LASER THERMAL ABLATION

Phoiochembtry and Photobiology Vol. 53, No. 6, pp. 815-823, 1991 Printed in Great Britain. All rights reserved

0031-8655/91 S03.OO t 0.00 Copyright @ 1991 Pergarnon Press plc

REVIEW ARTICLE

LASER THERMAL ABLATION A. J. WELCH'*, MASSOUD MOTAMEDI~, SOHI RASTEGAR', GERALD L. LECARPENTIER' and

Duco JANSEN* IBiomedical Engineering Program, The University of Texas at Austin, Austin, Tx 78712-1084, 2Department of Internal Medicine, Wayne State University, Detroit, MI and 3BioEngineering

Department, Texas A&M University, College Station, TX, USA

(Received 20 August 1990; accepted 15 Octof~er 1990)

Abstract4ontinuous wave and pulsed laser ablation of tissue is described as an explosive event. A subsurface temperature maximum and superheated tissue produce high pressures that eject fragments from the tissue. Decreased water content due to dehydration and vaporization decreases thermal conductivity which reduces heat conduction. Also, a decrease in water content dramatically alters the local rate of heat generation of laser radiation above 1.3 pm since water is the primary absorber, In contrast, at UV wavelengths protein and DNA are the primary absorbers so destruction of tissue bonds is due to direct absorption of the laser light rather than heat transfer from water.

NOMENCLATURE

A = area[cm2] c = heat capacity [J/g "C] H = heat of vaporization [Jig]

HI = heat of ablation [J/g] k = thermal conductivity [W/cm "C] t = characteristic length [crn] Q = radiant energy [J] S = laser heat source (W/cm3] T = temperature ["C] r, = pulse duration [s]

W,, = threshold radiant energy density Z, = ablation front position at time I 2 = ablation interface a = thermal diffusivity [cm2/s] 6 = optical depth [cm] p, = absorption coefficient [cm-'1 CL, = scattering coefficient [cm-*I p., = total attenuation coefficient [cm-)] p = density [glcm'] 7 = diffusion time IS]

7, = time constant for radiant heat flow [s] T, = time constant for axial heat flow [s] 4" = irradiance [W/cm*] $I,,, = threshold fluence [J/cm2] O, = l/e2 radius of gaussian beam [cm].

In the text, some results are presented in terms of mm. Units different from tables are clearly marked.

INTRODUCTION

Reponses of tissue to laser irradiation are typically classified as photochemical, photomechanical, photothermal and photoablative. Photochemical reactions are associated with low fluence rates that do not produce a significant temperature increase in the irradiated tissue, but do interact with a natural or exogenous photosensitizer to produce the desired

reaction. An example of medical importance is the treatment of cancer by injection of hematoporphy- rin derivative and irradiation with a red light. In contrast, photomechanical responses occur during application of extremely high fluence rates (greater than 1On Wlcm2), which produce shockwaves and plasmas. Shockwaves generated with exposure durations of 1.0 ps and less, give rise to the photomechanical reaction that fragments kidney stones. At short wavelengths (less than 200 nm), photon energy is sufficient to break molecular bands and lead to photoablation. Even though there are a number of important photochemical and photomechanical applications for lasers, a majority of the current medical procedures involve photothermal reactions. Desired end points are either coagulation or, at higher fluence rates, ablation.

Although there are many examples for the ablation of biological media using continuous wave (CW) lasers such as argon, Nd:YAG and C02 lasers, there has been a growing interest in the use of short duration (less than a few ms) pulses to ablate tissue. The rationale for favoring pulsed to continuous irradiation are (1) the desire to produce a thermal event that is shorter than the thermal relaxation time of the tissue to reduce the zone of thermal damage; (2) the delivery of sufficient energy to ablate tissue with each pulse which removes the hot tissue before heat is transferred to surrounding tissue; (3) precise tissue cutting control; and (4) ablation of dense tissue such as bone or tooth by rapid heating andlor plasma formation.

The sequence of events that leads to laser ablation of tissue may vary depending on the optical and mechanical properties of tissue as well as the tem- poral and spatial characteristics of the laser beam. For instance, the ablation of aortic specimens using *To whom correspondence should be addressed.

815

816 A. J. WELCH et a!.

a CW argon laser is often initiated by the occurrence of an explosion resulting in tissue tearing, and is followed by the process of tissue pyrolysis. In contrast, when a pulsed excimer laser is used, rapid and clean ablation of tissue is obtained. In this paper we discuss a number of factors that influence the ablation process. By examining existing exper- imental data, it is possible to delineate mechanisms that are associated with tissue ablation. Here the term ablation simply means the removal of tissue. This may be by vaporization of the liquid and solid components by pyrolysis of macromolecules and/or spallation where tissue fragments are ejected during the ablation process. In this article, we will first address issues related to CW laser ablation of tissue. This will be followed by a section on pulsed laser ablation which includes issues related to delivery of high power pulsed laser via optical fibers.

VAPORIZATION OF WATER

Since tissue is 75430% water, it is tempting to describe ablation in terms of water vaporization. If water is at some temperature Ti, vaporization requires sufficient heat to bring the water to a critical temperature T, (usually 100°C for water) plus the heat of vaporization necessary for the phase change. That is, the threshold radiant energy density w [ h is

For water at 37”C, 0.25 J/mm3 are necessary to heat the water to 100°C and pH = 2.25 J/mrn3. So approx. 2.5 J/mm3 are required to vaporize water. A generalization of the heat [J/mm3] required to vaporize water can be formulated by assuming that the laser beam has a uniform irradiance profile, &, [W/rnm2) and pulse duration t0 [s ] with area A [mm2], and the laser energy is uniformly distributed to an optical penetration depth S [mm] which is also assumed to be the ablation depth. With these assumptions the threshold fluence +th [J/rnm2] for vaporization is

4 t h = Wth8 = @do = 2.5 8 [J/mrn2] (2) and the energy that must be supplied by the laser radiation is

QL.=~ = J’td [JI (3) Although tissue is not water, the 2.5 J/mm3 is a

reasonable benchmark. However, assuming that the ablation depth is equal to the optical penetration depth is a generalization for which there are many counter-examples.

One counter-example is continuous wave ablation of low irradiances with a wavelength such as 1060 nm that penetrates several mm into tissue. Typically a large portion of the heat generated by the absorption of the laser light is conducted axially

and radially away from the ablation front. The diffusion of heat into surrounding tissue can be minim- ized by increasing the irradiance until the velocity of the ablation front is equal to the diffusion time of the tissue. In practice, irradiation times are kept below 100 ms to reduce diffusion losses and unnecessary thermal damage of surrounding tissue. Another example is structural strength of tissue. Walsh (1988) has hypothesized and has demonstrated that at high irradiances mass loss is due to laser-induced explosive removal of tissue , and that weaker tissues ablate more rapidly than stronger tissues. Walsh has irradiated liver, myocardium, aorta, and skin with a TEA C 0 2 laser and has plotted “mass loss per pulse” vs “fluence per pulse.” At fluences above 5 J/cmZ the largest mass loss is from liver, followed by myocardium, aorta and skin. A significantly deeper ablation depth is recorded for the “mechanically weaker” liver than the other tissues. In this paper we explore the various thermal events that are associated with CW and pulsed ablation. The importance of wavelength and pulse structure is examined in relation to these thermal events.

CONTINUOUS WAVE ABLATION

It is not difficult to postulate that ablation requires sufficient heat to produce phase changes of liquid and solid components. This heat is supplied by absorption of laser Light in the tissue. Basic thermodynamic principles for 1-D laser ablation have been suggested by Rastegar et al. (1989). Their theoretical ablation model uses the heat conduction equation which has the form

a~ a aT at az az pc- = - k - + S, (4)

a critical ablation temperature T, and after the onset of ablation an interface of energy balance at the ablation interface

(5 ) kaT(z,r) dz -= pH,-;i; a t t = Z az

where k [W/cm°C] is thermal conductivity, S [W/cm3] is the laser heat source, H , [J/g] is the heat of ablation, and Z is the ablation front position at time t . The heat of ablation H , is the sum of the latent heat of vaporization of tissue water content and the energy necessary to initiate pyrolysis of tissue components.

According to Eq. (5 ) a positive ablation velocity requires a positive temperature gradient. That is, the temperature just inside the tissue must be higher than the temperature at the irradiated surface. Sim- ultaneous solution of Eqs. (4) and ( 5 ) produces the temperature response shown in Fig. 1. Initially, maximum temperatures are at the surface. Once ablation starts, the maximum temperature shifts inside the tissue and a constant ablation velocity is

Review Article 817

0 I000 2000 3000 4oao 5Mw) boo0 Position (nucronr)

0 IS00 3000 4sw boo0 7500 woo Poriuon (mcmnr) Figure 1 . Computed I-D temperatures as a function of

time using Eqs. (4) and ( 5 ) shown in non-dimensional units. Subsurface tmperature peak and motion Of the computed ablation front is observed after the onset of

ablation.

Figure 2. Axial temperature protile during ablation of (a) a polyacrylamide rod, and (b) collagen fibers with a 60 W/cm* CW argon laser irradiation, In each graph, the origin represents the initial front surface position of the sample (LeC’arpentier el a / . . 1990). The * indicates the

location of the ablation front at time of measurement. Dredicted. Once this hypothesis was postulated, it -. kas possible to design experiments to determine the validity of a temperature maximum inside the tissue or of a constant ablation velocity.

Rastegar et al. (1986) measured ablation velocity and temperatures of thin agar gel rods during CW laser ablation. They demonstrated a peak temperature just distal to the ablation front and a constant ablation velocity once ablation was initiated. Le- Carpentier et al. (1990) and Verdaasdonk et al. (1990) independently conducted similar studies to chronicle the response of tissue during CW ablation. LeCarpentier et al. (1990) recorded 1-D axial surface temperature profiles of thin rods (lateral view) and 2-D temperature profiles of flat tissue samples (top view) with a thermal camera while simul- taneously recording visual changes with a high frame rate video camera.

Figure 2 displays the lateral temperature profile along a thin (< 2mm) rod of polyacrylamide gel that is being irradiated and ablated coaxially by an almost uniform CW argon laser beam. During the ablation process, the peak temperature is distal to the irradiated front surface. Also shown in the figure is the temperature along the side of a collagen fiber bundle during argon faser ablation.

Although ablation with pulsed lasers is growing in prominence, the events associated with CW ablation provide clues for understanding the mechanisms of pulsed ablation. By using a moderate irradiance from an argon laser and measuring surface temperature with an IR thermal camera, it is possible to document the initial surface heating and explosive ejection of tissue (popcorn effect). This process is illustrated in Fig. 3 for argon laser irradiation of aorta. The rather low irradiance used in this experiment permits a clear visualization of the thermal events. The onset of ablation occurs above 100°C due to subsurface superheating. Ablation is a non-

equilibrium thermodynamic event where fluid within the tissue matrix is converted into steam at a pressure of over 1 atm. As hot tissue is removed, a cooler layer is exposed to the laser irradiation. The temperature of the exposed layer remains at approx. 100°C while continued irradiation dehydrates the tissue. The removal of water decreases the local thermal conductivity reducing heat conduction to the surrounding area. Continued radiation rapidly increases the tissue temperature until it reaches about 300°C; the tissue burns and car- bonizes as nucleation sites form on the surface. At this point a rather constant ablation velocity is established.

The following thermal events illustrated in Fig. 3 are typical for CW ablation. These events have been observed in our laboratories (LeCarpentier ef af., 1990; Welch el al . . 1987) and elsewhere (Verdaasdonk et al., 1990).

(1) Initial temperature rise due to absorption of laser irradiation.

(2) With continued irradiation, tissue temperature is driven over 100°C. Tissue fluid is

425 .f

- . . . 0 i Ib IS 20 u

Time (s)

Figure 3. Surface temperatures of porcine aorta samples at center o f X W argon irradiation. The beam diameter is

1.8 mm (LeCarpentier er d., 1990).

818 A. J. WELCH et al.

superheated and there is a rapid subsurface pressure increase. The ensuing explosion, often called the “popcorn” effect, ejects tissue; tears and dissections are seen in the crater. Removal of heated tissue exposes a cooler layer of tissue. Continued irradiation dehydrates the tissue while surface temperature remains at 100°C. As the tissue dehydrates, the thermal conductivity, k, and density, p, decrease (Spells, 1960). Diffusion time is increased and there is a rapid increase in temperature once the layer is dehydrated. Temperatures increase until burning and charring occur at nucleation sites. Sites of the irradiated surface burn and expose cooler tissue producing a variation in temperature from 350 to 450°C. Heat not directly associated with ablation is conducted to cooler regions. The extent of thermal denaturation surrounding the crater is due to the temperature rise associated with the conducted heat.

During the process there may be changes in optical properties. Denaturation may produce a whitening of the tissue (increased scattering) whereas carbonization dramatically increases absorption in the 400 nm to 2.0 pm band of wavelengths.

All of these events are evident when the tissue is ablated with low level CW irradiation. However, when ablation is accomplished with pulses less than a few ms, some of these events are difficult to identify. If the pulse repetition rate is high and there is not enough energy in a single pulse to produce ablation, then the process has the appearance of CW ablation as temperature responses for each pulse become superimposed as was shown by Deckelbaum et al. (1985).

PENETRATION DEPTH

The ablation process is strongly influenced by the wavelength dependent penetration depth of the laser beam and its fluence. The penetration depth at any wavelength is governed by the chromophores that absorb the laser light. In the UV region the important absorbers are proteins and amino acids whereas water, OH and amines are the dominant chromophores for IR wavelengths above 1.3 pm. In between, chromophores such as carotenoid and melanin pigments and hemoglobin govern the absorption process. At IR wavelengths where water is believed to be the dominant chromophore, the concentration of water may significantly affect absorption and the penetration depth of the laser

PENETRATION DEPTH ()I, =l/altenuation

coefficient) ArF KrF XeCI Dyr Argon Nd:YAG Ho:YAQ EvYAG CO2 193 248 308 465 514.5 1084 2100 2940 10600 -.

<I pm I <I pm I 15-20 pm 15-20 pm

30-50 pm

185-285 pm 330 um

185-285 pm I 330 um I 300.400 pm

I I I

Swrra l mm*

Figure 4. Measured penetration depths for various wavelengths in aorta. Depth at which the total fluence rate [W/cmZ] is reduced to e-I of the surface irradiance [W/cm2] value. When light is highly scattered, the measured l/e penetration depth is greater than the optical penetration depth which is defined as Upl (Furzikov ei al., 1987; Cross ei al., 1990). ‘Measured l/e penetration depth which is

much larger than p1 when p,$ >> p,,.

beam. Typical penetration depths for several of the lasers considered for medical applications such as laser angioplasty are presented in Fig. 4. It is important to realize that for every penetration depth associated with a UV wavelength, there is at least one IR wavelength that will give exactly the same penetration depth although the chromophores are different.

ULTRAVIOLET PULSED ABLATION

There has been considerable interest in the ArF excimer lasers for corneal reshaping (A = 193 nm) as described by Cotliar et al. (1985) and the XeCl and XeF excimer lasers at 308 and 251 nm respectively for angioplasty as described by Singleton et al. (1987). The 0.5-1.0 pm penetration depth at 193 nm produces a sharp and extraordinarily fine cut with less than 1 pm of thermal damage (Marshall et al., 1985). At this wavelength there is sufficient photon energy (6.2 eV) that molecular bond breaking is expected and ablation is not totally due to thermal processes. Even though photoablation is not termed a thermal process, temperatures of several thousand degrees may occur in tissue vaporized by the nanosecond pulses from a 193 nm excimer laser (Srinivasan, 1986). Heat is carried away by the vaporized material before the residual tissue is heated. Although this wavelength has been very successful for removal of corneal material (Cotliar et al., 1985), its general use is limited by (1) problems of fiber optic delivery, and (2) mutagenicity of 193 nm radiation and fluor- escence that includes wavelengths in the mutagenic

Review Article 819

+ aWnm 0 11- . amnm 0 h .runrn.n!s - . Iolor*ll I o n

VICROV FLUEWCC IlnYmmq

Figure 5 . Ablation as a function of wavelength and fluence. Number of pulses to perforate 2.5 % 0.5 mm thick human cadaver atherosclerotic segments in air using a beam spot of 1.0 mm' area (Litvack et al . , 1988, with

permission. )

band from 200 to 300 nm. To avoid these problems. excimer laser angioplas-

ty systems have been developed around the 308 nm XeCl laser. Lengthening the wavelength decreases the absorption coefficient and increases the penetration depth. One consequence is an increase in the fluence per pulse to ablate the tissue. However, the increased penetration depth and higher fluence removes more tissue per pulse at 308 nm than at 193 nm. An excellent summary by Litvack el a/. (1988) of the effect of wavelength and fluence per pulse on the removal of tissue is illustrated in Fig. 5. Notice the extreme difference in fluence required to perforate a 2.5 mm thick sample of human cadaver atherosclerotic samples at 193 and 1060 nm; the penetration depths for these two wavelengths are 1.0 pm and a few mm respectively. Although satis- factory angioplasty results have been reported using 308 nm irradiation, two problems that are common to most wavelengths are: (1) difficulty in removing calcified plaque, and (2) dissections and tears on the wall of the crater (Bonner ef a l . , 1990).

MID-INZ'WARED PULSED ABLATION

Because of the high cost of excimer lasers and the possibility of achieving similar absorption properties in the IR, a concentrated effort has been devoted to the development of solid state lasers. Two lasers receiving considerable attention are the Ho:YAG at 2.1 pm and the Er:YAG at 2.94 Fm. The penetration depth at these wavelengths in tissue, based on the absorption properties of water and assuming tissue has a water content of 75%, is about 300 and 1.0 pm. respectively. Both lasers can be operated in either a Q-switched mode (standard == 2OO ns pulse) or normal pulse mode (-200 p s pulse). The normal pulse mode is made up of a series of micropulses with individual pulse durations of about a few microseconds. The Ho:YAG wavelength at 2.1 pm can be transmitted by standard low OH silica fibers for endoscopic applications; in

progress to determine the feasibility of this wavelength for angioplasty. I n vifro tissue studies have shown that the 2.1 prn wavelength efficiently ablates plaque; however, Bonner ef al. (1990) noted considerable tearing and dissections. Typically. pops are heard each time a pulse removes tissue. As with CW irradiation, ablation is associated with a subsurface explosive event that tears the tissue. An example of the surface temperature response and corresponding histology for €Io:YAG ablation of normal aorta is shown in Fig. 6.

The temperatures were measured using an IR camera aimed at the tissue surface during pulsed Ho:YAG irradiation in air. The surface temperature response at the spot center and at different distances away from the spot center are shown. Peak temperatures during the 200 pm irradiation are unknown and not shown in Fig. 6. Temperatures wcre measured at the video frame rate of 30 images per second.

The effectiveness of photothermal ablation during microsecond exposures depends on deposition of

(a) Ho-VAG irradlrlion of human aorta

3 Hz, 500 mJ/p

0 I 2 3

lime (sl

Figure 6 . (a) Temperature response of hunian aorta irradiated with a pulsed Ho:YAG laser in air. Plotted iire the surfacc temperatures iit the spot center and 2.2 and 5.3 mm away from the spot center, respectively. 'Thc spot diameter is 0.4 mm. Peak temperatures during irradiation are unknown and not shown in the ligurc. (h) Corresponding histological samplc ' Ihc i n t i m l side o f this sample of human aorta was irriidiated in ;iir with a pulsed Ho:YACi laser using the laser par;inictcrs as

fact, animai and human peripheral studies are in mentioned in Fig h(,t)

820 A. J. WELCH el al.

heat within the tissue: rapid enough to minimize heat losses due to conduction, yet slow enough to prevent formation of shockwaves due to the fact that vaporization cannot consume all the energy deposited at the surface.

As an example for these conditions, we have compared the response of tissue to Nd:YAG and Ho:YAG long pulse lasers (= 200 ps) by measuring crater depth induced in fresh porcine aorta as target, and delivering laser pulses via a 300 pm optical fiber to a spot size of 1 mm. Each experiment was repeated three times, and the mean value for crater depth measured microscopically is shown in Fig. 7.

Since Nd:YAG laser radiation is very poorly absorbed by vascular tissue (pa = O.l/cm), the energy supplied by each long pulse is subthreshold for tissue vaporization. Thus cumulative effects of individually subthreshold Nd:YAG laser pulses at relatively higher repetition rate (10Hz) are needed to initiate tissue ablation after a relatively long period of surface heating. In contrast, Ho:YAG laser radiation is more strongly absorbed by tissue (p, 25/cm) and heat is deposited at a faster rate due to the higher absorption. Tissue ablation is initiated much faster than when the Nd:YAG laser is employed, although the same energy per pulse is applied. Histologically, the extent of thermal injury in tissue irradiated with the pulsed Nd:YAG laser is significantly more than that observed in tissue exposed to Ho:YAG laser radiation. Yet, excessive mechanical injury and dissections surround the crater produced with a Ho:YAG laser.

A laser waiting for a fiber delivery system is the 2.94 pm Er:YAG laser. Two possible fibers are the zirconium fluoride and crystalline IR fibers. The respective transmission bands for these fibers are 0.3-4.5 pm and 3-15 pm. To date these fibers have been too large, brittle, and do not deliver sufficient power for most endoscopic ablation applications.

Yet initial in vifro ablation results using direct

I2

1 0 ]

EWWW n m ( u o ~ n d ~ )

Figure 7. Comparison of crater depth induced in porcine aorta during application of pulsed Ho:YAG and Nd:YAG

lasers at constant fluence level.

zoo aoo 400 coo so0 aioo a700 aaw aew JOOO

Wavelength (nm)

Figure 8. Mass removal of beef shank bone in air per joule as a function of wavelength. Assuming tissue is water the ablation yield would be approx. 0.4 mm3/JN. (Izatt el a/.,

1990, with permission).

irradiation have been encouraging. During each 200 pm pulse there appears to be a sequence of ablations. Walsh and Deutsch (1989) have reported ablation depths in bovine aorta of 700 pm for a fluence of 400 J/cm2 per pulse and a spot diameter of 1.1 mm. Walsh (1988) notes that each micropulse ejects a separate bolus of tissue from the surface, and ablation depth depends upon the mechanical properties of the medium. For Q-switched Er:YAG irradiation the explosive ablation at 24.5/cm2 ejects fragments of tissue at speeds of approx. 1250 m / s .

Izatt ef af. (1990) postulate a two component ablation mechanism in which bone and calcified plaque consist of apatite salts deposited in a soft component matrix. Using a high enough irradiance, the soft component is heated fast enough to gener- ate superheated vapor. The rapidly expanding vapor exerts force on the salt granules through momentum transfer and viscous drag. Particles as big as a few microns in diameter are ejected with velocities up to 600 m / s or Mach 2. Their experiments on bovine bone at 2.7-2.9 p.m with a hydrogen fluoride, 350 ns pulse width laser demonstrating mass removals in excess of the amount predicted by the simple vaporization of water support this hypothesis. Their results are presented in Fig. 8. Thus lasers with wavelengths around the water absorption peak at 2.94 such as the Er:YAG appear to be ideal for tissue removal including bone and calcified plaque.

Yet, reports of thermal damage extending 50 pm or more beyond the ablated crater in in v i m experiments using human aorta samples are unexpected in view of the shallow penetration depth of this wavelength (Walsh, 1988; Walsh and Deutsch, 1989). One possible explanation is suggested in the following analysis of thermal ablation.

ANALYSIS OF THERMAL tNJURY

Ablation of tissue is often accompanied using a series of pulses. Reports of the depth or volume of tissue removal per pulse always represent an average value for a specified sequence of pulses. There are scant data on the pulse to pulse variations of

Review Article 82 1

the ablation process and no data exist on the dynamics of ablation during a single pulse.

As an example, we will examine the thermodynamic events that take place during a Er:YAG pulse laser irradiation. First consider the thermal response of tissue to a single, short duration pulse such as a single Q-switched pulse. A high pulse energy vapor- izes the tissue and materal is ejected for a period of microseconds. Striking examples of the Q-switched Er:YAG ablation of tissue and water have been photographed by Walsh (1988) and Jacques et al. (1990), respectively. The depth of the crater produced by the Q-switched pulse corresponds to the penetration depth of the 2.94 pm wavelength in water. Heat remaining in the tissue just beyond the crater boundary reaches a maximum temperature at the end of the laser irradiation and then decays. If we assume that heat is generated in a 1.0 pm layer that is not vaporized, it is possible to estimate the temperature decay by estimating a diffusion time. The classic diffusion time is

In laser irradiated tissue, e is the smaller of either the penetration depth or beam radius (Furzikov, 1987). For a shallow penetration depth such as the 1.0 p n of the Er:YAG laser, thermal conduction is in the direction of the laser beam. However, if the beam radius is small relative to the penetration depth, conduction is normal to the beam. The conduction directions are illustrated in Fig. 9. van Gemert and Welch (1989) have approximated the thermal response with a simple but restrictive time constant model. The time constant is also based on the penetration depth and beam radius and thus closely resembles the diffusion time. They define an overall time constant T as

1 1 1 - = - + - 7 7, 7,

(7)

Thermal diffusion time 18 a function 01 spot dlameter and/or penetration depth

r-L'/4a a I thermal dlffurivlly

t - 1.3 x Idcm'/a I1 - 0.1 cm, then r - 20 ms

?J* ----.

Figure 9. Dominant diffusion times for shallow penetration (disk on the left) and deep penetration (cylinder on the right) of laser light in tissue. Beam diameter is represented by diameters of cylinders. For shallow penetration L is equal to the penetration depth ( 1 1 ~ ~ ) . For deep penetration L is equal to the radius (l/e2) of the

laser beam.

Lam Inadlanco

10 nr a awltch puir

't

4 PJ time eonstant (Ponehatlon depth 8 - 1 . 0 ~

Zono of thomi damago t on the order of a fow

mkrona

Figure 10. Impulse temperature response for Q-switched Er:YAG irradiation of tissue.

Based on the diffusivity of water, the value of the time constant is approximately related to values of squared image radius and penetration depth in mm.

7, [s] 22 w4 where wo is in mm

3 1 (8) T~ [s] (E)* where is in mm

For an absorption coefficient of loo0 mm-' which approximates the Er:YAG wavelength in soft tissue, the axial time constant 7, is 9 ks. This thermal response to a low energy Q-switched pulse is depicted in Fig. 10. This temperature response to an impulse of energy will not change as long as the laser pulse duration is less than approx. 1/10 of the time constant.

Having established the temperature response for a single impulse of light, consider the thermal response of a 200 ps pulse that consists of a number (-20) of pulses of about 1.0 ps duration. Typically, the pulses are not evenly spaced as depicted in Fig. ll(a) and the average time between micropulses is about 10 ps. Thus on an average, there is some overlap in the temperature response to each micropulse. However, in the region where the time between pulses is less than 10 ps, there may be considerable superposition of the temperature responses. If the tissue dehydrates owing to repeated temperature insults, the 2.94 pm wavelength will penetrate deeper into the tissue. This increases the time constant and decreases the rate

t

Figure 11. Tissue temperature response during 200 ps macropulse from an Er:YAG laser.

822 A. J . WELCH el al.

4.0 - 67

E

& 3.0 E B 3 2.0- n

a

e v

L U

1.0.- n -

0 . 0 9 s

of temperature decay. As a result of the increased penetration depth, the temperature of Fig. l l(b) may occur as an increase in remnant temperature from either micropulse to micropulse and macropulse to macropulse as water is vaporized from the tissue. If dehydration reduces the 2.94 p,m absorption coefficient to 100 mm-’, then the time constant will increase to 900 ps. Thus a critical factor in the ablation process with an IR wavelength will be the dynamics of water transport. The lengthening of time constant can be seen in the temperature responses of aorta to the sequence of Ho:YAG irradiations in Fig. 6.

Thus an important difference between excimer and IR ablation is the chromophore that absorbs the laser light. For UV wavelengths the chromophores are protein and DNA whereas the absorbing chromophore for IR irradiation is water. Dehy- dration will not affect UV ablation but it may drasti- cally alter the penetration depth of IR wavelengths. As dehydration takes place, the penetration depth will depend upon the concentration of OH and amine chromophores in the tissue.

0 No pbamo obaemd 0

Plaama obsemd / ’ .. - - Unaor ragreadon

B c. 0 .

HIGH IRRADIANCE DELIVERY

A significant problem encountered in the delivery of the 308 nm, 10 ns radiation via an optical fiber was the destruction of the front surface of the fiber. The fluence [J/cm2] sufficient for angioplasty required a relatively large irradiance [W/cm’] that chipped the proximal surface of the fiber. This problem was lessened by stretching the laser pulse length to 300 ns which decreased the required irradiance by a factor of greater than 20. The threshold for front surface damage is approximately directly pro- portional to the square root of pulse length as shown from the results of Singleton et al. (1987) presented in Fig. 12. Another solution is to increase the surface area of the surface by expanding the diameter of the fiber in a tapering fashion at the front tip

Fluence per pulse (J/crn2)

Figure 13. The mass of guinea pig skin removed per unit area by a TEA CO, laser focused to a 1.6 x 2.0 mm spot on the tissue surface. The dotted line represents a linear relationship between the mass of tissue removed and the logarithm of the fluence (Walsh and Deutsch, 1988 with

permission).

also described by Singleton et al. (1987). For a fixed pulse energy, the level of irradiance on the fiber can be decreased by increasing the surface area. Tapered fibers have a larger surface area while maintaining a small core diameter that is needed for flexibility.

If the irradiance delivered to the target is above lo8 W/cm2, the strong electric field forms a plasma. The plasma can develop in a few ns and once it forms, the laser irradiation is absorbed in the plasma and does not penetrate the tissue. As a result, the amount of tissue ablated as a function of fluence per pulse plateaus. The phenomena occur at all wavelengths and the threshold for plasma formation is a function of the absorption properties of the material. An example from Walsh and Deutsch (1988) of tissue removed per pulse as a function of fluence provided by a pulsed CO2 laser is presented in Fig. 13. The effect of plasma formation is clearly shown in the figure.

10’ N

B 3

I! I

c 2 so*

a LI

a 0

X

a a 10

Laaer Pular Our~t lon - nr

Figure 12. Threshold irradiance for front surface damage of a fiber optic as a function of XeCl (308 nm) laser pulse duration (Singleton et al., 1987). The dashed curve represents a PZ relationship for damage threshold. (With

permission.)

CONCLUSION

Thermal ablation is a non-equilibrium thermodynamic process. Both CW and pulsed ablation are characterized by subsurface heating, fluids become superheated, pressure increases until an explosive event removes the tissue. Dehydration of tissue alters the absorption properties at IR wavelengths, increasing penetration depths and the zone of thermal damage. When penetration depths are several hundred microns, considerable dissection and tearing of tissue is noted.

Acknowledgements-This work was supported in part by the Free Electron Laser BiomedicaVMaterials Science Pro- gram: ONR Contract No. N-14-86-K-0875, and in part by the Albert and Clemmie Caster Foundation.

Review Article 823

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LASER THERMAL ABLATION