Lasers in Surgery and Medicine 11:141-151 (1991)
Co:MgF, Laser Ablation of Tissue: Effect of Wavelength on Ablation Threshold and Thermal Damage Kevin T. Schomacker, PhD, Yacov Domankevitz, MS, Thomas J. Flotte, MD, and Thomas F. Deutsch, PhD Wellman Laboratories of Photomedicine, Department of Dermatology, Massachusetts General Hospital, Boston 021 14
The wavelength dependence of the ablation threshold of a variety of tissues has been studied by using a tunable pulsed Co:MgF, laser to determine how closely it tracks the optical absorption length of water. The Co:MgF, laser was tuned between 1.81 and 2.14 pm, a wavelength region in which the absorption length varies by a decade. For soft tissues the ablation threshold tracks the optical absorption length; for bone there is little wavelength dependence, consistent with the low water content of bone. Thermal damage vs. wavelength was also studied for cornea and bone. Thermal damage to cornea has a weak wavelength dependence, while that to bone shows little wavelength dependence. Framingcamera pictures of the ablation of both cornea and liver show explosive removal of material, but differ as to the nature of the explosion. Key words: pulsed infrared laser albation, tissue damage, cornea, bone, aorta, tissue denaturation, optical properties of water
first approximation, the optical and thermal properties of tissue can be taken t o be those of liquid We report measurements of ablation thresh- water. The simplest models treat ablation as the olds and tissue damage vs. wavelength by using rapid (explosive) vaporization of water [ll, while the tunable, pulsed Co:MgF, laser. Tissue abla- more complex models consider the dynamics of tion with pulsed infrared (IR) lasers has been vaporization [161 and may include a layer of liqstudied by a number of workers using a variety of uefied material between the tissue being ablated laser sources ranging in wavelength from 2 to and the surrounding atmosphere [5,6]. The pre10.6 pm and in pulsewidth from less than 100 ns dictions of ablation models can be compared to to over 100 ps [l-151. Studies using pulsed and measurements of a number of parameters, includcontinuous-wave CO, lasers have shown that ing the ablation threshold, the mass-loss rate, and thermal damage can be reduced by using pulsed the momentum transferred to target tissue. lasers. Thermal damage zones as small as a few Thermal damage caused by denaturation of microns thick can be obtained at wavelengths tissue at the bottom and sides of the ablation crathat are strongly absorbed in tissue by using ter is of considerable clinical interest, since it can pulses that are short compared t o the thermal re- affect the speed of wound healing. The simplest laxation times of the laser-heated volume; for ex- thermal models of tissue damage calculate the ample, damage zones of 5-10 pm have been found depth to which tissue is heated to the denaturin the cornea following ablation using a Qswitched Er:YAG laser operating at 2.94 pm [41, the wavelength of the peak of the water absorp- Accepted for publication December 11, 1990. tion band. Address reprint requests to Thomas F. Deutsch, Wellman Modelling of pulsed IR laser ablation has Laboratories of Photomedicine, Department of Dermatology, generally proceeded on the assumption that, t o a Massachusetts General Hospital, Boston, MA 02114. INTRODUCTION
0 1991 Wiley-Liss, Inc.
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cient of liquid water, taken from data of Curcio
-rl and Petty 1191, together with the tuning range of
i -
C o Mgl', Range
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Fig. 1. Mid-infrared absorption spectrum of water showing the 1.94-)~m peak of interest (taken from ref. 19).The dashed line shows the range covered in these experiments.
the pulsed Co:MgF, laser. The absorption coefficient has a peak of 114 cm-' at 1.94 pm, and drops t o less than 10 cm-' at 1.85 pm and t o about 16 cm-' at 2.2 pm. Thus the optical penetration depth can be varied by over a decade by tuning the output wavelength of the Co:MgF, laser. The primary goal of these experiments was t o examine the effect of optical penetration depth on ablation threshold and thermal damage for a variety of tissues. In addition, since some of the experiments employed a high-speed framing camera in the determination of ablation threshold, information about the dynamics of material removal was also obtained. MATERIALS AND METHODS
ation temperature. This approach predicts damage only at the bottom of the crater, while in practice a layer of damaged tissue roughly uniform in thickness is observed at the walls and the bottom of the crater. More complex thermal models include a layer of liquefied material between the tissue and the surrounding atmosphere; the radial forces on the liquid layer due t o a nonuniform laser beam profile cause the liquid layer to be ejected from the crater, flowing along the crater walls. This liquid layer lining the crater acts as a thermal reservoir which can transfer heat t o the crater walls, leading t o damage [17]. Thus the existence of a liquid layer during the ablation process accounts for the geometry of the observed thermal damage. Most of the ablation studies to date have been performed with a laser operating at a single frequency; some studies using the line-tunable HF laser have also been reported [10,111. Comparisons between such studies are sometimes difficult because of variations in beam profile and pulse shape, as well as because of possible differences in methods of beam characterization. Measurements using a broadly tunable IR laser can eliminate some of these difficulties and provide additional data to improve our understanding of ablation, provided the absorption coefficient of water varies significantly, by at least a decade, over the laser's tuning range. The pulsed Co:MgF, laser, which can be tuned from 1.75 t o 2.5 pm, is well suited for studying the way in which the optical penetration depth affects tissue ablation and tissue damage 1181. Figure 1shows the optical absorption coeffi-
The Co:MgF, laser (Schwartz Electro-Optics, Concord, MA) was optically pumped by a Nd:YAG laser emitting at 1.338 pm. Typical output energies were 70 mJ in a train of 2-ps pulses approximately 100 ps long. The details of the pulse shape varied with wavelength and pump conditions. For example, at 1.91 pm the pulse train consisted of 5 pulses, each returning t o baseline, while at 2.0 pm several micropulses were superimposed on the envelope of a 100-ps macropulse. Such differences can be important in quantitative modelling of thermal damage, but were not documented in detail here. The spot size at the tissue was obtained from the beam profile, which was determined by scanning a pyroelectric detector (model 53, Molectron, Campbell, CA), with a 50pm pinhole in front of it, in the target plane. Readings were taken at 25-pm intervals and processed by computer t o obtain a three-dimensiona (3-D) plot of the beam profile, which was nearly Gaussian. As discussed below, a cross-section through this 3-D plot was used to obtain the Gaussian spot size of the beam. Two methods were used to obtain Fth, the threshold fluence for ablation. The first, shown schematically in Figure 2a, was an optical pumpprobe technique; a He-Ne laser beam, focused to a spot size of 60 pm, passed approximately 80 pm in front of and parallel to the surface of the sample and was detected by a Si photodiode. An x-y-z stage was used t o accurately position the sample surface with respect to the He-Ne probe beam. The onset of tissue ablation resulted in a reduction of the photodiode signal, due to either the deflection of the beam by refractive index
Co:MgF, Laser Ablation of Tissue
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Fig. 2. a: Experimental schematic of the pump-probe system for determining ablation threshold. b: Experimental schematic of the framing-camera system for determining ablation threshold. The He-Ne laser is used to position the tissue sample reproducibly.
changes in the air near the tissue or to a blocking of the beam by ejected tissue. The temporal profile of the Co:MgF, laser pulse was obtained using
a fast HgCdTe photoconductive detector. The output of both detectors was fed to a digital oscilloscope (model 9400, LeCroy Corp., Chestnut Ridge,
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NY); comparison of the two traces allowed one t o determine when during the laser pulse ablation started. The signal-processing capabilities of the digital oscilloscope were used t o measure the integrated energy delivered at the time ablation started. The Co:MgF, laser energy was focussed onto the tissue surface by a 20-cm-focal-length lens, resulting in a l/e2 spot diameter, 20, of 255 pm as determined by fitting a cross-sectional slice through the 3-D beam profile to a Gaussian function, f(r) = exp(-2r2/a2). Since, at threshold, ablation begins at the peak of the Gaussian beam, the peak fluence, which can be shown t o be twice the total energy in the pulse divided by m r 2 , is plotted in the figures below. The second technique for measuring ablation threshold, shown schematically in Figure 2b, involved the use of a framing camera (Imacon HE 700, Hadland Photonic, Cupertino, CA) t o obtain a series of pictures of the tissue target, viewed at 90" to the ablating laser beam axis, during the ablation process [81. The target was illuminated by a pulsed flash lamp having an output pulse several hundred microseconds long. The He-Ne laser and associated lenses were used t o position the sample reproducibly. The magnification of the camera was tenfold. The exposure time of a single frame was approximately 200 ns; a series of 10 or more photographs separated by intervals of 1 ps was taken and the onset of ablation determined from the photographs. That is, the time of the first photo showing material removal was determined. A delay circuit was used to vary the time delay between the onset of the laser pulse and the start of picture-taking. The output of the HgCdTe detector was sent into one channel of the digital oscilloscope and pulses from the camera, synchronized with the gating of the electronic shutter, were sent to the other channel. The Co:MgF, laser energy was focussed onto the tissue surface by a 30-cm-focal-length lens, resulting in a lie2 spot diameter of 368 pm, as determined by fitting the measured profile to a Gaussian function. The threshold energy was determined by integrating the laser pulse profile up t o the appropriate camera "sync" pulse. Ablation threshold measurements were made in vitro on samples of bovine calf cornea, chicken liver, human aorta, and rabbit bone. In addition, threshold measurements were also performed on collagen gels (gelatin) having 85% water content; such gels are often used as models for tissue. The ablation thresholds of fatty plaque and calcific plaque were also examined. All abla-
01 1.8
1.9
2.0
2.1
2.2
Wavelength (pm)
Fig. 3. The dependence of ablation threshold on wavelength for collagen gels containing 85% water. The data were recorded with the optical pump-probe technique only. The solid line represents a least-squares fit assuming that F,, = a h .
tion threshold measurements were made using a single laser pulse, thereby avoiding changes in the tissue, such as dessication, which can occur with multiple pulses. Thermal damage measurements were made on samples of calf cornea and rabbit ulna bones. Single-pulse ablations were performed at different wavelengths and fluences. The zone of thermal damage at the bottom of the ablation crater was determined by light microscopic examination. Cornea specimens were fixed in either formalin or Bouin's solution, routinely processed, and stained by using hematoxylin and eosin (H&E). Bone samples were fixed in 80% ethanol for 48 hours at 4"C, transferred t o a saturated aqueous solution of ethylenediaminetetraacetic acid (EDTA), and kept there until the sample was decalcified, approximately 3 weeks. Samples were then embedded in paraffin, sectioned, and stained with H&E. RESULTS
Figures 3-8 show the ablation thresholds vs. wavelength for gelatin, calf cornea, chicken liver, human aorta, diseased human aorta, and rabbit ulna bone respectively. The data cover a region in which the absorption coefficient of water, a,varies by over a decade, and include measurements on both sides of the 1.94 pm peak. As has been discussed in the literature [ll, Ftha is the energy per unit volume deposited at the surface at ablation threshold, Fth, and can be identified with a heat of ablation, Habl.This heat of ablation is en-
Co:MgF, Laser Ablation of Tissue
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Fig. 4. The dependence of ablation threshold on wavelength for resected calf corneas in situ. The data were recorded using the optical pump-probe technique (open circles) and framing camera technique (solid circles). The solid curve represents a least-squares fit to the optical pump-probe data assuming that Fth = a h .
ergy/volume that must be deposited at the surface in order t o cause a deflection of the surface (in the framing-camera measurements) or to expel some tissue fragments (in the pump-probe measurements). Provided the heat of ablation is wavelength independent and that scattering does not significantly change the fluence below the surface, Fth should vary as l/a.Both these assumptions will be considered in more detail below. The solid lines in Figures 3-7 and the dashed line in Figure 8 are plots of a/a, with “a”obtained from a least-squares fit to the data using published values of (Y [191. Figures 4 and 5 combine the results of the pump-probe and framing-camera techniques for obtaining the ablation threshold for calf cornea and chicken liver. In particular, the results using cornea show excellent agreement between the two measurements. For liver, there is a systematic increase in the difference between the two techniques with increasing wavelength in the region 1.95 to 2.15 pm. This may be explained by examination of framing-camera pictures (vide infra) for ablation with these wavelengths. In particular, as the wavelength increases beyond 1.95 pm, the time between the observation of a deflection of the tissue surface, considered threshold for framing camera measurements, and the expulsion of tissue fragments, increases. The determination of thresholds using the pump-probe technique is sensitive to this time delay, resulting in larger threshold values. The absorbed energy per unit volume decreases as the laser is tuned off the wa-
1.9
2.0
2.1
2.2
Wavelength ( p m ) Fig. 5 . The dependence of ablation threshold on wavelength for resected chicken livers. The data were recorded using the optical pump-probe technique (open circles) and framing camera technique (solid circles). The solid curve represents a least-squares fit to the framing camera data assuming that Fth = a h
h
1.8
1.9
2.0
2.1
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Wavelength ( p m ) Fig. 6. The dependence of ablation threshold on wavelength for resected human aortas. The data were recorded with the optical pump-probe technique only. The solid line represents a least-squares fit assuming that Fth = a h .
ter absorption band, leading one to expect slower ablation dynamics. The threshold for bone, in which water is a minority constituent, shows a relatively weak variation with wavelength; the actual water content of bone has been reported to vary from 34% by weight for young bone to between 16.6 and 22% for established bone of varying age [ZOI. We expect that rabbit ulna is probably close to young bone in water content. The dashed curve in Figure 8 is proportional t o (Y-’ and shows that, as expected, the threshold data do not track the water absorption. The solid curve represents a leastsquares fit of the data by a simple model assum-
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TABLE 1. Heats of Ablation for Gelatin and Various Tissue TvDes ____ ~
Tisssue type Gelatin Calf cornea Chicken liver Human aorta Fatty plaque Calcified plaque Rabbit bone (ulna) 1.8
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2.2
Wavelcngth ( p m )
a” (J/cm3) 2,901 865 562 738 1,002 1,205 4,200b
%water ( x 0.01) 0.85 0.80 0.79 0.79 ? ? 0.30
al%water (J/cm3) 3,410 1,080 712 932 -
14,000
“Found by fitting the data to F,, = a/a(X). bThe equation, Fth = 2,0401(0.3xa(h)+a’), where a is the absorption coefficient of water and a’ ( = 18 crn-l) is from an unknown chromophore, gives a better fit.
Fig. 7. The dependence of ablation threshold on wavelength for resected diseased human aortas. The data were recorded with the optical pump-probe technique only. The solid line represents a least-squares fit to the fatty plaque data assuming that Fth = a h .
fitting the Fth vs. wavelength curves for the materials discussed. As discussed above “a” may be identified with a heat of ablation for the initiation of material removal. The last column gives “a” 300 . _ . . corrected for the varying water content of the difFth(h) = 4200/a(h) ferent tissues and should approximate the heat of F,,(h) = 2040/(.3xa(h)+18) vaporization of water, 2500 J/cm3, if ablation is essentially explosive vaporization of the tissue 200 water. Note that the heat of ablation for normal aorta is lower than that of either fatty or calcific plaque, indicating a potentially undesirable selectivity for normal tissue in angioplasty applications. However, all values were determined from single-pulse measurements and the results obtained from multiple pulses used clinically may O C converge due t o the sample dessication [9]. 1.8 1.9 2.0 2.1 2.2 Figures 9 and 10 display thermal damage t o Wavelength (pm) calf cornea as a function of wavelength at a fluFig. 8. The dependence of ablation threshold on wavelength ence of 7.3 ? 2.6 times Fth and as a function of for the resected rabbit ulna bones. The data were recorded fluence at a wavelength of 1.94 pm, the peak of with the optical pump-probe technique only. The dashed line the water absorption curve. Damage zones as represents a least-squares fit assuming that F,, = a h . The small as 150 pm are obtained at 1.94 pm. Note solid curve represents a modified fit to the data assuming that that the use of Bouin’s fixative rather than fora second chromophore having a constant absorption coeffimalin led t o a systematically higher damage zone cient exists (see text). measurement. The thermal damage t o bone varied between 160 and 205 pm in the wavelength ing two different absorbers in bone. Specifically, region 1.81 t o 2.14 pm, with an average value of the model assumes that 30% of the laser energy is 180 15 pm; there was no significant dependence absorbed by water with the remaining energy ab- on wavelength. Figure l l a and b shows two series of framsorbed by a second chromophore, such as collagen or hydroxyapatites, which has a wavelength-in- ing-camera pictures of the ablation of cornea, dependent absorption coefficient in the wave- while Figure 12 shows a similar series for chicken length region studied. The fitting procedure gives liver. Time increases in 2-ps increments from left Fth = 2040/[0.3 X &(A) + 181 J/cm2. The first t o right and in 1 - p s increments from bottom to term, the water contribution, is about 34 cmpl at top; the beam comes from the right. One striking 1.94 p m , compared t o 18 cm-’ for the wave- difference in the pictures for ablation at threshold length-independent term. is that the ablation of the cornea shows the forTable 1 tabulates the value of “a” found by mation of a cavity, which finally breaks. By con-
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Co:MgF, Laser Ablation of Tissue DISCUSSION
147
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A primary goal of this study was to use the tunability of the Co:MgF, laser t o study the wavelength dependence of the ablation threshold and the thermal damage in order to determine how well simple thermal models based on the vaporization of water describe tissue ablation. For the case of gelatin, cornea, chicken liver, human aorta, fatty plaque and calcific plaque the ablation threshold tracks l/a reasonably well. No such relationship is observed for bone, which contains 17 I to 34% water, and none would be expected. This 19 20 21 result is in agreement with data presented in refWavelength (,urn) erence 11, which show negligible threshold deFig. 9. The dependence of thermal damage on wavelength for pendence .with wavelength- fir bone within the resected calf corneas. The measurements were taken from 2.7-2.9-pm absorption band of water. Note that tissue samples fixed in formalin (open triangles and open while both gelatin and cornea track l/a, the acsquares) and Bouin's solution (filled traingles). Except for the tual value of Hablis almost a factor of three higher open squares, all fluences were 7.3 2 2.6 times the ablation for gelatin, indicating that gelatin does not bethreshold found at each wavelength (see Fig. 4). The open squares used fluences of 2.6, 3.0, 12.8, and 36 times the ab- have exactly like cornea, despite the fact that lation thresholds found at 1.81, 1.85, 1.89, and 1.93 pm re- their collagen and water contents are similar. spectively. The solid line represents a least-squares fit to the The analysis of ablation used here has asopen and filled triangles, assuming that the damage depth sumed that 1) the effects of scattering are neglivaries as 01-l. gible at the wavelengths used and that 2) the heat of ablation does not vary with wavelength. The effect of scattering will vary from tissue t o tissue. Cornea, which is optically clear in the visible porE 3 500 400 tion of the spectra, would not be expected to exhibit significant scattering in the 2-pm region. The role of scattering in liver can be examined by 300 using recent data. The absorption and scattering coefficients for liver have been measured in the 200 range 350-2,200 nm [all. For situations where the absorption coefficient,l pa, is equal to or 100 greater than the scattering coefficient, ps, the propagation of light is described by the transport 0 coefficient, pt = pa ps(l-g), where g is the anisotropy parameter [22]. Scattering can be neglected if the second term is small compared to the first one. At 1.94 pm, pa = 73 cm-l and Fig. 10. The dependence of thermal damage found at 1.94 pm = 3.8 cm-'. Here the condition for nepJ1-g) on fluence for resected calf corneas. The line shown is a linear regression analysis of the data showing a small (0.17 km- glecting scattering is well satisfied. However at cm2/J) increase in thermal damage with increasing fluence. 1.8 pm, pa = 11 cm-' and pJ1-g) = 3.8 cmP1 The dotted line denotes the averaged thermal damage. and the condition is only weakly satisfied. The values of ps obtained from reference 21 should be used with caution, since diffusion theory, used to trast, the pictures of chicken liver ablation show analyze the data, is not appropriate near 1.94 pm what may be multiple bubbles forming. The picwhere absorption is greater than scattering. One tures do not allow us t o decide if the cavities are consequence of scattering is t o increase the subempty, full of vaporized water, or full of liquid (water or liquefield collagen). Both figures illustrate how the onset of surface disturbances, pre- 'pa is the same as a used earlier in the text. The change of sumably the onset of ablation, can be determined notation here is to agree with the notation in references 21 and 22. t o within about 1 ps.
,-.
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Fig. 11. Two series of framing-camera photographs of the ablation of cornea taken early (a)and late (b) in the ablation process. The wavelength was 1.91 pm and the fluence 71.2 J/cm2. The exposure time for each photograph was 200 ns.
The bottom left frame corresponds to the earliest time; the laser beam enters from the right. Time increases from left to right and from bottom to top with the odd frames depicted in the bottom row.
surface irradiance in the tissue above the incident in the wavelength region of interest have not irradiance, an effect that would lead t o an under- been measured. estimation of the true ablation threshold. For the to the ablation The fits of Fth = H,,/a case of bone and plaque the scattering coefficients threshold vs. wavelength curves resulted in val-
Co:MgF, Laser Ablation of Tissue
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Fig. 12. A series of framing-camera photographs of the ablation of chicken liver. The wavelength was 1.95 km and the fluence 61.4 Jicm'. The exposure time for each photograph was 200 ns. The bottom left frame corresponds to the earliest time; the laser beam enters from the right. Time increases from left t o right and from bottom to top with the odd frames depicted in the bottom row.
ues of Hab, on the order of 1 kJ/cc, considerably less than the 2.5 kJ/cc required t o vaporize water. This result is similar to that found in studies of the ablation of skin using pulsed CO, TEA lasers; Ftha obtained from a linear fit t o mass loss vs. fluence curves was 0.57 kJ/cc, compared 2.37 kJ/ cc from a Beer's law fit [l].The reason for the difference is that the linear fit approach measures the lowest fluence at which some material is removed. The Beer's law fit to mass-loss data results in a higher threshold value which represents the threshold for explosive, "blow-off" material removal. Our measurements of threshold using either the pump-probe or framingcamera approach detect the beginning of ablation and so correspond to the threshold determined by using a linear fit to mass loss vs. fluence data. The fact that the Habl value for liver is the lowest of all the tissues studies is consistent with studies of the effect of mechanical properties on pulsed CO, laser ablation of tissue 1151. These studies found a heat of ablation for liver, calculated from the slope of a mass-loss curve, of 606 J/gm. This heat of ablation was the lowest of the tissues studied and was correlated with a very low tensile strength compared t o these other tissues. The cornea was used as a model system for thermal damage to tissue because its homoge-
neous nature eliminated some of the variability found in studies of damage t o skin. The damage is less than 100 pm at 1.92 pm and increases as the wavelength is tuned off the absorption peak t o 400 pm at 2.1 pm. It is important t o note that these damage measurements are for single-pulse exposures; multiple pulses have been shown to result in increased damage [6,171. The simplest model of damage, which assumes that damage occurs when the initial temperature distribution in the tissue reaches the denaturation temperature and calculates that depth by using the heat of ablation deposited at the surface, predicts a damage depth given by A/cx 121. In reference 2 it is shown that A = In [Fthcx/(T,-T0)pc], where To is the initial tissue temperature, T, is the critical temperature for denaturation, p is the density, and c is the specific heat. Using the heat of ablation for cornea given in Table 1, T, = 65"C, To = 20°C, and taking the absorption coefficient of cornea t o be 80%of that of water, one finds that A = 1.9. The solid line in Figure 9, given by l/a,is a fit to the data points considered most reliable, given by the filled and open traingles. This fit corresponds to A = 0.8 when the value of cx for cornea is taken to be 0.8 that of water. The measured damage zone shows a weaker wavelength dependence than predicted by the model. Note that the
Schomacker et al. damage at wavelengths shorter than the 1.94-pm tological sections. We thank Peter Moulton and absorption peak does not increase as rapidly as David Welford for many helpful technical discusthe l/a curve would require. Since the cornea ini- sions about the Co:MgF, laser and David Rines tially has negligible scattering, it seems unlikely for technical support. This work was supported by that the simple model fails because it neglects a subcontract from Schwartz Electro-Optics under scattering, although collagen denaturation dur- a National Institutes of Health SBIR program ing the ablating pulse could introduce a time-de- and by the SDIO-MFEL program under Office of pendent scattering. Some evidence for such time- Naval Research Contract N00014-86-K-0017. dependent changes in the optical properties of tissue does exist; the reflectance of liver has been shown to change during ablation with a pulsed REFERENCES 2.1-pm holmium laser 1231. The slight increase in 1. Walsh JT, Deutsch TF. Pulsed CO, laser tissue ablation: Measurement of the ablation rate. Lasers Surg Med 1988; damage with increasing fluence shown in Figure 8:264-275. 10 is not predicted by the model of reference 2. 2. Walsh JT, Flotte TJ, Anderson RR, Deutsch TF. Pulsed More extensive measurements are needed t o obCO, laser tissue ablation: Effect of tissue type and pulse tain sufficient data t o test a more complex abladuration on thermal damage. Laser Surg Med 1988; 8: tion model; however, the usefulness of the tunable 108-118. Co:MgF, laser in obtaining data to test models of 3. Walsh JT, Deutsch TF. Er:YAG laser ablation of tissue: Measurement of ablation rates. Lasers Surg Med 1989; thermal damage is clear. Clinically, the ability to 9:327-337. vary the damage by tuning the wavelength may 4. Walsh JT, Flotte TJ, Deutsch TF. Er:YAG laser ablation be useful, allowing the damage zone t o be adof tissue: Effect of pulse duration and tissue type on therjusted t o the minimum required to obtain hemomal damage. Lasers Surg Med 1989; 9:314-326. 5. Zweig AD, Weber HP. Mechanical and thermal paramestasis. ters in pulsed laser cutting of tissue. IEEE J Quant ElecThe framing-camera pictures demonstrate tron 1987; QE-23(10):1787-1793. the role of tissue mechanical properties in abla6. Frenz M, Romano V, Zweig AD, Weber HP, Chapliev NI, tion. Liver, with its low tensile strength, is clearly Silenok AV. Instabilities in laser cutting of soft media. J removed explosively. In cornea the strength of Appl Phys 1989; 66:4496-4503. 7. Nishioka NS, Domankevitz Y, Flotte TJ, Anderson RR. the surface membrane, Bowman’s layer, appears Ablation of rat liver, stomach and colon with a pulsed to be sufficient t o delay explosive removal until holmium laser. Gastroenterology 1989; 96:831-837. it is ruptured by the expansion of underlying 8. Domankevitz Y, Nishioka NS. Measurement of laser absteam. lation threshold with a high-speed framing camera. IEEE In summary, the use of framing-camera and J Quantum Elec 1990; In press. pump-probe techniques t o determine tissue abla- 9. Nishioka NS, Domankevitz Y. Comparison of tissue ablation with pulsed holmium and thulium lasers. IEEE J tion thresholds for wavelengths near the 1.94-pm Quantum Elec 1990; In press. water absorption band where the absorbance var10. Izatt JA, Albagli D, Itzkan I, Feld MS. Pulsed laser abies tenfold has been investigated. The results lation of calcified tissue: Physical mechanisms and funshow that the ablation thresholds for soft tissue damental processes. In: Jacques SL, ed. “Laser-Tissue Intrack l/a reasonably well. The equivalence in teraction.” Proc SPIE 1202, 1990; pp 133-140. thresholds obtained from these two techniques 11. Izatt JA, Sankey ND, Partovi F, Fitzmaurice M, Rava RP, Itzkan I, Feld MS. Ablation of calcified biological should be stressed because of the simplicity of the tissue using pulsed hydrogen fluoride laser radiation. pump-probe experiment. The effects of waveIEEE J Quantum Elec 1990; In press. length on thermal damage have also been inves- 12. Nelson JS, Yow L, Liaw LH, MacLeay L, Zavar RB, Orentigated for cornea and show only a weak depenstein A, Wright WH, Andrews JJ, Berns MW. Ablation of bone and methacrylate by a prototype mid-infrared erdence on l/a; for bone the damage zone is bium:YAG laser. Lasers Surg Med 1988; 8:494-500. wavelength independent. While simple thermal 13. Nuss RC, Fabian RL, Sarkar R, Puliafito CA. Infrared models predict the dependence of ablation threshlaser bone ablation. Lasers Surg Med 1988; 8:381-391. old on absorption coefficient reasonably well, an 14. Stern D, Puliafito CA, Dobi ET, Reidy WT. Infrared laser understanding of thermal damage will require surgery of the cornea: Studies with a Raman-shifted Nd: YAG laser at 2.80 and 2.92 pm. Ophthalmology 1988; further experimentation and modelling. 150
ACKNOWLEDGMENTS
We gratefully acknowledge the technical support of Margo Goetschkes in preparing the his-
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Co:MgF, Laser Ablation of Tissue MJ, Harpole GM, eds. “Thermal and Optical Interactions with Biological and Related Composite Materials.” Proc SPIE 1064,1989; pp 133-140. 17. Zweig AD, Meierhofer B, Muller OM, Mischler C, Romano V, Frenz M, Weber HP. Lateral thermal damage along pulsed laser incisions. Lasers Surg Med 1990; 10(3):262-274. 18. Welford D, Moulton PF, Room-temperature operation of a Co:MgF, laser. Opt Lett 1988; 13:975-977. 19. Curcio JA, Petty CC. The Near-Infrared Absorption Spectrum of Liquid Water. J Opt SOCAm 1951; 41:302-304.
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20. Stein I, Stein RO, Beller ML. “Living Bone in Health and Disease.” Philadelphia: J.B. Lippincott, 1955: 52. 21. Parsa P, Jacques SL, Nishioka NS. Optical properties of rat liver between 350 and 2200 nm. Appl Optics 1989; 28(12):2325 -2330. 22. Cheong WF, Prahl SA, Welch AJ. A review of the optical properties of biological tissues. IEEE J Quantum Elec 1990; In press. 23. Nishioka NS, Domankevitz Y. Reflectance during pulsed holmium laser irradiation of tissue. Lasers Surg Med 1989; 9:375-381.
Co:MgF, Laser Ablation of Tissue: Effect of Wavelength on Ablation Threshold and Thermal Damage Kevin T. Schomacker, PhD, Yacov Domankevitz, MS, Thomas J. Flotte, MD, and Thomas F. Deutsch, PhD Wellman Laboratories of Photomedicine, Department of Dermatology, Massachusetts General Hospital, Boston 021 14
The wavelength dependence of the ablation threshold of a variety of tissues has been studied by using a tunable pulsed Co:MgF, laser to determine how closely it tracks the optical absorption length of water. The Co:MgF, laser was tuned between 1.81 and 2.14 pm, a wavelength region in which the absorption length varies by a decade. For soft tissues the ablation threshold tracks the optical absorption length; for bone there is little wavelength dependence, consistent with the low water content of bone. Thermal damage vs. wavelength was also studied for cornea and bone. Thermal damage to cornea has a weak wavelength dependence, while that to bone shows little wavelength dependence. Framingcamera pictures of the ablation of both cornea and liver show explosive removal of material, but differ as to the nature of the explosion. Key words: pulsed infrared laser albation, tissue damage, cornea, bone, aorta, tissue denaturation, optical properties of water
first approximation, the optical and thermal properties of tissue can be taken t o be those of liquid We report measurements of ablation thresh- water. The simplest models treat ablation as the olds and tissue damage vs. wavelength by using rapid (explosive) vaporization of water [ll, while the tunable, pulsed Co:MgF, laser. Tissue abla- more complex models consider the dynamics of tion with pulsed infrared (IR) lasers has been vaporization [161 and may include a layer of liqstudied by a number of workers using a variety of uefied material between the tissue being ablated laser sources ranging in wavelength from 2 to and the surrounding atmosphere [5,6]. The pre10.6 pm and in pulsewidth from less than 100 ns dictions of ablation models can be compared to to over 100 ps [l-151. Studies using pulsed and measurements of a number of parameters, includcontinuous-wave CO, lasers have shown that ing the ablation threshold, the mass-loss rate, and thermal damage can be reduced by using pulsed the momentum transferred to target tissue. lasers. Thermal damage zones as small as a few Thermal damage caused by denaturation of microns thick can be obtained at wavelengths tissue at the bottom and sides of the ablation crathat are strongly absorbed in tissue by using ter is of considerable clinical interest, since it can pulses that are short compared t o the thermal re- affect the speed of wound healing. The simplest laxation times of the laser-heated volume; for ex- thermal models of tissue damage calculate the ample, damage zones of 5-10 pm have been found depth to which tissue is heated to the denaturin the cornea following ablation using a Qswitched Er:YAG laser operating at 2.94 pm [41, the wavelength of the peak of the water absorp- Accepted for publication December 11, 1990. tion band. Address reprint requests to Thomas F. Deutsch, Wellman Modelling of pulsed IR laser ablation has Laboratories of Photomedicine, Department of Dermatology, generally proceeded on the assumption that, t o a Massachusetts General Hospital, Boston, MA 02114. INTRODUCTION
0 1991 Wiley-Liss, Inc.
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cient of liquid water, taken from data of Curcio
-rl and Petty 1191, together with the tuning range of
i -
C o Mgl', Range
i
CZ'avrlrngth ( / i m )
Fig. 1. Mid-infrared absorption spectrum of water showing the 1.94-)~m peak of interest (taken from ref. 19).The dashed line shows the range covered in these experiments.
the pulsed Co:MgF, laser. The absorption coefficient has a peak of 114 cm-' at 1.94 pm, and drops t o less than 10 cm-' at 1.85 pm and t o about 16 cm-' at 2.2 pm. Thus the optical penetration depth can be varied by over a decade by tuning the output wavelength of the Co:MgF, laser. The primary goal of these experiments was t o examine the effect of optical penetration depth on ablation threshold and thermal damage for a variety of tissues. In addition, since some of the experiments employed a high-speed framing camera in the determination of ablation threshold, information about the dynamics of material removal was also obtained. MATERIALS AND METHODS
ation temperature. This approach predicts damage only at the bottom of the crater, while in practice a layer of damaged tissue roughly uniform in thickness is observed at the walls and the bottom of the crater. More complex thermal models include a layer of liquefied material between the tissue and the surrounding atmosphere; the radial forces on the liquid layer due t o a nonuniform laser beam profile cause the liquid layer to be ejected from the crater, flowing along the crater walls. This liquid layer lining the crater acts as a thermal reservoir which can transfer heat t o the crater walls, leading t o damage [17]. Thus the existence of a liquid layer during the ablation process accounts for the geometry of the observed thermal damage. Most of the ablation studies to date have been performed with a laser operating at a single frequency; some studies using the line-tunable HF laser have also been reported [10,111. Comparisons between such studies are sometimes difficult because of variations in beam profile and pulse shape, as well as because of possible differences in methods of beam characterization. Measurements using a broadly tunable IR laser can eliminate some of these difficulties and provide additional data to improve our understanding of ablation, provided the absorption coefficient of water varies significantly, by at least a decade, over the laser's tuning range. The pulsed Co:MgF, laser, which can be tuned from 1.75 t o 2.5 pm, is well suited for studying the way in which the optical penetration depth affects tissue ablation and tissue damage 1181. Figure 1shows the optical absorption coeffi-
The Co:MgF, laser (Schwartz Electro-Optics, Concord, MA) was optically pumped by a Nd:YAG laser emitting at 1.338 pm. Typical output energies were 70 mJ in a train of 2-ps pulses approximately 100 ps long. The details of the pulse shape varied with wavelength and pump conditions. For example, at 1.91 pm the pulse train consisted of 5 pulses, each returning t o baseline, while at 2.0 pm several micropulses were superimposed on the envelope of a 100-ps macropulse. Such differences can be important in quantitative modelling of thermal damage, but were not documented in detail here. The spot size at the tissue was obtained from the beam profile, which was determined by scanning a pyroelectric detector (model 53, Molectron, Campbell, CA), with a 50pm pinhole in front of it, in the target plane. Readings were taken at 25-pm intervals and processed by computer t o obtain a three-dimensiona (3-D) plot of the beam profile, which was nearly Gaussian. As discussed below, a cross-section through this 3-D plot was used to obtain the Gaussian spot size of the beam. Two methods were used to obtain Fth, the threshold fluence for ablation. The first, shown schematically in Figure 2a, was an optical pumpprobe technique; a He-Ne laser beam, focused to a spot size of 60 pm, passed approximately 80 pm in front of and parallel to the surface of the sample and was detected by a Si photodiode. An x-y-z stage was used t o accurately position the sample surface with respect to the He-Ne probe beam. The onset of tissue ablation resulted in a reduction of the photodiode signal, due to either the deflection of the beam by refractive index
Co:MgF, Laser Ablation of Tissue
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Fig. 2. a: Experimental schematic of the pump-probe system for determining ablation threshold. b: Experimental schematic of the framing-camera system for determining ablation threshold. The He-Ne laser is used to position the tissue sample reproducibly.
changes in the air near the tissue or to a blocking of the beam by ejected tissue. The temporal profile of the Co:MgF, laser pulse was obtained using
a fast HgCdTe photoconductive detector. The output of both detectors was fed to a digital oscilloscope (model 9400, LeCroy Corp., Chestnut Ridge,
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NY); comparison of the two traces allowed one t o determine when during the laser pulse ablation started. The signal-processing capabilities of the digital oscilloscope were used t o measure the integrated energy delivered at the time ablation started. The Co:MgF, laser energy was focussed onto the tissue surface by a 20-cm-focal-length lens, resulting in a l/e2 spot diameter, 20, of 255 pm as determined by fitting a cross-sectional slice through the 3-D beam profile to a Gaussian function, f(r) = exp(-2r2/a2). Since, at threshold, ablation begins at the peak of the Gaussian beam, the peak fluence, which can be shown t o be twice the total energy in the pulse divided by m r 2 , is plotted in the figures below. The second technique for measuring ablation threshold, shown schematically in Figure 2b, involved the use of a framing camera (Imacon HE 700, Hadland Photonic, Cupertino, CA) t o obtain a series of pictures of the tissue target, viewed at 90" to the ablating laser beam axis, during the ablation process [81. The target was illuminated by a pulsed flash lamp having an output pulse several hundred microseconds long. The He-Ne laser and associated lenses were used t o position the sample reproducibly. The magnification of the camera was tenfold. The exposure time of a single frame was approximately 200 ns; a series of 10 or more photographs separated by intervals of 1 ps was taken and the onset of ablation determined from the photographs. That is, the time of the first photo showing material removal was determined. A delay circuit was used to vary the time delay between the onset of the laser pulse and the start of picture-taking. The output of the HgCdTe detector was sent into one channel of the digital oscilloscope and pulses from the camera, synchronized with the gating of the electronic shutter, were sent to the other channel. The Co:MgF, laser energy was focussed onto the tissue surface by a 30-cm-focal-length lens, resulting in a lie2 spot diameter of 368 pm, as determined by fitting the measured profile to a Gaussian function. The threshold energy was determined by integrating the laser pulse profile up t o the appropriate camera "sync" pulse. Ablation threshold measurements were made in vitro on samples of bovine calf cornea, chicken liver, human aorta, and rabbit bone. In addition, threshold measurements were also performed on collagen gels (gelatin) having 85% water content; such gels are often used as models for tissue. The ablation thresholds of fatty plaque and calcific plaque were also examined. All abla-
01 1.8
1.9
2.0
2.1
2.2
Wavelength (pm)
Fig. 3. The dependence of ablation threshold on wavelength for collagen gels containing 85% water. The data were recorded with the optical pump-probe technique only. The solid line represents a least-squares fit assuming that F,, = a h .
tion threshold measurements were made using a single laser pulse, thereby avoiding changes in the tissue, such as dessication, which can occur with multiple pulses. Thermal damage measurements were made on samples of calf cornea and rabbit ulna bones. Single-pulse ablations were performed at different wavelengths and fluences. The zone of thermal damage at the bottom of the ablation crater was determined by light microscopic examination. Cornea specimens were fixed in either formalin or Bouin's solution, routinely processed, and stained by using hematoxylin and eosin (H&E). Bone samples were fixed in 80% ethanol for 48 hours at 4"C, transferred t o a saturated aqueous solution of ethylenediaminetetraacetic acid (EDTA), and kept there until the sample was decalcified, approximately 3 weeks. Samples were then embedded in paraffin, sectioned, and stained with H&E. RESULTS
Figures 3-8 show the ablation thresholds vs. wavelength for gelatin, calf cornea, chicken liver, human aorta, diseased human aorta, and rabbit ulna bone respectively. The data cover a region in which the absorption coefficient of water, a,varies by over a decade, and include measurements on both sides of the 1.94 pm peak. As has been discussed in the literature [ll, Ftha is the energy per unit volume deposited at the surface at ablation threshold, Fth, and can be identified with a heat of ablation, Habl.This heat of ablation is en-
Co:MgF, Laser Ablation of Tissue
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Fig. 4. The dependence of ablation threshold on wavelength for resected calf corneas in situ. The data were recorded using the optical pump-probe technique (open circles) and framing camera technique (solid circles). The solid curve represents a least-squares fit to the optical pump-probe data assuming that Fth = a h .
ergy/volume that must be deposited at the surface in order t o cause a deflection of the surface (in the framing-camera measurements) or to expel some tissue fragments (in the pump-probe measurements). Provided the heat of ablation is wavelength independent and that scattering does not significantly change the fluence below the surface, Fth should vary as l/a.Both these assumptions will be considered in more detail below. The solid lines in Figures 3-7 and the dashed line in Figure 8 are plots of a/a, with “a”obtained from a least-squares fit to the data using published values of (Y [191. Figures 4 and 5 combine the results of the pump-probe and framing-camera techniques for obtaining the ablation threshold for calf cornea and chicken liver. In particular, the results using cornea show excellent agreement between the two measurements. For liver, there is a systematic increase in the difference between the two techniques with increasing wavelength in the region 1.95 to 2.15 pm. This may be explained by examination of framing-camera pictures (vide infra) for ablation with these wavelengths. In particular, as the wavelength increases beyond 1.95 pm, the time between the observation of a deflection of the tissue surface, considered threshold for framing camera measurements, and the expulsion of tissue fragments, increases. The determination of thresholds using the pump-probe technique is sensitive to this time delay, resulting in larger threshold values. The absorbed energy per unit volume decreases as the laser is tuned off the wa-
1.9
2.0
2.1
2.2
Wavelength ( p m ) Fig. 5 . The dependence of ablation threshold on wavelength for resected chicken livers. The data were recorded using the optical pump-probe technique (open circles) and framing camera technique (solid circles). The solid curve represents a least-squares fit to the framing camera data assuming that Fth = a h
h
1.8
1.9
2.0
2.1
2.2
Wavelength ( p m ) Fig. 6. The dependence of ablation threshold on wavelength for resected human aortas. The data were recorded with the optical pump-probe technique only. The solid line represents a least-squares fit assuming that Fth = a h .
ter absorption band, leading one to expect slower ablation dynamics. The threshold for bone, in which water is a minority constituent, shows a relatively weak variation with wavelength; the actual water content of bone has been reported to vary from 34% by weight for young bone to between 16.6 and 22% for established bone of varying age [ZOI. We expect that rabbit ulna is probably close to young bone in water content. The dashed curve in Figure 8 is proportional t o (Y-’ and shows that, as expected, the threshold data do not track the water absorption. The solid curve represents a leastsquares fit of the data by a simple model assum-
Schomacker et al.
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TABLE 1. Heats of Ablation for Gelatin and Various Tissue TvDes ____ ~
Tisssue type Gelatin Calf cornea Chicken liver Human aorta Fatty plaque Calcified plaque Rabbit bone (ulna) 1.8
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2.1
2.2
Wavelcngth ( p m )
a” (J/cm3) 2,901 865 562 738 1,002 1,205 4,200b
%water ( x 0.01) 0.85 0.80 0.79 0.79 ? ? 0.30
al%water (J/cm3) 3,410 1,080 712 932 -
14,000
“Found by fitting the data to F,, = a/a(X). bThe equation, Fth = 2,0401(0.3xa(h)+a’), where a is the absorption coefficient of water and a’ ( = 18 crn-l) is from an unknown chromophore, gives a better fit.
Fig. 7. The dependence of ablation threshold on wavelength for resected diseased human aortas. The data were recorded with the optical pump-probe technique only. The solid line represents a least-squares fit to the fatty plaque data assuming that Fth = a h .
fitting the Fth vs. wavelength curves for the materials discussed. As discussed above “a” may be identified with a heat of ablation for the initiation of material removal. The last column gives “a” 300 . _ . . corrected for the varying water content of the difFth(h) = 4200/a(h) ferent tissues and should approximate the heat of F,,(h) = 2040/(.3xa(h)+18) vaporization of water, 2500 J/cm3, if ablation is essentially explosive vaporization of the tissue 200 water. Note that the heat of ablation for normal aorta is lower than that of either fatty or calcific plaque, indicating a potentially undesirable selectivity for normal tissue in angioplasty applications. However, all values were determined from single-pulse measurements and the results obtained from multiple pulses used clinically may O C converge due t o the sample dessication [9]. 1.8 1.9 2.0 2.1 2.2 Figures 9 and 10 display thermal damage t o Wavelength (pm) calf cornea as a function of wavelength at a fluFig. 8. The dependence of ablation threshold on wavelength ence of 7.3 ? 2.6 times Fth and as a function of for the resected rabbit ulna bones. The data were recorded fluence at a wavelength of 1.94 pm, the peak of with the optical pump-probe technique only. The dashed line the water absorption curve. Damage zones as represents a least-squares fit assuming that F,, = a h . The small as 150 pm are obtained at 1.94 pm. Note solid curve represents a modified fit to the data assuming that that the use of Bouin’s fixative rather than fora second chromophore having a constant absorption coeffimalin led t o a systematically higher damage zone cient exists (see text). measurement. The thermal damage t o bone varied between 160 and 205 pm in the wavelength ing two different absorbers in bone. Specifically, region 1.81 t o 2.14 pm, with an average value of the model assumes that 30% of the laser energy is 180 15 pm; there was no significant dependence absorbed by water with the remaining energy ab- on wavelength. Figure l l a and b shows two series of framsorbed by a second chromophore, such as collagen or hydroxyapatites, which has a wavelength-in- ing-camera pictures of the ablation of cornea, dependent absorption coefficient in the wave- while Figure 12 shows a similar series for chicken length region studied. The fitting procedure gives liver. Time increases in 2-ps increments from left Fth = 2040/[0.3 X &(A) + 181 J/cm2. The first t o right and in 1 - p s increments from bottom to term, the water contribution, is about 34 cmpl at top; the beam comes from the right. One striking 1.94 p m , compared t o 18 cm-’ for the wave- difference in the pictures for ablation at threshold length-independent term. is that the ablation of the cornea shows the forTable 1 tabulates the value of “a” found by mation of a cavity, which finally breaks. By con-
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Co:MgF, Laser Ablation of Tissue DISCUSSION
147
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A primary goal of this study was to use the tunability of the Co:MgF, laser t o study the wavelength dependence of the ablation threshold and the thermal damage in order to determine how well simple thermal models based on the vaporization of water describe tissue ablation. For the case of gelatin, cornea, chicken liver, human aorta, fatty plaque and calcific plaque the ablation threshold tracks l/a reasonably well. No such relationship is observed for bone, which contains 17 I to 34% water, and none would be expected. This 19 20 21 result is in agreement with data presented in refWavelength (,urn) erence 11, which show negligible threshold deFig. 9. The dependence of thermal damage on wavelength for pendence .with wavelength- fir bone within the resected calf corneas. The measurements were taken from 2.7-2.9-pm absorption band of water. Note that tissue samples fixed in formalin (open triangles and open while both gelatin and cornea track l/a, the acsquares) and Bouin's solution (filled traingles). Except for the tual value of Hablis almost a factor of three higher open squares, all fluences were 7.3 2 2.6 times the ablation for gelatin, indicating that gelatin does not bethreshold found at each wavelength (see Fig. 4). The open squares used fluences of 2.6, 3.0, 12.8, and 36 times the ab- have exactly like cornea, despite the fact that lation thresholds found at 1.81, 1.85, 1.89, and 1.93 pm re- their collagen and water contents are similar. spectively. The solid line represents a least-squares fit to the The analysis of ablation used here has asopen and filled triangles, assuming that the damage depth sumed that 1) the effects of scattering are neglivaries as 01-l. gible at the wavelengths used and that 2) the heat of ablation does not vary with wavelength. The effect of scattering will vary from tissue t o tissue. Cornea, which is optically clear in the visible porE 3 500 400 tion of the spectra, would not be expected to exhibit significant scattering in the 2-pm region. The role of scattering in liver can be examined by 300 using recent data. The absorption and scattering coefficients for liver have been measured in the 200 range 350-2,200 nm [all. For situations where the absorption coefficient,l pa, is equal to or 100 greater than the scattering coefficient, ps, the propagation of light is described by the transport 0 coefficient, pt = pa ps(l-g), where g is the anisotropy parameter [22]. Scattering can be neglected if the second term is small compared to the first one. At 1.94 pm, pa = 73 cm-l and Fig. 10. The dependence of thermal damage found at 1.94 pm = 3.8 cm-'. Here the condition for nepJ1-g) on fluence for resected calf corneas. The line shown is a linear regression analysis of the data showing a small (0.17 km- glecting scattering is well satisfied. However at cm2/J) increase in thermal damage with increasing fluence. 1.8 pm, pa = 11 cm-' and pJ1-g) = 3.8 cmP1 The dotted line denotes the averaged thermal damage. and the condition is only weakly satisfied. The values of ps obtained from reference 21 should be used with caution, since diffusion theory, used to trast, the pictures of chicken liver ablation show analyze the data, is not appropriate near 1.94 pm what may be multiple bubbles forming. The picwhere absorption is greater than scattering. One tures do not allow us t o decide if the cavities are consequence of scattering is t o increase the subempty, full of vaporized water, or full of liquid (water or liquefield collagen). Both figures illustrate how the onset of surface disturbances, pre- 'pa is the same as a used earlier in the text. The change of sumably the onset of ablation, can be determined notation here is to agree with the notation in references 21 and 22. t o within about 1 ps.
,-.
v
5
+
148
Schomacker et al.
Fig. 11. Two series of framing-camera photographs of the ablation of cornea taken early (a)and late (b) in the ablation process. The wavelength was 1.91 pm and the fluence 71.2 J/cm2. The exposure time for each photograph was 200 ns.
The bottom left frame corresponds to the earliest time; the laser beam enters from the right. Time increases from left to right and from bottom to top with the odd frames depicted in the bottom row.
surface irradiance in the tissue above the incident in the wavelength region of interest have not irradiance, an effect that would lead t o an under- been measured. estimation of the true ablation threshold. For the to the ablation The fits of Fth = H,,/a case of bone and plaque the scattering coefficients threshold vs. wavelength curves resulted in val-
Co:MgF, Laser Ablation of Tissue
149
Fig. 12. A series of framing-camera photographs of the ablation of chicken liver. The wavelength was 1.95 km and the fluence 61.4 Jicm'. The exposure time for each photograph was 200 ns. The bottom left frame corresponds to the earliest time; the laser beam enters from the right. Time increases from left t o right and from bottom to top with the odd frames depicted in the bottom row.
ues of Hab, on the order of 1 kJ/cc, considerably less than the 2.5 kJ/cc required t o vaporize water. This result is similar to that found in studies of the ablation of skin using pulsed CO, TEA lasers; Ftha obtained from a linear fit t o mass loss vs. fluence curves was 0.57 kJ/cc, compared 2.37 kJ/ cc from a Beer's law fit [l].The reason for the difference is that the linear fit approach measures the lowest fluence at which some material is removed. The Beer's law fit to mass-loss data results in a higher threshold value which represents the threshold for explosive, "blow-off" material removal. Our measurements of threshold using either the pump-probe or framingcamera approach detect the beginning of ablation and so correspond to the threshold determined by using a linear fit to mass loss vs. fluence data. The fact that the Habl value for liver is the lowest of all the tissues studies is consistent with studies of the effect of mechanical properties on pulsed CO, laser ablation of tissue 1151. These studies found a heat of ablation for liver, calculated from the slope of a mass-loss curve, of 606 J/gm. This heat of ablation was the lowest of the tissues studied and was correlated with a very low tensile strength compared t o these other tissues. The cornea was used as a model system for thermal damage to tissue because its homoge-
neous nature eliminated some of the variability found in studies of damage t o skin. The damage is less than 100 pm at 1.92 pm and increases as the wavelength is tuned off the absorption peak t o 400 pm at 2.1 pm. It is important t o note that these damage measurements are for single-pulse exposures; multiple pulses have been shown to result in increased damage [6,171. The simplest model of damage, which assumes that damage occurs when the initial temperature distribution in the tissue reaches the denaturation temperature and calculates that depth by using the heat of ablation deposited at the surface, predicts a damage depth given by A/cx 121. In reference 2 it is shown that A = In [Fthcx/(T,-T0)pc], where To is the initial tissue temperature, T, is the critical temperature for denaturation, p is the density, and c is the specific heat. Using the heat of ablation for cornea given in Table 1, T, = 65"C, To = 20°C, and taking the absorption coefficient of cornea t o be 80%of that of water, one finds that A = 1.9. The solid line in Figure 9, given by l/a,is a fit to the data points considered most reliable, given by the filled and open traingles. This fit corresponds to A = 0.8 when the value of cx for cornea is taken to be 0.8 that of water. The measured damage zone shows a weaker wavelength dependence than predicted by the model. Note that the
Schomacker et al. damage at wavelengths shorter than the 1.94-pm tological sections. We thank Peter Moulton and absorption peak does not increase as rapidly as David Welford for many helpful technical discusthe l/a curve would require. Since the cornea ini- sions about the Co:MgF, laser and David Rines tially has negligible scattering, it seems unlikely for technical support. This work was supported by that the simple model fails because it neglects a subcontract from Schwartz Electro-Optics under scattering, although collagen denaturation dur- a National Institutes of Health SBIR program ing the ablating pulse could introduce a time-de- and by the SDIO-MFEL program under Office of pendent scattering. Some evidence for such time- Naval Research Contract N00014-86-K-0017. dependent changes in the optical properties of tissue does exist; the reflectance of liver has been shown to change during ablation with a pulsed REFERENCES 2.1-pm holmium laser 1231. The slight increase in 1. Walsh JT, Deutsch TF. Pulsed CO, laser tissue ablation: Measurement of the ablation rate. Lasers Surg Med 1988; damage with increasing fluence shown in Figure 8:264-275. 10 is not predicted by the model of reference 2. 2. Walsh JT, Flotte TJ, Anderson RR, Deutsch TF. Pulsed More extensive measurements are needed t o obCO, laser tissue ablation: Effect of tissue type and pulse tain sufficient data t o test a more complex abladuration on thermal damage. Laser Surg Med 1988; 8: tion model; however, the usefulness of the tunable 108-118. Co:MgF, laser in obtaining data to test models of 3. Walsh JT, Deutsch TF. Er:YAG laser ablation of tissue: Measurement of ablation rates. Lasers Surg Med 1989; thermal damage is clear. Clinically, the ability to 9:327-337. vary the damage by tuning the wavelength may 4. Walsh JT, Flotte TJ, Deutsch TF. Er:YAG laser ablation be useful, allowing the damage zone t o be adof tissue: Effect of pulse duration and tissue type on therjusted t o the minimum required to obtain hemomal damage. Lasers Surg Med 1989; 9:314-326. 5. Zweig AD, Weber HP. Mechanical and thermal paramestasis. ters in pulsed laser cutting of tissue. IEEE J Quant ElecThe framing-camera pictures demonstrate tron 1987; QE-23(10):1787-1793. the role of tissue mechanical properties in abla6. Frenz M, Romano V, Zweig AD, Weber HP, Chapliev NI, tion. Liver, with its low tensile strength, is clearly Silenok AV. Instabilities in laser cutting of soft media. J removed explosively. In cornea the strength of Appl Phys 1989; 66:4496-4503. 7. Nishioka NS, Domankevitz Y, Flotte TJ, Anderson RR. the surface membrane, Bowman’s layer, appears Ablation of rat liver, stomach and colon with a pulsed to be sufficient t o delay explosive removal until holmium laser. Gastroenterology 1989; 96:831-837. it is ruptured by the expansion of underlying 8. Domankevitz Y, Nishioka NS. Measurement of laser absteam. lation threshold with a high-speed framing camera. IEEE In summary, the use of framing-camera and J Quantum Elec 1990; In press. pump-probe techniques t o determine tissue abla- 9. Nishioka NS, Domankevitz Y. Comparison of tissue ablation with pulsed holmium and thulium lasers. IEEE J tion thresholds for wavelengths near the 1.94-pm Quantum Elec 1990; In press. water absorption band where the absorbance var10. Izatt JA, Albagli D, Itzkan I, Feld MS. Pulsed laser abies tenfold has been investigated. The results lation of calcified tissue: Physical mechanisms and funshow that the ablation thresholds for soft tissue damental processes. In: Jacques SL, ed. “Laser-Tissue Intrack l/a reasonably well. The equivalence in teraction.” Proc SPIE 1202, 1990; pp 133-140. thresholds obtained from these two techniques 11. Izatt JA, Sankey ND, Partovi F, Fitzmaurice M, Rava RP, Itzkan I, Feld MS. Ablation of calcified biological should be stressed because of the simplicity of the tissue using pulsed hydrogen fluoride laser radiation. pump-probe experiment. The effects of waveIEEE J Quantum Elec 1990; In press. length on thermal damage have also been inves- 12. Nelson JS, Yow L, Liaw LH, MacLeay L, Zavar RB, Orentigated for cornea and show only a weak depenstein A, Wright WH, Andrews JJ, Berns MW. Ablation of bone and methacrylate by a prototype mid-infrared erdence on l/a; for bone the damage zone is bium:YAG laser. Lasers Surg Med 1988; 8:494-500. wavelength independent. While simple thermal 13. Nuss RC, Fabian RL, Sarkar R, Puliafito CA. Infrared models predict the dependence of ablation threshlaser bone ablation. Lasers Surg Med 1988; 8:381-391. old on absorption coefficient reasonably well, an 14. Stern D, Puliafito CA, Dobi ET, Reidy WT. Infrared laser understanding of thermal damage will require surgery of the cornea: Studies with a Raman-shifted Nd: YAG laser at 2.80 and 2.92 pm. Ophthalmology 1988; further experimentation and modelling. 150
ACKNOWLEDGMENTS
We gratefully acknowledge the technical support of Margo Goetschkes in preparing the his-
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