefficient p.,a (1/cm), the scatteringcoefficient p's (1/cm), and the average cosine of thescattering angle associated with the single-scatteringphase functiong. The products of each ofthe formertwo parameters with path length As give the probabil-ities that a photon will be absorbed or conservativelyscattered in As as As -> 0. The phase functiondescribes the angular distribution for a single-scattering event, i.e., the probability per unit solidangle that a photon will be scattered from an angle ('into an angle 0". When the average cosine of thephase function g is equal to 1, scattering is purelyforward; when g = 0 the scattering is isotropic.

For an in vitro tissue sample the computation of pa,[.,, and g requires the measurement of thickness of auniform sample and three optical measurementssuch as total reflection, total or diffuse transmission,and collimated transmission. Assuming a preset gvalue, the similarity relations permit the determina-tion of an absorption coefficient pa and a reducedscattering coefficient [L,' = LS(l - g) from two opticalmeasurements: typically diffuse reflection Rd andtotal transmission Tt, which are measured by usingan integrating sphere system.13

Unfortunately there is no direct relation betweenpermitting the calculation of pa and p,' from reflec-tion and transmission measurements. If Rd and Ttare measured, it is necessary to form an iterativecomputation that varies p.a and p.' until the mea-sured values of Rd and Tt are computed within apreselected error criteria. A one-dimensional diffu-sion approximation model with a 8-Eddington phasefunction' 4 has been used for this purpose.

Materials and MethodsA Varian 2300 UV-visible-near-IR spectrophotome-ter was used to measure diffuse reflection and totaltransmission. The experiments within the contextof this study covered a spectrum ranging from 300 to1800 nm.

Data were taken at 5-nm intervals, and a two-pointcalibration was performed by using experimental andstandard data from two stimulated Raman scatteringseries reflectance standards (2% and 99%) manufac-tured by LabSphere Inc. The diffusion approxima-tion with a -Eddington phase function, which as-signs forward-scattered light into a function, wasused in an iterative program developed by S. Prahl to

compute diffuse reflection and total transmission foran assumed pair of values for absorption and reducedscattering coefficients.' 4 This program consideredmultiple reflections that occurred at air-slide-tissue-slide-air interfaces. New values of [La and p,' wereautomatically computed until reflection and transmis-sion matched measured values. An outline of theexperimental protocol is shown in Fig. 1.

Tissue PreparationAorta segments were harvested from human cadav-ers within 24 h postmortem and delivered in anairtight container. To remove residual blood, wesoaked the segments in 0.9% isotonic saline solutionfor at least 1 h. Although the cleaning did not alterthe in vitro hydration levels, placing tissue in fluidmay have slightly increased the in vitro hydrationlevel relative to the in vivo level. Segments werechecked for any visible plaque and fatty streaks.Those with clean tunica intima were prepared forexperiments. The outer part of the adventitia withmost blood vessels was removed.

Squares of 1 cm x 1 cm were cut. Some sampleswere stripped to a thickness of 500-700 m leavingmostly intimal and medial layers. A microtome wasused for most samples, and thicknesses down to 250jum were obtained without freezing the samples.Microtome cuts were more uniform both in thicknessand in surface quality. The thin samples were moist-ened with saline and carefully sandwiched betweentwo slides so that no air bubbles were trapped be-tween the tissue and each slide. Finally this assem-bly was mounted in a custom-made frame for spectro-photometric measurements as shown in Fig. 2.Quartz slides (GM Associates Inc., part 7525-02) wereused to minimize losses in the UV spectrum. Pre-pared but not immediately used samples were wrappedin saline-moistened gauze, sealed, and refrigerated at40C. Except for dehydration studies, samples werenot kept longer than 24 h. The maximum time ofrefrigeration was limited to 72 h.' 5

Record a baseline between the wavelengthsof interest

Record calibration using two reflectionstandards (e.g. 2%, 99%)

Measure sample thickness, diffusereflection Rdand total transmission Tt

4Run post-processor programto calibrate experimental data

Run inverse delta-Eddington Programto obtain optical properties

Fig. 1. Experimental protocol.

478 APPLIED OPTICS / Vol. 32, No. 4 / 1 February 1993

group prepared thin samples were heated in a con-stant-temperature water bath after control measure-ments. The samples in this second group werewrapped in aluminum foil before being placed in aconstant-temperature water bath. The aluminumfoil was employed to prevent diffusion of saline intotissue and diffusion of chromophores into the bath17

and thus to investigate the influence of such diffusionon optical properties. After immersion cooling atroom temperature ( 220C) for 5 min, the opticalmeasurements were repeated.

Since the scattering is known to decrease as apower of wavelength, we assumed that

us, ° -n,

alwninumframe

Fig. 2. Tissue preparation for spectrophotometric measurements.Top: view of the assembled sample from the top; bottom: viewat the AA cross section.

To determine the effects of hydration levels, wescanned and weighed the samples. Then they weresubjected to slow dehydration in a refrigerator for atleast 20 h. To isolate the influence of dehydrationfrom that of thermal damage, we induced dehydra-tion in a refrigerator instead of in a low-temperatureoven. A weight loss of at least 40% was achievedwithin 2-3 days without any visible sign of tissuedecay. The samples were maintained between thesandwiching slides but with the seals loosened toprovide dehydration and to maintain smooth surfaces.The samples were weighed, and the thicknesses weremeasured after each scan. A Mettler PE 360 digitalscale ( 1 mg) and a Hommel Schnelltaster caliper(±25 jim) were used to weigh and measure thethicknesses of the samples, respectively.15

Thermal damage, as characterized primarily byprotein coagulation, was induced by bathing aortasamples in a constant-temperature isotonic salinebath (Nesleb RTE-190 constant-temperature bath).The evaporation of water in the bath was minimizedby covering the bath. Saline levels at the start andfinish of the experiment were approximately thesame. Samples were either exposed directly to heator wrapped in aluminum foil at temperatures of 60°,70°, and 1000C for 300 ± 10 S.15 Since the thermaldiffusion time for a 1.0-mm sample is 1.0 S,16 the300-s heating time ensured that the temperature atthe center of a 1-mm-thick tissue sample approxi-mated the surface temperature. For thermal dam-age tests at 60° and 70'C aorta segments were placedin the constant-temperature water bath. Followingimmersion cooling, to avoid further thermal damage,we prepared thin samples as described before.Different samples from the same specimen were usedfor the control and thermal damage experiments.Two groups of experiments were conducted at 10000.In one group samples were processed in the samemanner as discussed above, whereas in the other

ln pus' = C - n n X,

(la)

(lb)

where C is a constant. Least-squares values of nwere computed for normal, dehydrated, and damagedtissue samples in the 500-1200-nm range, where theslope of the reduced scattering spectra is the mostlinear in the logarithmic scale. A t-test was used tocompare the significance between paired conditions.The hypothesis that samples from two sets of datacould come from the same population, i.e., the ex-pected value of one set is equal to the expected valueof another set, is tested.

ResultsA total of 22 tissue segments was used in the experi-mentation. In the experiments involving the directheating of tissue, these segments were cut in two ormore pieces; one or two samples were used for controlmeasurements, and the other(s) was placed in thetemperature bath. Typical curves for transmission,reflection, and the corresponding optical propertiesare shown in Figs. 3 and 5 for dehydrated and heatedsamples, respectively.

Effects of DehydrationNine sets of experiments were carried out. Typicaltransmission and reflection curves are presented inFigs. 3(a) and 3(b). The corresponding absorptionand reduced scattering spectra are shown in Figs. 3(c)and 3(d). It was observed that the absorption coeffi-cient of the human aorta increased, particularly inthe 400-1300-nm range, when at least 40% of thetotal weight of tissue was lost because of dehydration.The average weight loss of 46.4 ± 7.6% was accompa-nied with an average thickness shrinkage of 19.5 ±4.8%. The loss of water decreased the sample thick-ness. Primarily because of shrinkage the absorptioncoefficient was increased by 20-50% in this range.Even though data points in the water absorptionband were occasionally lost because of low levels ofdetected signals, there was an increase in the transmis-sion of dehydrated samples between 1400 and 1550nm as seen in Fig. 3(a). Yet the absorption coeffi-cient was not affected by water loss. The fraction of


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(b) (d)Fig. 3. Raw transmission (a) and reflection (b) data from the human aorta dehydrated for 48 h, corresponding absorption (c), and reducedscattering spectra (d). In control measurements the thickness was 300 jim, the weight was 162 mg. The 24-h-dehydrated sample was250 pum thick and weighed 98 mg. Following the 48-h dehydration the sample was 225 pum thick and weighed 74 mg. Note that thediscontinuity at 800 nm is an instrument artifact caused by detector change. The dips in the reduced scattering spectra are artifactscaused by the S-Eddington approximation that cannot give accurateabsorption band.14'18

light absorbed at near-IR wavelengths beyond 1550nm decreased by 10-20%.

By assuming that the aorta is composed of 60-80%water, an average weight loss of 46% corresponded to58-77% dehydration. The optical properties of nor-mal and 50-90% dehydrated human aorta at selectedwavelengths are summarized in Table 1. Thechanges in the reduced scattering coefficient of hu-man aorta were not as consistent and as pronouncedas the changes in the absorption coefficient as seen inFig. 4. Yet there was a slight increase of 2-15% inthe visible range.

It became more and more difficult to prepare tissuesamples for measurements for dehydration levelsbeyond 90% (i.e., a weight loss over 60%) sincesurfaces were no longer smooth and air bubbles weretrapped between slides and tissue samples.

properties for regions of very high absorption, such as the water

Effects of Thermal DamageFourteen sets of experiments were carried out.They were divided into three groups according to thetemperature of the isotonic saline bath. Sampleswere heated at 60'C (three sets), 70'C (two sets), or100 0C (nine sets) for 300 + 10 s. Typical curves fortransmission, reflection, the absorption coefficient,and the reduced scattering coefficient for thermaldamage studies at 100 0C are shown in Fig. 4.

In our initial experiments on the influence ofthermal damage at 60° and 70'C on the opticalproperties of the human aorta, we obtained resultswith large standard deviations as seen in Tables 2 and3. Therefore experiments were continued on theeffects of thermal damage at 100 0C.15

When a human aorta was exposed to a temperatureof 100TC for 300 10 s, a slight decrease was


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Table 1. Optical Properties of Normal and Dehydrated Human Aorta at Selected Wavelengths: nnorma = ndehydrated = 9a

A .La Normal pa Dehydrated Change in La A Lu' Normal Lu' Dehydrated Change in p,,'(nm) (cm-1) (cm-') (%) (nm) (cm-') (cm-i) (%)

350 20.5 + 2.8 25.7 + 4.1 +25.9 350 53.0 ± 13.9 60.3 ± 12.9 +15.7490 4.7 + 0.8 6.5 + 1.8 +35.1 490 35.8 ± 4.9 39.1 ± 5.7 +9.3515 3.7 + 0.8 5.2 ± 1.7 +35.3 515 33.3 ± 4.8 36.4 ± 5.3 +9.4630 2.5 + 0.7 3.6 ± 1.4 +42.2 630 25.4 ± 3.9 27.6 ± 3.9 +9.1

1065 3.2 + 0.7 4.4 ± 1.1 +35.8 1065 14.2 ± 2.2 14.8 ± 2.0 +5.41320 5.4 ± 0.6 6.5 ± 1.0 +19.1 1320 10.7 ± 1.7 10.9 ± 1.6 +3.01400 26.7 ± 3.5b 26.5 ± 2.7b +11.7c 1400 6.2 ± 3.2c 6.2 ± 2.9b -35.3c1450 43.1 ± 4.1d 42.0 ± 0.7d +5.9e 1450 4.1 ± 1.4d 3.1 ± 1.3d -20.4e1500 28.6 ± 3.2f 29.5 ± 3.2f +12.4g 1500 4.1 ± 3.3f 5.7 ± 2.5f -14.2g1550 16.7 ± 1.9 17.5 ± 1.7 +5.5 1550 6.8 ± 1.8 6.6 ± 1.8 +5.51650 9.4 ± 0.8 10.2 ± 1.0 +8.8 1650 7.8 ± 1.3 7.7 ± 1.2 +0.91750 12.2 ± 1.1 13.3 ± 1.3 +9.8 1750 6.7 ± 1.3 6.4 ± 1.4 -0.4

aThe values are given as a mean ± standard deviation. The corresponding tissue thicknesses for normal and dehydrated samples were575 ± 138 and 464 ± 116 m, respectively. The percentage weight loss and thickness shrinkage caused by dehydration were 46.4 ± 7.6%and 19.5 ± 4.8%, respectively. % change = {[p. (dehydrated)/,u (control)] - 1}100.

bFrom only five control and five dehydrated samples, which are not necessarily the same samples.CFrom scans of only two samples, which did not miss data points in the water absorption band.d From only two control and two dehydrated samples, which are not necessarily the same samples.eFrom scans of only one sample, which did not miss data points in the water absorption band.f From only five control and four dehydrated samples, which are not necessarily the same samples.gFrom scans of only two samples, which did not miss data points in the water absorption band.

observed in the absorption coefficient in the visiblespectrum above 500 nm. However, the reduced scat-tering coefficient increased at all wavelengths as seenin Tables 4 and 5. Representative spectra of changesin the optical properties are shown in Figs. 4(c) and4(d).

It was observed that direct and indirect (i.e., aswrapped in aluminum foil) exposure to a hot salinebath had similar results qualitatively as seen in Figs.5 and 6, respectively. Yet the changes in the opticalproperties were less pronounced when tissue sampleswere first wrapped in aluminum foil and then bathedin an isotonic saline bath. It was also observed thatthe thickness of the samples wrapped in aluminumfoil increased by 16% on average. Microscopic stud-ies revealed increased tissue thickness accompaniedby a reduced surface area. It was found that thereduced scattering coefficient of human aorta changedby 10-40% in the range from 400 to 800 nm and by 30to over 90% above 800 nm when it was bathed in a hotsaline solution wrapped in aluminum foil, whereasthe reduced scattering coefficient of a human aortachanged by 15-45% in the range from 300 to 800 nmand by 30% to over 150% above 800 nm when it was indirect contact to the hot saline.

Reduced Scattering Coefficient as a Function of WavelengthThe average value of n [for Eq. (1)] and the correspond-ing standard deviation for normal, dehydrated, andheated samples is given in Table 6. The significanceof dehydrated and heated samples relative to thecontrol data is also shown in the table. Typically, ifthe significance is < 5%, we reject the hypothesis thatboth sets of data come from the same population.According to this criterion, there is a significantincrease in the reduced scattering coefficient for 70°and 100°C heated specimens. Another t-test per-

formed for 100°C heated samples revealed that thereis no significant difference in the reduced scatteringcoefficient for directly and wrapped heated samples.

DiscussionIn this study we analyzed the effects of dehydrationand protein coagulation on the optical properties ofthe human aorta in vitro. It was found that whenthe average weight loss resulting from dehydrationwas 46%, the absorption coefficient of the humanaorta increased by 20-50% in the 500-1100-nm rangeand somewhat less in the 300-500- and 1100-1350-nm ranges. The relatively large increase in theabsorption coefficient is basically a result of denserpacking of cells owing to shrinkage of the tissuesamples while the number of chromophores remainedconstant. However, at 1350-1550 nm (a water ab-sorption band) and up to 1730 nm, the absorptioncoefficient remained largely unaffected by water loss.Although water is the primary chromophore in thisband, the decrease in water content is balanced bytissue shrinkage yielding approximately the sameabsorption coefficient at the levels of dehydrationthat we tested.

Since dehydration did not significantly change thereduced scattering coefficient, the authors assumedthat the observed optical property changes wereattributed only to water loss; actually dehydrationmay also change collagen cross linking. Indeed itwas reported that water is intimately involved in thecollagen structure.' 9 Electron micrographs of thereplicas of wet collagen showed relatively smoothcylindrical fibrils in contrast to the corrugated appear-ance of dry collagen.20 Furthermore, in our analysisof the power relationship the value of n for normaland dehydrated samples was 1.15 ± 0.10 and 1.22 ±0.13, respectively. Since the significance level was


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Fig. 4. Raw transmission (a) and reflection (b) data from a human aorta heated in a saline bath at 100°C for 300 ± 10 s, correspondingabsorption (c), and reduced scattering spectra (d). There are two different damaged samples. The control sample was 325 pLm thick,whereas the damaged samples were 350 (0) and 450 (+) m thick. Note that the discontinuity at 800 nm is an instrument artifact causedby the detector change. The dips in the reduced scattering spectra are artifacts caused by the -Eddington approximation that cannot giveaccurate properties for regions of high absorption such as the water-absorption band.' 4' 18

-15%, dehydration did not affect the power-lawrelation between the wavelength and the reducedscattering coefficient.

We observed that exposure to a temperature of60-70°C for 300 ± 10 s did not result in predictableand distinct changes in the optical properties as seenin Fig. 515 Our findings are consistent with the factthat collagen denaturation starts dominating tissuebehavior between 550 and 70 0C.2 Tissue progressesfrom normal to denatured states between 60° and75 0C.2123 Thus heterogeneous tissue samples suchas the aorta may have reached different end points atthe end of the 300-s heating period. Although heat-ing tissue to temperatures of 60°C for more than 10s leads to protein denaturation in cells,24 it has beenreported that some changes in the optical propertiescaused by thermal damage are still reversible eventhough the thermal threshold for protein coagulationis exceeded.25' 26

Although coagulation was first suspected of increas-ing both the absorption and the reduced scatteringcoefficients of the human aorta,25 exposure to hotsaline at 100°C for 300 ± 10 s resulted in a slightdecrease in the absorption coefficient in the visiblespectrum and up to 1400 nm and no dramatic changesover 1400 nm. Only when tissue samples werewrapped in aluminum foil before they were bathedwas there a distinct increase in the absorption coeffi-cient at wavelengths of <400 nm. Yet when wecompared the curves representing the average changein the absorption coefficient as a result of thermaldamage at 70° and 100°C, they were similar, as seenin Fig. 5(a). We believe that a slight decrease from500 to 1300 nm may be due to denaturation of tissuechromophores. We have also observed that the influ-ence of saline and chromophore diffusion on theabsorption coefficient for 100°C heated samples wasnegligible above 500 nm. On the other hand the


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Table 2. Optical Properties of Normal and Thermally Damaged (at 60°C for 300 + lOs) Human Aorta at Selected Wavelengths:nnormal = 4, ndamaed = 4

A jp,u Normal [La Damaged Change in p1a X us Normal is' Damaged Change in As'(nm) (cm-,) (cm-1) (%) (nm) (cm-') (cm-') (%)

350 22.7 ± 1.8 24.6 ± 3.6 +11.7 350 46.3 ± 3.5 58.2 + 12.9 +16.2490 5.4 ± 0.5 6.4 ± 1.0 +14.6 490 31.7 ± 2.1 36.5 ± 7.3 +16.0515 4.3 ± 0.6 5.2 ± 1.0 +14.9 515 29.3 ± 2.2 33.7 ± 6.8 +16.8630 2.7 ± 0.8 3.6 ± 0.8 +28.0 630 21.9 ± 2.0 25.0 + 4.9 +18.1

1065 3.3 ± 0.6 4.1 ± 0.8 +23.7 1065 11.9 ± 1.8 13.1 + 2.7 +19.01320 5.5 ± 0.7 5.9 ± 0.8 +9.8 1320 9.0 ± 1.5 9.8 + 2.2 +16.21400 27.1 ± 4 .8 b 22.9 ± 3 .7 b -4.9e 1400 4.3 ± 2.lb 8.5 +.± b +185.0e1450 37.6 ± 0.OC 37.8 ± 7.8c +15.2f 1450 1.2 ± 0.0c 4.4 1.5c +179.2f1500 29.1 ± 4 .6 d 23.5 ± 4 .2 d -6.59 1500 3.1 ± 1.8d 7.7 0.6d +392.991550 15.6 ± 1.3 14.9 ± 2.2 +0.3 1550 6.2 ± 1.3 7.3 ± 2.1 +15.01650 9.0 ± 0.6 9.0 ± 1.3 +1.6 1650 6.9 ± 1.3 7.6 ± 1.8 +14.91750 11.7 ± 0.8 11.5 ± 1.8 +0.9 1750 5.9 ± 1.2 6.8 ± 1.7 +17.3

aThe samples were directly exposed to hot saline. The values are given as mean ± standard deviation. The corresponding tissuethicknesses for normal and thermally damaged samples were 575 ± 115 and 425 ± 75 pum, respectively. Note that normal and damagedsamples were different samples from the same specimen. % change = {[pu (damaged)/,u (control)] - 11100.

bFrom only three control and three damaged samples, which are not necessarily from the same specimen.cFrom only one control and three damaged samples, which are not necessarily from the same specimen.dFrom only three control and three damaged samples, which are not necessarily from the same specimen.eFrom scans of only two specimens, which did not miss data points in the water absorption band.fFrom scans of only one specimen, which did not miss data points in the water absorption band.gFrom scans of only two specimens, which did not miss data points in the water absorption band.

reduced scattering coefficient of the human aortaincreased by 10-45% from 400 to 1300 nm and 30% toover 100% above 1500 nm and considerably less in the300-400-nm range. Assuming that other experimen-tal conditions remained the same, saline and chromo-phore diffusion resulted in an increased reducedscattering coefficient for 100'C directly heated sam-ples, especially above 1300 nm, even though ouranalysis of the power relationship between the wave-length and reduced scattering coefficient as given inEq. (1) revealed no statistical difference in n for 100'Cheated samples. There was a consistent decrease inn with thermal damage at 70° and 100'C as can beseen in Table 6. Since a t-test showed a significance

level of < 5% between n values for the control and 700and 100'C heated samples, this decrease in n wouldrepresent a change in the shape factor for scatterersand scattering anisotropy in Mie theory. Our hypoth-esis is supported by the observations of another studyon optical-property changes during the thermal coag-ulation of myocardium. In this study the analysesbased on Mie theory suggested that the appearance ofcoarse and small thermally coagulated granular cellu-lar proteins could be responsible for the increase inthe reduced scattering coefficient.27 Recently Essen-preis et al.2 8 ,29 pointed to changes in the scatteringphase function in coagulated rat liver. Essenpreis etal. observed a nearly 35% decrease in the anisotropy

Table 3. Optical Properties of Normal and Thermally Damaged (at 70°C for 300 + lOs) Human Aorta at Selected Wavelengths:norma_ = 2, ndamaged = 3

X Iot Normal L, Damaged Change in pi, X R,' Normal As' Damaged Change in ps'(nm) (cm-') (cm-') (%) (nm) (cm-') (cm-') (%)

350 21.0 ± 0.3 17.0 ± 1.8 -18.8 350 65.6 ± 17.1 68.5 ± 6.1 +1.1490 5.6 + 0.8 4.7 + 0.4 -16.3 490 38.9 ± 7.8 41.5 ± 5.4 +1.7515 4.8 + 0.8 4.0 + 0.3 -18.1 515 36.3 ± 7.6 38.4 ± 4.9 +0.8630 3.5 + 0.8 2.9 + 0.1 -18.8 630 27.7 ± 6.2 29.2 ± 3.6 +0.6

1065 3.6 + 0.8 2.9 + 0.1 -21.9 1065 14.1 ± 3.2 16.9 ± 3.1 +13.71320 5.7 + 0.8 4.9 + 0.2 -15.7 1320 10.3 ± 2.2 13.7 + 3.0 +24.81400 24.1 + 1.2 22.6 ± 2.8 -5.1 1400 7.1 ± 3.0 11.2 ± 4.7 +38.51450 39.0 + 0 .Ob 34.0 ± 0 1.b - 12.8c 1450 5.4 ± O.Ob 10.9 + 3 .7 b +101.0c1500 26.1 ± 1.0 23.0 ± 2.7 -11.1 1500 5.7 ± 2.9 10.2 + 4.4 +57.91550 15.4 ± 0.2 13.6 ± 1.1 -11.8 1550 7.3 ± 2.2 11.3 + 3.4 +41.01650 9.4 + 0.6 8.4 ± 0.6 -11.5 1650 7.7 ± 1.8 11.0 ± 3.0 +33.41750 11.8 ± 0.7 10.7 ± 0.8 -9.7 1750 6.9 ± 1.6 10.4 ± 3.2 +37.6

aThe samples were directly exposed to hot saline. The values are given as mean ± standard deviation. The corresponding tissuethicknesses for normal and thermally damaged samples were 388 ± 88 ,um and 458 ± 12 pum, respectively. Note that normal anddamaged samples were different samples from the same specimen. % change = {[u (damaged)/u (control)] - 11100.

bFrom only one control and two damaged samples, which are not necessarily from the same specimen.cFrom scans of only one specimen, which did not miss data points in the water absorption band.


Table 4. Optical Properties of Normal and Thermally Damaged (at 100°C for 300 + lOs) Human Aorta at Selected Wavelengths:nnormaI = 6, ndamag.d = 9

A col. Normal .a Damaged Change in p.a X pi.s Normal p.s' Damaged Change in pus'(nm) (cm-') (cm-,) (%) (nm) (cm-i) (cm'1) (%)

350 18.9 ± 2.8 20.3 ± 2.5 +7.6 350 56.1 ± 7.9 66.3 ± 7.0 +22.2490 6.2 ± 0.9 5.9 ± 0.9 -5.8 490 33.7 ± 3.6 41.4 ± 3.1 +28.9515 5.4 ± 0.7 4.6 ± 0.6 -14.0 515 30.1 ± 3.1 38.4 ± 2.8 +29.6630 4.3 ± 0.6 3.4 ± 0.6 -21.4 630 22.1 ± 2.1 29.9 ± 2.3 +36.7

1065 4.0 ± 0.4 3.4 ± 0.5 -15.1 1065 11.0 ± 1.4 17.5 ± 1.7 +62.01320 5.5 ± 0.4 5.2 ± 0.7 -8.1 1320 8.2 ± 1.6 13.8 ± 1.8 +76.51400 16.3 ± 5.1 17.4 ± 6.2 -2.0 1400 7.1 ± 3.4 12.4 ± 3.5 +127.01450 34.9 ± 4 .3 b 35.9 ± 8.4b - .gd 1450 7.5 ± 5.6b 10.1 8 .6 b +21.5d1500 26.2 3.3c 26.4 ± 5.7c +1.3e 1500 5.9 ± 4.7c 10.6 ± 6.3c +426.6e1550 16.2 ± 2.2 15.4 ± 2.2 -3.5 1550 6.3 + 3.0 11.7 + 3.4 +148.61650 9.6 ± 0.8 9.3 ± 1.0 -2.9 1650 6.3 ± 2.0 11.2 ± 2.2 +97.71750 11.1 ± 0.9 11.1 ± 1.4 -1.5 1750 5.8 ± 2.3 10.6 ± 2.5 +112.1

aThe samples were directly exposed to hot saline. The values are given as mean + standard deviation. The corresponding tissuethicknesses for normal and thermally damaged samples were 305 ± 37 pm and 420 ± 59 pum, respectively. Note that normal anddamaged samples were different samples from the same specimen. % change = ([p (damaged)/p (control)] -1)100.

bFrom only three control and seven damaged samples, which are not necessarily from the same specimen.cFrom only five control and nine damaged samples, which are not necessarily from the same specimen.dFrom scans of only two specimens, which did not miss data points in the water absorption band.eFrom scans of only five specimens, which did not miss data points in the water absorption band.

factor when rat liver was coagulated. His resultsmay well be applicable to the human aorta.

The thermal response of biological tissues to laserlight can be predicted by using an optical model suchas a -Eddington diffusion approximation to deter-mine the fluence rate of light in combination with theheat-conduction equation. However, none of thepublished models so far incorporates dynamic changesin the optical behavior of tissue during laser irradia-tion. The photothermal response of laser-irradiatedtissue is given by the so-called heat diffusion equa-tion4 with the heat source term given by

Q = -a'(z, r), (2)

where pIo is the absorption coefficient of the tissue(cm-') and (z, r) is the local fluence rate of the laserlight (W cm- 2).

Since the thermal response of tissue is directlyproportional to its absorption characteristics and thelocal fluence rates, changes in the optical propertieswill affect the source term. Therefore the localheating patterns will be changed as a function ofhydration and protein denaturation. Since the re-duced attenuation coefficient is given by

LAt = Pa + Ps X (3)

an increase in either of the optical properties will

Table 5. Optical Properties of Normal and Thermally Damaged (at 1 00C for 300 lOs) Human Aorta at Selected Wavelengths:fnoma ndamaged = 5a

X pl.a Normal [.,, Damaged Change in .,, X ps' Normal pu8' Damaged Change in ILs'(nm) (cm-') (cm-) (%) (nm) (cm-') (cm-') (%)

350 15.4 ± 2.3 17.0 ± 1.5 +11.7 350 62.3 ± 7.6 70.9 ± 7.6 +14.7490 4.3 ± 1.1 4.1 ± 0.7 -3.0 490 37.0 ± 4.5 45.0 ± 2.2 +23.0515 3.9 ± 1.0 3.4 ± 0.6 -9.8 515 34.3 ± 4.3 41.9 ± 2.0 +23.3630 3.1 ± 1.0 2.5 ± 0.6 -17.5 630 25.9 ± 3.4 32.5 ± 1.4 +26.6

1065 3.2 ± 0.9 2.6 ± 0.7 -18.7 1065 13.6 ± 1.6 19.2 ± 0.9 +42.41320 4.6 ± 0.9 4.2 ± 0.8 -8.6 1320 10.2 ± 1.2 15.5 ± 1.0 +53.61400 11.7 ± 1.3 11.7 + 0.6 +0.2 1400 9.0 ± 1.0 14.2 ± 1.2 +58.71450 38.1 ± 4 .2 b 42.1 5 .3 b +3 .3 d 1450 1.6 ± 2 0b 5.7 ± 3.8b +65.0d1500 28.6 ± 33c 30.2 + 2.6c +4.3e 1500 4.0 ± 2.7c 9.6 ± 2.4c +130.5e1550 16.3 ± 1.9 16.8 1.0 +3.9 1550 7.2 ± 0.8 12.1 ± 1.3 +69.51650 9.9 ± 1.1 9.2 ± 0.7 +3.2 1650 7.5 + 0.9 12.2 ± 1.1 +64.01750 10.0 ± 1.2 10.4 + 0.7 +5.6 1750 7.0 ± 0.8 11.5 + 1.1 +64.9

aThe samples were wrapped in aluminum foil before being immersed in a hot saline bath. The values are given as the mean ± standarddeviation. The corresponding tissue thicknesses before and after the hot saline bath were 460 ± 116 and 530 ± 124 pum, respectively.The percentage thickness increase caused by thermal damage was 16 ± 5.5%. % change = {[p. (damaged)/p (control)] - 1}100.

bFrom only three control and five damaged samples.cFrom only four control and five damaged samples.dFrom scans of only three specimens, which did not miss the data points in the water absorption band,eFrom scans of only four specimens, which did not miss the data points in the water absorption band.


60

40

20

0

-20

-40

-60' . I I . I I I I I , I , I ,

200 400 600 800 1000 1200 1400 1600 1800

Wavelength (nn)

(a)

160 -

140 -

120 -

100

80 _

60 -

40 -

20 -

0-200

Fig. 5. Effecreduced scattwavelengths.fit with a thirn

result in;attenuatiothe tissuechanges dibehavior, dosimetry l

In Tableples are co:

80 -

70 -

60 -

50 -

40 -

30 -

20

10I

* 490 nE0 515 nm11M| 630 nm3 1065 nmEl 1320 nm

MlM,M%MVA I

Fr

70%C 1000C (indirect) 100'C (direct)

Fig.6. Effects of thermal damage at 700 and 10000 on the reducedscattering coefficient at frequently used laser wavelengths.

and Nd:YAG laser wavelengths reported by van Ge-100

0C (dirmt) imert et al.33 and Lozano are significantly lower than

- -+ the values reported in this study. Whereas the, + - absorption coefficient at 632.8 nm reported by Keijzer

et al.31 is somehow close to the values found in this+ study, the absorption coefficient reported by Yoon35 is

7+ close to the values reported by van Gemert et al.7 < The reduced scattering coefficient at 476, 580, and

+, -'C 600 nm reported by Keijzer et al.3 ' is significantlyhigher than the values reported in this study. Even

_________________;_______________ )though there is reasonable agreement between the,_=_______, values of the reduced scattering coefficient found by

400 600 800 1000 1200 1400 1600 1800 Oraevski et al.32 and in our study at 488 nm, the valueWavelength (nm) at 514.5 nm obtained in our study almost doubles the(b) values reported by van Gemert et al.33 The values

ts of thermal damage on the (a) absorption and (b) for the reduced scattering coefficient at 515 and 633ering coefficients of the human aorta at selected nm found by Lozano35 are close to the values in this

The average change in optical properties was curve study. Values for the reduced scattering coefficientI-order polynomial. at 632.8 nm as reported by van Gemert et al.3 3 are

significantly lower than the values found in ourstudy, whereas Yoon36 and Keijzer et al.3 ' found

a reduced penetration depth, increased higher reduced scattering values at this wavelength.i, and thus higher temperatures closer to At Nd:YAG laser wavelength of 1064 nm, the values

surface. By incorporating dynamic found in this study doubled the reduced scatteringaring laser irradiation in tissue optical coefficient reported by van Gemert et al.,3 3 but theyit is possible to estimate the required are lower than the value reported by Lozano.35

more precisely. Why is there such a variation in the value of the7 optical properties of normal aorta sam- optical properties reported by various laboratories?

impared with published values from other Undoubtedly differences in the measurement tech-research groups. The absorption coefficient at 470nm reported by Prince et al.3 0 (as converted fromKubelka-Munk optical properties) and at 476 and580 nm as reported by Keijzer et al.3 1 are significantlyhigher than the absorption coefficient found in thisstudy. Yet the absorption coefficient at 600 nmreported by Keijzer et al. is in close agreement withthe value found in this study. At argon laser wave-lengths, 488 and 514.5 nm, there is reasonable agree-ment between values of the absorption coefficientfound in this study and those reported by Oraevski etal.32 and van Gemert et al.,3 3 respectively, who usedKubelka-Munk theory. Yet the absorption coeffi-cient at 515 nm reported by Lozano,35 and at He-Ne

Table 6. Power Relationship Between the Wavelength and the ReducedScattering Coefficient and the Significance of n Values for Control and

Experimental Reduced Scattering Spectra as Obtained from a t-Test

SignificanceDescription ncontrol nexperimental (%)

Dehydration 1.15 ± 0.10 1.22 ± 0.13 - 15Heatingat60'C 1.21 ± 0.12 1.28 ± 0.04 -25Heating at 70'C 1.30 ± 0.01 1.10 ± 0.10 < 5Heating at 1000C

(Direct heating) 1.38 ± 0.11 1.06 ± 0.07 <5Heating at 1000C

(Wrapped heating) 1.26 ± 0.08 1.03 ± 0.05 < 5


Table 7. Published Optical Properties of Normal Human Aorta atSelected Wavelengths Compared with Values Obtained In This Study

X (nm) I.La (cm' ) p,' (cm') Methods and Referencea

470 13 2 - KM3 0

5.9 + 1 .2b 37.8 6 .1 b DA476 7.5 39.20 DA (d = 250 pum) 31

26.1 58.76 DA(d = 233 pum)31

10.8 38.56 DA (d = 375 pm)31

5.7 ± 1 2b 36.8 + 6 4 b DA488 3.9 30.23 KM3

1

5.2 ± 2 .1 b 35.9 ± 5.8b DA514.5 7.0 16.67 KM (d = 1.4 mm)33

6.25 16.75 KM (d = 1.88 mm)38

5.0 17.67 KM (d = 1.23 mm)33

4.3 ± 1 0 b 33.0 ± 54b DA515 2.7 32.6 DA33

580 6.6 27.75 DA (d = 250 pM)31

9.2 45.75 DA (d = 233 pum)31

11.0 28.96 DA(d = 375 pLm)313.5 ± 1.1b 27.9 ± 4 .9 b DA

600 4.1 28.48 DA (d = 250 pm)31

4.7 28.32 DA (d = 233 pum) 31

3.3 26.85 DA (d = 375 pm)31

3.4 ± 1.1b 26.6 ± 4 6 b DA632.8 0.75 10.12 KM (d = 1.4 mm)33

0.70 9.17 KM (d = 1.88 mm)33

0.55 9.25 KM (d = 1.23 mm)33

2.3 40.95 DA30

0.52 31.0 ADT34

3.3 ± l.1b 24.6 ± 4.4b DA633 1.4 26.0 DA33

1060 1.17 17.0 DA33

1064 0.30 6.77 KM (d = 1.28 mm) 33

0.30 6.77 KM (d = 1.47 mm)33

0.45 4.15 KM (d = 2.87 mm)33

3.5 ± 0 .8b 13.2 ± 2 6 b DA

aKM, Kubelka-Munk properties converted by using34 AKM = 2p.a, SKM = (3 ps' - 11a)/4; DA, diffusion approximation; ADT,asymptotic diffuse transmission.

bMean ± standard deviation from 22 samples.

nique and the difficulty in performing an experimentthat matches the mathematical theory are a majorsource of error. Although little information hasbeen published on the sensitivity of a and p,' toerrors in the measurement of R and T, addingdoubling computation by Prahl'4 suggests that a5-10% change in total reflection can change thecalculated total attenuation coefficient, p = a + s,by a factor of 2 for g = 0.875, a = 0.9, a =

sll/(s + a). Thus a small error in the measure-ment of reflection can produce a factor-of-2 change inthe albedo a. In this situation it appears that thescattering coefficient is extremely sensitive to errorsin the measurement of reflection.

During this study we have noted that measure-ments with a spectrophotometer system predictgreater values of the absorption coefficient than anequivalent laser-based integrating sphere system.Furthermore there is considerable variability in themeasurement of the absorption coefficient below 3cm-1 as a function of tissue thickness and unac-counted losses associated with the integrating sphere

measurements. 3 353 7 The data of Keijzer et al.3 ' andvan Gemert et al.33 as presented in Table 7 andmeasurements in our laser laboratory support theobservation that the calculated values of the absorp-tion coefficient decrease as the sample thicknessincreases.13 1537 Undoubtedly, the variability in theseresults is partly due to the nonhomogeneity of thetissue. Yet we can improve the results by minimiz-ing the effects of other sources of error. With tun-able laser sources the beam divergence associatedwith spectrophotometer systems can be reduced.Integrating sphere measurements can also be im-proved with a double integrating sphere system.Measurements should be corrected for various lossesas described by Cheong37 and Pickering et al.3 8

Tissue samples must be fresh and uniform, anddehydration must be avoided, as noted above.

ConclusionThe optical properties of the human aorta are afunction of hydration levels and previous thermaldamage. A water loss of 50% increases the absorp-tion coefficient by 20-50% in the visible spectrum,whereas thermal damage induced at 100°C increasesthe reduced scattering coefficient by 10% to over150%. Neither the power of the wavelength depen-dency of the reduced scattering coefficient [see Eq. (1)]nor the reduced scattering coefficient is significantlyaltered by dehydration or heating at 60°C. However,the value of n decreased by 16-24% as a result ofheating at 70° and 100°C, and this change wasaccompanied by a significant increase in the reducedscattering coefficient.

The authors thank Roberto Bayardo and his assis-tants at the Travis County Medical Examiner's Officefor providing the tissue samples.

This work was supported in part by the U.S. Officeof Naval Research under grant N00014-91-J-4065and in part by the Albert and Clemmie Caster Foun-dation.

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