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High-Resolution Optical Coherence Tomography-Guided LaserAblation of Surgical Tissue1
Stephen A. Boppart, Ph.D.,*,†,2 Juergen Herrmann, Ph.D.,* Costas Pitris, M.S.E.E.,*,†Debra L. Stamper, Ph.D.,‡ Mark E. Brezinski, M.D., Ph.D.,§ and James G. Fujimoto, Ph.D.*
*Department of Electrical Engineering and Computer Science, Research Laboratory of Electronics, and †Harvard–MIT Division of HealthSciences and Technology, Massachusetts Institute of Technology, Cambridge, Massachusetts 02139; ‡Department of Biology,
King’s College, Wilkes-Barre, Pennsylvania 18711; and §Cardiac Unit, Harvard Medical School,
ournal of Surgical Research 82, 275–284 (1999)rticle ID jsre.1998.5555, available online at http://www.idealibrary.com on
Massachusetts General Hospital, Boston, Massachusetts 02114
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Background. Optical coherence tomography (OCT)s a compact high-speed imaging technology whichses infrared light to acquire cross-sectional images ofissue on the micrometer scale. Because OCT imagesre based on the optical backscattering properties ofissue, changes in tissue optical properties due to sur-ical laser ablation should be detectable using thisechnique. In this work, we examine the feasibility ofsing real-time OCT imaging to guide the placementnd observe the dynamics of surgical laser ablation invariety of tissue types.Materials and methods. More than 65 sites on five ex
ivo rat organ tissue types were imaged at eightrames per second before, during, and after laser ab-ation. Ablation was performed with a coincident con-inuous wave argon laser operating at 514-nm wave-ength and varying exposure powers and durations.ollowing imaging, tissue registration was achievedsing microinjections of dye followed by routine his-ologic processing to confirm the morphology of theblation site.Results. High-speed OCT imaging at eight frames per
1 This research is supported in part by the Medical Free Electronaser Program, Office of Naval Research Contract N000014-97-1-066; the National Institutes of Health, Contracts NIH-9-RO1-A75289-01, NIH-9-RO1-EY11289-13 (J.G.F.), NIH-RO1-AR44812-1 (M.E.B.), and NIH-1-R29-HL55686-01A1 (M.E.B.); the Whittakeroundation Contract 96-0205 (M.E.B.); Air Force Office of Scientificesearch Contract F49620-98-1-0139; and the Air Force Palacenight Program (S.A.B.).2 To whom correspondence should be addressed at Department of
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lectrical Engineering and Computer Science, Research Laboratoryf Electronics, Harvard–MIT Division of Health Sciences and Tech-ology, Massachusetts Institute of Technology, 77 Massachusettsve., Room 36-345, Cambridge, MA 02139. Fax: (617) 253-9611.-mail: [email protected].
275
September 17, 1998
econd permitted rapid tissue orientation and guidedblation in numerous organ specimens. Acquisitionates were fast enough to capture dynamic changes inptical backscatter which corresponded to thermalissue damage during laser ablation.Conclusions. The ability of high-resolution high-
peed OCT to guide laser ablation and image the dy-amic changes suggests a role in image-guided surgi-al procedures, such as the ablation of neoplasms.uture in vivo studies are necessary to demonstrateerformance intraoperatively. © 1999 Academic Press
Key Words: laser surgery; imaging; ablation; neopla-ia; tumor.
INTRODUCTION
The surgical laser has been effective in the surgicaluite because it provides a cutting and coagulation toolith controllable powers, a compact, flexible fiber-opticelivery system, and wavelength- and tissue-specificutting efficiencies [1, 2]. Lasers are becoming stan-ard instruments for the ablation of endometrial foci ofhe reproductive tract [3], interstitial laser ablation ofumors [4], and surgical ophthalmologic procedures [5,]. While laser use has been widespread, control andosimetry of the delivery of laser radiation have reliedargely on visual feedback of the surface ablation site.his is a limitation which, by not allowing full visual-
zation of the location and extent of subsurface tissueblation, has the potential to result in iatrogenic in-ury.
Advancements in high-resolution image-guidanceechnologies coupled with the ability to tightly focus aaser beam would permit precise ablation of tissue,uch as a neoplasm, while minimizing collateral injury.
0022-4804/99 $30.00Copyright © 1999 by Academic Press
All rights of reproduction in any form reserved.
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urrently, however, no intraoperative imaging tech-ique exists with sufficient resolution to guide thelacement and fast acquisition rates to image the dy-amics of surgical laser ablation. Magnetic resonance
maging and computed tomography can image wholeody and organs, but have resolutions between 500 mmnd 1 mm, insufficient to resolve tissue microstructure.n addition, their use for high-speed intraoperativemage guidance is limited by size, cost, and complexity7]. Ultrasound imaging can be performed intraopera-ively, either noninvasively or with endoscopic ultra-ound probes, but resolutions are typically no higherhan 100 mm and physical contact or an index-atching medium is required [8]. A compact, high-
esolution, high-speed imaging technique capable ofonitoring tissue coagulation, cutting, and ablation
ntraoperatively at a localized site would permit image-uided surgical laser procedures to be monitored forontrolled therapy and subsequent reduction in iatro-enic injury.Optical coherence tomography (OCT)3 is a recently
eveloped micron-scale, real-time imaging technology9]. OCT is analogous to ultrasound, measuring thentensity of backreflected infrared light rather thancoustic waves. OCT was first applied clinically to im-ge and track retinal diseases in the transparent struc-ures of the human eye [10]. Recently, OCT imagingas been developed for imaging nontransparent tissues11–13]. Imaging depths are limited to a few millime-ers because of optical attenuation due to absorptionnd scattering; however, these imaging depths are suf-cient for many diagnostic applications including per-orming may types of intraoperative monitoring. Stud-es have shown the feasibility of using OCT as aurgical diagnostic tool [14] and for intraoperativeonitoring of microsurgical anastomoses of vessels
nd nerves [15]. In vivo studies have been performed inn animal model which demonstrate the feasibility ofatheter–endoscope-based OCT imaging of internal or-an systems including the gastrointestinal and pulmo-ary tracts [16].Several features, in addition to high resolution,ake OCT attractive for guiding surgical ablation.irst, OCT imaging is fiber based, allowing it to beeadily integrated with hand-held surgical probes,aparoscopes, catheters, and endoscopes [17]. Second,nlike ultrasound, OCT imaging is noncontact and cane performed through air. Third, OCT systems areompact and portable, an important consideration forhe operating suite. Finally, OCT imaging is performedt high speeds, allowing image data to be acquiredapidly over large regions of tissue.
The pathologic effects of laser-induced hyperthermia
76 JOURNAL OF SURGICAL RESEA
ithin biological tissue [18] and the thermal and me-hanical effects of laser tissue ablation have been pre-
Tedga3 Abbreviation used: OCT, optical coherence tomography.
iously analyzed and well characterized [19]. The ma-ority of studies, however, have used histologicalreparations at single time points to document theissue changes that occur. Dynamic changes in tissueptical properties have been characterized [20, 21], butave relied on spectroscopic transmission and reflec-ion data rather than image data. OCT has been usedo characterize the changes in the retina following la-er injury; however, images were acquired at 5-s inter-als [22]. In tissues exhibiting birefringence, such asuscle and tendon, polarization-sensitive OCT has
een used to detect birefringence changes followinghermal heating [23, 24]. Birefringent changes mayndicate early thermal injury, but detection is limitedo specific tissue types. No studies have been per-ormed using high-speed OCT to image real-timehanges in optical backscatter in highly scattering tis-ues during laser ablation.In this study, we investigate the feasibility of using
igh-speed OCT imaging to guide the placement andollow the dynamics of surgical laser ablation. Argonaser ablation at 514 nm, a wavelength frequently usedurgically, is performed in five different ex vivo ratrgans to assess OCT imaging performance and varia-ions between tissue types. The use of OCT to monitorblative therapy in real time may enable more preciseontrol of laser delivery and a reduction in iatrogenicnjury.
MATERIALS AND METHODS
Optical coherence tomography. Optical coherence tomography istechnique for performing high-resolution imaging in biological
issue [9]. OCT is somewhat analogous to ultrasound B-mode imag-ng except backreflections of light are detected rather than acousticaves. Whereas ultrasound pulse propagation can be measured elec-
ronically, the echo time delay of light used in OCT imaging cannote measured directly because the velocity of light is extremely highnd the resulting echo delays are extremely short. Therefore, inCT, a technique known as interferometry is used. This techniqueeasures the magnitude and echo time delay of reflected or back-
cattered light by comparing the reflected light beam to a beamhich travels a reference path of known delay.Figure 1 shows a schematic of the OCT system. The OCT system
ses a fiber-optic implementation of a Michelson-type interferometer9]. A laser or other light source with a broad spectral bandwidthshort coherence length) is coupled into the fiber-optic interferometernd split by a fiber coupler into a sample arm and a reference arm.high-speed optical delay line is used in the reference arm to vary
he optical path length of the reference arm. The optical beam fromhe sample arm of the interferometer is directed onto the object orpecimen to be imaged. Light is backreflected or backscattered fromicrostructural features internal to the specimen and has a different
cho time delay depending on its depth. When the echo time delay ofhe light from the specimen matches the echo time delay of lightraveling the reference path to within the coherence length of theight, interference is detected at the output of the interferometer.
H: VOL. 82, NO. 2, APRIL 1999
he interference is detected by a photodetector and demodulatedlectronically to give a measurement of the magnitude and echoelay of backscattered light. A pair of steering mirrors controlled byalvanometer actuators is used to scan the OCT imaging beamcross the specimen. By measuring the magnitude and echo delay of
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ackscattered light at different transverse positions on the specimen,two-dimensional data set is generated which represents the back-
cattering through a cross-sectional plane of the specimen. This dataet is displayed as an image using a gray-scale or false-color repre-entation.The axial resolution of the imaging is determined by the coherence
ength of the light source. The coherence length is inversely propor-ional to the spectral bandwidth; therefore, light sources with aroad spectral bandwidth will enable high resolutions. The studieseported here were performed using a commercially available low-oherent light source which is based on a super-luminescent diode/mplifier (AFC Technologies Inc., Hull, Quebec, Canada). Otherources of low-coherence light can also be used including superlumi-escent optical fiber sources or short-pulse lasers. The light sourcesed for these experiments had a center wavelength of 1.3 mm and a
ree space axial resolution of 18 mm. The axial resolution was deter-ined by measuring the point spread function from a mirror placed
t the sample position in the OCT apparatus. The transverse reso-ution was determined by the spot size of the incident beam withinhe tissue. The spot size was 30 mm which yielded a 1.1-mm confocalarameter (depth of field). The confocal parameter for the beam waselected to closely match axial and transverse resolutions whileaintaining a sufficient depth of field.The signal-to-noise ratio was 115 dB using 5 mW of incident power
n the specimen. For typical tissues, this sensitivity permits imagingo depths up to 3 mm. A single axial scan was acquired for everyweep of the optical delay line in the reference arm. For high-speedmage acquisition, the length of the axial scan was 3 mm and 256ixels were acquired during each axial scan, corresponding to an
FIG. 1. Optical coherence tomography schematic. OCT uses a fiissue. Two- and three-dimensional subsurface images are generatedrames per second and saved either digitally or to Super-VHS tape. Ohe dynamic effects of laser tissue ablation. SLD, super-luminescent
BOPPART ET AL.: OCT IMAG
xial pixel sampling distance of 11.7 mm/pixel. Images were gener-ted by assembling adjacent axial scans to form a two-dimensionalross-sectional image of the optical backscatter from within the spec-men. A total of 256 axial scans were acquired across a transverseistance of 3 mm, corresponding to a transverse pixel sampling
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istance of 11.7 mm/pixel. Images were displayed as the logarithm ofhe backscattered intensity versus position in gray scale on a com-uter monitor and simultaneously recorded to Super-VHS videoape. The acquisition rate was eight frames per second or 125 ms forach image.Laser ablation was performed using a continuous wave argon laser
perating predominantly at a wavelength of 514 nm. As shown inig. 1, 1–3 W of output power from the argon laser was coupled intofiber and focused to a 0.8-mm-diameter spot on the tissue surface.he argon beam was aligned and centered within the OCT imaginglane. A mechanical shutter was used to control the argon laserxposures on the sample. Prior to laser ablation, each specimen wasanipulated under OCT image guidance to locate a relatively uni-
orm region of tissue for the ablation site. During exposures, both thergon beam and the specimen remained stationary. This setup per-itted image acquisition immediately prior to and after exposure to
rack the optical changes occurring within the tissue during laserblation.
Specimen preparation and imaging. Sprague–Dawley rats wereuthanized with 42 mg/kg intraperitoneal injections of phenobarbitalNembutal). Rat organs including brain, liver, kidney, lung, andectus abdominis muscle were immediately resected, placed in 0.9%aline, and stored at room temperature prior to imaging. More than5 sites on five ex vivo rat organ tissues were OCT imaged. Ahree-dimensional micron-precision computer-controlled stage wassed to position the specimen under the OCT imaging beam and wassed to acquire three-dimensional data sets. Argon laser ablationas performed at discrete sites on the organs using varying argon
-optic interferometer to localize optical backreflections from withincanning the beam across the tissue. Images can be acquired at eightimaging is coincident with a continuous wave argon laser to monitorode.
277GUIDED LASER ABLATION
E-owers and exposure durations. Immediately following image acqui-ition, the location of the image plane was marked with India ink foregistration between OCT images and histology. Specimens werelaced in a 10% buffered solution of formalin for standard histolog-cal preparation. Histological sections, 5 mm thick, were sectioned
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nd stained with hematoxylin and eosin for comparison with ac-uired OCT images. The protocol for the care and handling of ani-als used in this study has been approved by the Massachusetts
nstitute of Technology Committee on Animal Care.
RESULTS
Before performing dynamic imaging studies, a singleblation site was imaged and thoroughly investigatedo determine what optical changes within tissue, fol-owing laser ablation, were detectable with OCT. A0-s, 3-W argon laser exposure was used to form anblation crater in rat rectus abdominis muscle. Thepecimen was then translated on a multiaxis transla-ional stage to perform 3-D imaging of the crater. Sixtyross-sectional images were acquired at 100-mm inter-als to produce a 3-D data set. One projection from thisata set is shown in Fig. 2A. The deep crater can beeen in the center of the projection. Surrounding therater is an elevated region of damage which decreasesadially outward from the center.
A sequence of cross-sectional images from this dataet is shown in Fig. 2B. These provide information onhe depth- and distance-dependent distribution of thehermal energy. The number in each image refers tohe distance from the center of the crater at which theross-sectional image was acquired. At a distance of.75 mm from the center, small distortions in the mus-le layers (arrow) are observed, although little changeas occurred at the tissue surface. At 3.0 mm, thermal
njury has elevated the tissue surface and distorted annternal low-backscattering layer. At 2.25 mm, a regionf increased optical backscatter has appeared. Thisegion appears deeper in the 1.5-mm section. At 0.75m from the crater center, carbonization of the tissue
esults in significant shadowing of underlying struc-ures. A cross-section through the center of the craters shown in the last image. The vertical bands of lowackscatter adjacent to the crater are due to shadowingy the vertical crater walls which are lined with car-onized tissue. Carbonized tissue scatters and absorbshe incident light, decreasing imaging penetration.
Surrounding the deep crater in Fig. 2A lie concentricings of tissue damage. These are the result of a radialhermal distribution outward from the site of the inci-ent beam. The differences between each ring areikely the result of different thermal damage mecha-isms, if one assumes the tissue is homogeneous. Toptimally assess these radial distributions, the 3-Data set is resectioned in the en face plane, as would beiewed from the surface and with increasing depth intohe tissue. Resectioned slices are shown in Fig. 2C. Theumber in each figure refers to the depth of the recon-
78 JOURNAL OF SURGICAL RESEA
tructed plane from the surface. The shallow planes150–600 mm) show some empty spaces (lower whiteegions) because the axes of the tissue block were notxactly orthogonal to the translation stage axes. The
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mage at 150 mm shows a relatively uniform ring con-istent with the elevated region appearing in the 3-Drojection of Fig. 2A. At greater depths (450–900 mm),owever, a multiple-ring pattern is evident, possibly
ndicating different zones of tissue damage. An alter-ative explanation may be a concentric elevation ofirefringent tissue layers presenting as regions of lownd high optical backscatter. At a depth of 1350 mm,he crater bottom is approached and the concentricings have diminished. At 1500 mm, the region of highackscatter in the center represents tissue immedi-tely below the crater bottom and the white ring ofeduced optical backscatter is due to the attenuation ofhe incident beam by the carbonized tissue lining theearly vertical crater walls.Ablation dynamics represented as changes in optical
ackscatter within the OCT images are shown in theemaining figures. Figure 3 is an image sequencehowing thermal injury in kidney tissue from 1-W, 3-srgon laser exposure. The OCT images illustrate theelatively homogeneous nature of the outer cortex ofhe kidney. Imaging penetration in this tissue is lim-ted to '1 mm. An increased region of optical backscat-er is first observed at 0.1 s. This is followed by anutward propagating front of increased backscatterhown at 0.6 s. A region of low backscatter develops athe center of the lesion (1.6 s). Backscatter from thisegion then increases over time (3.0 s). This exposureas below the threshold for surface membrane rup-
ure, tissue ejection, and crater formation. Based onmpirical ablation observations, membrane ruptureould likely have occurred within the following second
f the exposure had continued. OCT image guidancend feedback enabled laser exposure to be terminatedmmediately prior to membrane rupture. The corre-ponding histology is also shown, indicating a region ofoagulated tissue (arrows) with no tissue fragmenta-ion.
The ablation threshold at which tissue is ejected isocumented by a sequence of images showing thermalnjury in rat liver. A superthreshold ablation sequenceith corresponding histology is shown in Fig. 4. Thexposure is allowed to continue for 6 s, resulting injection of tissue and crater formation. The changes inptical backscatter observed prior to membrane rup-ure are similar to those observed for the subthresholdxposure in Fig. 3. The superthreshold histology showsarked tissue ablation and fragmentation within the
esion crater. Below the crater extends a zone of coag-lated tissue (arrows) which is not fully imaged withCT due to the poor imaging penetration through the
arbonized crater wall.
H: VOL. 82, NO. 2, APRIL 1999
In contrast to the ablation of tissues such as livernd kidney which have a high absorption coefficient atrgon laser wavelengths, Fig. 5 shows the ablation ofrain tissue. The lower absorption coefficient implies
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279BOPPART ET AL.: OCT IMAGE-GUIDED LASER ABLATION
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280 JOURNAL OF SURGICAL RESEARCH: VOL. 82, NO. 2, APRIL 1999
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BOPPART ET AL.: OCT IMAGE-
hat longer exposure durations are required for theame incident power to produce similar effects. Thisrend was observed for 1 W incident on brain tissue for0 s. In this sequence, significant vacuolization andissue heating occur (5.5–13 s) before the surface mem-rane is ruptured at 14 s. Membrane rupture is fol-owed by ejection of tissue from the lesion, evolution of
smoke plume (16–17 s), and crater formation (20 s).Figure 6 shows the ablation of lung tissue. Inflated
ir-filled alveolar spaces represent a large portion ofhe tissue space making this tissue more inhomoge-eous compared to previous specimens. A 1-s exposuref 1-W argon collapses most of the alveoli (0.5–1.0 s)esulting in the rapid deflation of the lung followed byhe formation of a shallow crater.
DISCUSSION
We have demonstrated the use of high-speed, real-ime OCT imaging for guiding the placement and mon-toring the dynamic changes of surgical laser ablationn a variety of tissues. OCT performs high-resolution,ross-sectional imaging and also gives information onhe optical scattering properties of tissue. Reflectionsnd optical backscattering occur in tissue due to differ-nces in index of refraction and organization of tissueicrostructure including architectural morphology
nd cellular and subcellular organization. OCT canccurately monitor distributions of ablative damageecause the thermal injury disrupts the normal opticalroperties of the tissue. These changes were illustratedn Fig. 2. In Fig. 2B, the left side of the image at 3.75
m shows normal rectus abdominis muscle. The lay-red structure of the muscle is disrupted near theblation site and replaced by a more homogeneousegion. The concentric rings of tissue damage in Fig. 2Chow interesting patterns corresponding to the radialistribution of thermal energy, whereas normal tissueould show a relatively homogeneous pattern at eachepth.Carbonization at the surface of the tissue is an un-
esirable effect which hinders controlled surgical abla-ion. This effect also represents a limitation for OCTmaging. The high temperatures char tissue, creating aarbonized layer at the surface which rapidly absorbsnd scatters both the incident argon and OCT imagingeams. Because of this, OCT imaging penetration iseduced and shadowing artifacts, identified by verticalow-backscattering streaks, appear in the images (Fig.B, 0 mm). Fortunately, tissue carbonization is limitedo applications of continuous high-power irradiationnd studies have identified laser parameters and oper-tive procedures to reduce this effect [25].
281GUIDED LASER ABLATION
The tissue response to thermal injury is highly de-pendent on the wavelength of the incident light and thetissue absorption coefficient. We have demonstratedOCT imaging in a variety of tissue types to illustrate
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his effect. Hemoglobin is one of the dominant absorb-rs of visible wavelengths in biological tissue [26]. Forhe same incident argon power (1 W), the hemoglobin-ich rat liver and kidney (absorption coefficients ma 52 and 1.21 cm21, respectively) exhibited vacuolization,blation, and surface membrane rupture within 1–3 s
FIG. 5. Brain ablation. Ablation sequence from 1-W, 20-s argonemoglobin-rich tissue such as liver and kidney results in longer
llustrates subsurface tissue vacuolization followed by membrane ru
82 JOURNAL OF SURGICAL RESEA
FIG. 6. Lung ablation. Ablation sequence from 1-W, 1-s argon exorphological differences from more solid organs. Thermal laser ablati
epresents 1 mm.
ompared to 14 s for brain tissue (ma 5 0.19 cm21). Thiss important clinically because often it is difficult toredict the tissue response to incident laser radiation.bsorption and scattering coefficients, which vary be-
ween tissue types, will vary the efficiency of laserblation. The use of OCT imaging in Fig. 3 provided
osure. The lower absorption coefficient of brain tissue compared toposures required to produce equivalent tissue ablation. Sequencere and tissue ejection. Bar represents 1 mm.
H: VOL. 82, NO. 2, APRIL 1999
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eedback to terminate laser ablation prior to mem-rane rupture and tissue ejection, demonstrating thebility to monitor tissue response during laser abla-ion.
The resolutions of OCT are sufficient to resolvehanges in optical backscatter and identify regions ofblated tissue. However, the kidney, liver, and brainpecimens appear relatively homogeneous at the cur-ent resolutions of OCT. Improvements in resolutionill improve the ability to identify tissue microstruc-
ure and detect more subtle thermal damage in theseissues. The axial and transverse resolutions in OCTre independent. The axial resolution is inversely pro-ortional to the bandwidth of the optical source. Hence,arger spectral bandwidths can improve axial resolu-ion. Axial resolutions as high as 1.9 mm have beenchieved using broader bandwidth laser sources athorter wavelengths near 800 nm [27, 28]. Shorteravelengths, however, are absorbed and scatteredore in biological tissue, resulting in decreased imag-
ng penetration. The transverse resolution is deter-ined by the beam-focusing optics. Higher transverse
esolutions are possible, but at the expense of de-reased depth of field. New laser sources and image-rocessing techniques for improving imaging resolu-ion are topics of active investigation.
The eight frames per second acquisition rate used inhis study was sufficient to image the dynamic changesuring thermal ablation. More rapid events, such asembrane rupture and tissue ejection, occurred on
ime scales too fast for OCT imaging. Still, OCT hastility for image guidance of faster processes becauseCT can image the integrated effects of ablation. It isot surgically important to capture explosive or disrup-ive effects on faster time scales as long as the result-nt effects can be accurately monitored. Nonetheless,CT imaging at faster rates is interesting for investi-ational purposes. Video rate acquisition (30 fps) isossible by either reducing the image size or scanninghe optical delay line faster. A resonant scanner can bencorporated into the delay line; however, the nonlin-ar sinusoidal output would require real-time signalnd image processing to correct the images.The use of an argon laser for surgical ablation is only
epresentative of a broad range of interventional tech-iques that can be guided using OCT. The 514-nmavelength of the argon laser is close to the 532-nmavelength frequently used for tissue cutting and co-gulation during surgery and is readily absorbed inissue. In principle, OCT can be used with other sur-ical laser wavelengths and with both pulsed and con-inuous wave laser systems. Because these techniques
BOPPART ET AL.: OCT IMAG
re optic based, there is potential for integration into aingle system with coincident beams for imaging andurgery. In addition, the fiber-optic-based design ofCT allows compact delivery and permits OCT imag-
ng to be integrated with scalpels, scissors, and biopsyorceps. Real-time OCT imaging with these integratedevices could be used to guide the operator in order tovoid injury to sensitive tissue structures such as ad-acent nerves, to differentiate tissue or tumor bound-ries, or to guide the operator toward specific morphol-gies. These integrated systems are currently undernvestigation.
OCT guidance during surgical interventions has aide range of application across surgical specialties.ecause of the micron-scale resolutions afforded byCT, not only surgical image guidance, but also intra-perative diagnostics of surgical tissue without havingo physically resect or biopsy specimens may be possi-le. Neurosurgery and cardiovascular surgery are twopecialties where OCT may be most helpful [29]. Iden-ification and surgical treatment of pathologies muste performed with minimal tissue resection, sparingensitive adjacent tissue. The single optical fiber deliv-ry of the OCT beam is suitable for integration withinimally invasive techniques. Fiber-based OCT im-
ging may be possible at physically restrictive sitesrequently encountered in the head and neck, the lowerespiratory tract, the fallopian tubes, and within theascular system [30].In conclusion, these results demonstrate the ability
f high-resolution, high-speed OCT to guide the place-ent of laser ablation sites and image the dynamic
hanges that occur during thermal tissue ablation.hese results suggest that OCT may play a role in
mage-guided surgical procedures. Future in vivo stud-es are necessary to demonstrate the performance ofCT during intraoperative scenarios.
ACKNOWLEDGMENTS
Dr. Juergen Herrmann was visiting from the University of Erlan-en. The authors thank Drs. Brett E. Bouma and Gary J. Tearney ofhe Wellman Laboratory of Photomedicine, Massachusetts Generalospital, for their technical input. We greatly appreciate the assis-
ance of Ms. Christine Jesser of Massachusetts General Hospital ands. Jennifer Libus of King’s College for handling tissue specimens
nd Ms. Cindy Kopf for her assistance in the preparation of theanuscript.
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