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Chapter 34. Laser Medicine and Biomedical Imaging

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Laser Medicine and Biomedical Imaging RLE Groups Laser Medicine and Medical Imaging Group, Optics and Quantum Electronics Group Academic and Research Staff Professor James G. Fujimoto Visiting Scientists and Research Affiliates Dr. Yu Chen, Mariana Carvalho, Dr. Robert Huber, Dr. Norihiko Nishizawa (Nagoya University, Japan), Adelina L. Paunescu (New England Eye Center), Dr. Alphan Sennaroglu (Koc University, Turkey), Dr. Kenji Taira, Dr. Maciej Wojtkowski, Dr. Rebecca Younkin Graduate Students Desmond C. Adler, Aaron D. Aguirre, Paul R. Herz, Pei-Lin Hsiung, Tony H. Ko, Andrew M. Kowalevicz, Vikas Sharma, Vivek J. Srinivasan, Aurea Zare Technical and Support Staff Dorothy A. Fleischer, Donna L. Gale 1. OCT Technology Optical coherence tomography (OCT) is an emerging diagnostic technology, developed by our research group and collaborators in 1991 [1], and it has been investigated for applications in several medical fields [2]. OCT enables the visualization of tissue microstructure in situ and in real time with resolutions in the 1-10 µm range, and previous studies have demonstrated that changes in tissue architectural morphology associated with neoplasia can be identified [3-6]. Clinical OCT systems often use superluminescent diodes (SLDs) that enable imaging with 10-15 µm axial resolution. These resolutions are typically insufficient for identifying neoplastic changes for cancer detection or tissue morphological and structural features for the visualization of other pathologies. Advances in solid-state lasers and nonlinear fiber technology have enabled the development of ultrahigh resolution and spectroscopic OCT techniques that promise to improve tissue differentiation and image contrast. The recent development of spectral domain OCT techniques enables high speed in vivo imaging of biological tissues, which has strong potential for imaging large tissue volumes in three dimensions for clinical diagnosis. References [1] D. Huang, E.A. Swanson, C.P. Lin, J.S. Schuman, W.G. Stinson, W. Chang, M.R. Hee, T.

Flotte, K. Gregory, C.A. Puliafito, and J.G. Fujimoto, "Optical coherence tomography," Science 254(5035): 1178-1181 (1991).

[2] J.G. Fujimoto, "Optical coherence tomography for ultrahigh resolution in vivo imaging," Nature Biotechnology 21(11): 1361-1367 (2003).

[3] J.G. Fujimoto, M.E. Brezinski, G.J. Tearney, S.A. Boppart, B. Bouma, M.R. Hee, J.F. Southern, and E.A. Swanson, "Optical biopsy and imaging using optical coherence tomography," Nature Medicine 1(9): 970-972 (1995).

[4] M. Brezinski, G. Tearney, B. Bouma, J. Izatt, M. Hee, E. Swanson, J. Southern, and J. Fujimoto, "Optical coherence tomography for optical biopsy: properties and demonstration of vascular pathology," Circulation 93: 1206-13 (1996).

[5] J.G. Fujimoto, C. Pitris, S.A. Boppart, and M.E. Brezinski, "Optical coherence tomography: an emerging technology for biomedical imaging and optical biopsy," Neoplasia 2(1-2): 9-25 (2000).

[6] W. Drexler, U. Morgner, R.K. Ghanta, F.X. Kaertner, J.S. Schuman, and J.G. Fujimoto, "Ultrahigh resolution ophthalmic optical coherence tomography," Nature Medicine 7: 502-507 (2001).


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1.1 Ultrahigh Resolution OCT Imaging The need for simple, robust sources of broadband light exists in fields such as spectroscopy as well as biomedical imaging applications such as Optical Coherence Tomography (OCT).[1, 7, 8] The achievable longitudinal resolution in OCT is inversely related to the bandwidth of the source. Standard, ~10 µm axial resolution OCT imaging can be performed using superluminescent diodes (SLDs) that have ~20-30 nm FWHM bandwidths centered near 800 nm for ophthalmic imaging, while SLDs at longer wavelengths ~1300 nm are used for imaging of tissue. These sources are relatively inexpensive, portable, and have turn-key operation suitable for clinical use. However, they provide limited image resolutions due to their narrow bandwidths. Recently, OCT imaging with an axial resolution of ~1 µm has been demonstrated by our group using a Ti:Al2O3 laser with a bandwidth of ~300 nm.[9] This femtosecond laser source was developed by our group and produces record pulse durations of ~5 fs, which correspond to approximately two optical cycles.[10] In ophthalmology, ultrahigh resolution OCT imaging with axial resolutions of 2-3 µm has been shown to significantly improve the visualization of retinal morphology[9, 11] and retinal pathologic changes.[12-15] Improved axial resolution enables improved visualization of intraretinal layers such as the photoreceptor layers, ganglion cell layer, plexiform and nuclear layers. Studies have also shown the potential of ultrahigh resolution to provide additional information over standard techniques on the pathogenesis or management of macular diseases.[12, 14-17] Ultrahigh resolution OCT endoscopy also has the potential as an imaging tool for early neoplastic changes in human tissue.[18] However, ultrahigh resolution systems using femtosecond lasers are expensive and complex, limiting their widespread use. Our group has conducted ongoing work to develop novel low-coherence light sources. We have introduced several alternative sources for ultrahigh resolution imaging. These sources focus on reducing the cost, increasing the reliability and portability of systems, while maintaining high performance capability. 1.2 Low-threshold Ti:Sapphire laser for Ultrahigh Resolution Ophthalmic Imaging Sponsors National Institutes of Health – R01-CA75289-06, R01-EY11289-18 National Science Foundation – ECS-01-19452, BES-0119494 Air Force Office of Scientific Research – F49620-01-1-0186, F49620-01-01-0084 The Whitaker Foundation Project Staff Andrew Kowalevicz, Tony H. Ko, Dr. Maciej Wojtkowski, Professor James G. Fujimoto, Dr. Jay S. Duker (NEEC), Dr. Joel Schuman (UPMC) Kerr lens modelocked (KLM) Ti:Al2O3 lasers can generate extremely short pulse durations with broad bandwidths that are particularly useful in biomedical imaging [19]. A standard Kerr lens modelocked laser operating with a 5W pump can produce output powers of 500 mW and bandwidths in excess of 150 nm [20]. Unfortunately, the high cost of today’s femtosecond lasers severely limits their widespread use. The cost of femtosecond Ti:Al2O3 lasers is strongly dependent on the pump power requirements. Diode pumped solid-state lasers capable of generating 5 W can be prohibitively expensive, while lasers generating several hundred mW are considerably more affordable. We have developed a low-threshold, low-cost, compact laser system. By replacing standard mirrors and prisms with 3rd generation double-chirped mirrors (DCMs), we are able to develop a broadband portable source to be used in the clinical setting for ultrahigh resolution OCT imaging with 3 µm axial resolution [21]. Figure 1 shows the spectrum and point spread function of the


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low-threshold source. Figure 2 shows a comparison between standard (a) and ultrahigh resolution (b) OCT images of the normal human fovea.

Figure 1. Spectrum and point spread function of the low-threshold Ti:Sapphire laser developed by our group for OCT imaging in the ophthalmology clinic.

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Figure 2. In vivo standard resolution OCT3 (top) and ultrahigh resolution OCT images (bottom) of a normal human fovea at approximately the same site. Resolutions were ~10-15 µm axial x 15 µm transverse (above) and ~3 µm axial x 15 µm transverse (below) respectively.



1.3 Broadband SLD for Ultrahigh Resolution Ophthalmic Imaging Sponsors National Institutes of Health – R01-CA75289-06, R01-EY11289-18 National Science Foundation – ECS-01-19452, BES-0119494 Air Force Office of Scientific Research – F49620-01-1-0186, F49620-01-01-0084 The Whitaker Foundation Project Staff Tony H. Ko, Dr. Maciej Wojtkowski, Professor James G. Fujimoto, Vladimir Shidlovski (Superlum Diodes Ltd.), Sergei Yakubovich (Superlum Diodes Ltd.), Dr. Jay S. Duker (NEEC), Dr. Joel Schuman (UPMC) While superluminescent diode sources have proven to be inexpensive and sufficiently stable for use in a clinical setting, conventional SLD sources typically have limited bandwidths and output powers. Broadband SLD sources are made possible by multiplexing two or more diodes with different center wavelengths. These sources are more inexpensive and user-friendly than Ti:Sapphire sources which are typically used to achieve ultrahigh axial image resolutions. For applications where light exposure levels are limited by safety considerations, such as retinal imaging, broadband SLD sources provide sufficient power levels for OCT imaging. Recently, our group and others have demonstrated ophthalmic retinal imaging with broadband SLD light sources.[16, 22] The broadband SLD source was a new prototype developed by Superlum Diodes, Ltd. and consists of two spectrally multiplexed, independently driven SLD diodes operating at 840 nm and 920 nm, respectively. Figure 3 shows individual (a) and combined (b) SLD spectra, along with the point spread function on a linear (c) and logarithmic (d) scale. An SLD diode based on a SQW (GaAl)As heterostructure [23] and a second recently-developed SLD diode based on a SQW (InGa)As/(AlGa)As heterostructure with graded-index waveguides[24] were used. The SLDs have optical outputs with overlapping spectral bands that can be combined with a broadband fiber coupler. A custom built broadband single-mode fiber Y-coupler based on fiber GRIN microlenses and miniature partially transmitting mirrors with dielectric multilayer coatings was developed to spectrally multiplex the SLD outputs. In the spectral range of 750-1100 nm, a coupling ratio of close to 50% and an insertion loss of less than 1 dB were achieved with this coupler. The combined SLD output provides 155 nm of bandwidth at a center wavelength of 890 nm with > 4 mW of CW output power. Compared to a solid-state laser, this broadband SLD source is quite compact and has a total footprint of only 31 x 26 x 15 cm including the power supply, or 31 x 15 x 5 cm for the SLD and multiplexer package alone. The total cost of this light source is about an order of magnitude less than that of a typical commercial Ti:Sapphire femtosecond laser. An image of the normal human macula taken with the Broadlighter source is shown in Figure 4.


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Figure 3. (a) Individual output spectra of the two superluminescent diodes. (b) Fiber-coupled multiplexed spectrum of the broadband SLD source. (c) Coherence point spread function of the broadband SLD source. (d) Logarithmic demodulated coherence point spread function. A 3.0 OD filter was used to prevent detector saturation.

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Figure 4. In vivo ultrahigh resolution OCT image of the human retina taken with a broadband SLD light source. Image axial resolution in the retina was about 3.2 µm and transverse resolution was about 15-20 µm. All the major intraretinal layers can be clearly seen in this ultrahigh resolution OCT image.



1.4 Cr:Forsterite for Endoscopic Imaging Sponsors National Institute of Health – R01-CA75289-06, R01-EY11289-18 National Science Foundation – ECS-0119452, BES-0119494 U.S. Air Force Office of Scientific Research – F49620-01-01-0084, F49620-01-01-0186 Project Staff Paul Herz, Dr. Yu Chen, Pei-Lin Hsiung, Aaron D. Aguirre, Professor James G. Fujimoto (MIT), Dr. Hiroshi Mashimo, Dr. Saleem Desai, Dr. Mihr Wagh (VAMC), Amanda Koski, Dr. Joseph Schmitt (LightLab Imaging) Epithelial cancers of the gastrointestinal tract, reproductive tract, and the respiratory tract comprise the majority of cancers encountered in internal medicine. Many epithelial cancers are preceded by pre-malignant changes, such as dysplasia. Conventional screening methods often rely on the gross morphological characteristics of tissues. Biopsy and histopathology is the standard for the diagnosis of dysplasia or carcinoma, but can suffer from sampling errors and is cumbersome for screening and surveillance applications. Our hypothesis is that OCT can function as a form of “optical biopsy” which can perform microstructural imaging of tissue morphology in situ without excision. The specific aim of this work is to perform imaging of architectural morphology of the human upper gastrointestinal tract to explore the feasibility of ultrahigh resolution OCT for the identification of morphologies characteristic of early neoplastic changes. A compact, broadband Cr4+:Forsterite laser combined with a nonlinear fiber was used to generate a spectral bandwidth greater than 200 nm at a center wavelength of 1250 nm. An output power of 50 mW was coupled into a single-mode fiber. The Cr4+:Forsterite laser was somewhat similar to one previously demonstrated [25, 26]; however, it used broadband double-chirped mirrors [27] to compensate intra-cavity dispersion and achieve a compact, prismless laser design. A compact Yb fiber laser was used as the pump source. Figure 5(a) shows the optical spectrum generated by the Cr:Forsterite laser and nonlinear fiber. The bandwidth is 210 nm FWHM. Due to bandwidth limitations in the optical circulator, shorter wavelengths were attenuated and the transmitted spectrum was reduced to 150 nm bandwidth (also shown). The measured axial point spread function shown in Figure 5(b) has a resolution of 5 µm in air, close to the calculated theoretical value of 4.6 µm for the transmitted bandwidth. This corresponds to an axial resolution of ~3.7 µm in tissue, assuming an index of refraction of ~1.37, and is a three-fold improvement in resolution when compared to previous endoscopic OCT systems. The system detection sensitivity was measured to be 102 dB with 10 mW of power on the sample. Previous data has demonstrated that OCT can image features of architectural morphology which are relevant for the diagnosis and monitoring of early neoplastic tissue changes. We have performed in vitro survey imaging and histopathologic correlations of normal as well as abnormal human tissue pathologies including Barrett’s esophagus, adenocarcinoma of the esophagus, ulcerative colitis, colonic dysplasia, adenomatous polyps, colorectal carcinoma, cervical intraepithelial neoplasia, cervical carcinoma, uterine dysplasia, and endometrial carcinoma in order to evaluate the ability of OCT to differentiate early neoplastic changes and other tissue abnormalities with a predisposition to malignancy. We have performed imaging of normal tissue using OCT endoscope-catheters in an in vivo rabbit animal model[18], as shown in Figure 6. Survey imaging studies were performed on the gastrointestinal, urinary, reproductive, and respiratory tracts and correlated with histology. We also have performed endoscopic imaging of patients with Barrett’s esophagus. OCT imaging demonstrates the ability to differentiate changes in esophageal architectural morphology associated with the development of Barrett’s.


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Figure 5. Figure 5(a) shows the optical spectrum generated by the Cr4+:Forsterite laser and nonlinear fiber. The bandwidth is 210 nm FWHM. Due to bandwidth limitations in the optical circulator, shorter wavelengths were attenuated and the transmitted spectrum was reduced to 150 nm bandwidth (also shown). The measured axial point spread function shown in Fig. 5(b) has a resolution of 5 µm in air, close to the calculated theoretical value of 4.6 µm for the transmitted bandwidth.

Figure 6. (a) In vivo OCT image and (b) histology of rabbit esophagus and trachea viewed intraluminally from the esophagus. Tracheal hyaline cartilage (hc) between the tracheal mucosa and trachealis muscle is well defined. The image demonstrates the ability of the endoscopic OCT system to image deeply within the tissue. Trichrome stain was used to highlight cartilage and muscle layers.



1.5 All-Fiber Continuous-Wave Raman Continuum Light Source Sponsors National Institutes of Health – R01-CA75289-06, R01-EY11289-18 National Science Foundation – ECS-01-19452, BES-0119494 Air Force Office of Scientific Research – MFEL F49620-01-1-0186, F49620-01-01-0084 Poduska Family Foundation Fund for Innovative Research in Cancer Philanthropy of Mr. Gerhard Andlinger Project Staff Pei-Lin Hsiung, Dr. Yu Chen, Tony H. Ko, Professor James G. Fujimoto High performance, short coherence length light sources with broad bandwidths and high output powers are critical for high-speed, ultrahigh resolution OCT imaging. We demonstrate a new, high performance light source for ultrahigh resolution OCT.[28] Bandwidths of 140 nm at 1300 nm center wavelength with high output powers of 330 mW are generated by an all-fiber Raman light source based on a continuous-wave Yb-fiber laser-pumped microstructure fiber. The schematic of the light source and output bandwidth are shown in Figure 7. The light source is compact, robust, turnkey and requires no optical alignment. In vivo, ultrahigh resolution, high-speed, time domain OCT imaging with <5 µm axial resolution is possible using a high-power, continuous-wave, all-fiber pump light source. This new light source promises to enable ultrahigh resolution, high-speed OCT imaging with lower cost and complexity than with femtosecond laser-based light sources. Microstructure fibers typically have been pumped with femtosecond lasers to provide the peak powers necessary to initiate nonlinear effects for continuum generation under conditions of anomalous or near-zero dispersion. However, the use of high intensity femtosecond pulses for broadband continuum generation under these conditions leads to severe spectral modulation of the continuum in the vicinity of the pump wavelength. This spectral modulation produces sidelobes and reduced contrast in the interferometric axial point spread function, which determines the axial resolution in OCT. Femtosecond pumping of microstructure fibers may also result in excessive temporal instability of the continuum and nonlinearly amplified quantum noise, which can lead to excess intensity noise [29]. An alternative to using high peak powers is to increase the effective nonlinear interaction length of the Raman interaction, which is governed by optical losses in the fiber and dispersive walk-off between the pump and continuum pulses. The use of longer pump pulses reduces this dispersive walk-off effect. Stimulated Raman scattering has been shown to be the principle nonlinearity for continuum generation by using nanosecond-scale pump pulses [30]. Recently, the possibility of low peak power and even continuous-wave, multiwatt Raman continuum generation in highly nonlinear fibers has been demonstrated [31]. Continuous-wave pumping of nonlinear fibers can enable the development of robust and turnkey continuum light sources which require no optical alignment, enabling high-speed, ultrahigh resolution OCT imaging in a wide range of applications outside the laboratory. The axial resolution of <5 µm is comparable to that achieved using much more complex femtosecond solid state sources such as the Cr:Forsterite laser. Output powers are comparable to or better than those available from femtosecond lasers, and more than one order of magnitude higher than superluminescent diode sources. Since this light source is all-fiber, it requires no alignment and provides completely turnkey operation. The source is extremely compact, measuring only 25 x 25 x 20 cm, and can easily be integrated into portable OCT systems. Excess noise is relatively high, but can be reduced using dual-balanced detection. Further reduction in noise should be possible by reducing parasitic feedback effects in the fiber splices and by techniques such as active seeding of the Raman process. Broader output bandwidths should be achievable using higher pump power Yb-fiber lasers and different fiber geometries. The bandwidth will ultimately be limited by the water absorption of microstructure fibers which reduces the effective interaction length of the Raman scattering. Rendering and segmentation of sweat ducts in the human fingerpad are shown in Figure 8.


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The current cost of the source is still relatively high due to the cost of the microstructure fiber. However, since demand for microstructure fiber is increasing, the price of these fibers is decreasing and can be expected to approach that of other specialty fibers. The cost of continuous-wave Yb-fiber lasers has also dramatically decreased in recent years. The cost of fiber Raman continuum light sources will therefore be significantly lower than that of bulk solid-state femtosecond laser-based light sources. The high performance and ease of use of the fiber Raman continuum source promises to enable a wide range of new ultrahigh resolution, high-speed, OCT imaging applications.

Figure 7. Left: Schematic of Raman continuum source. SMF: Single-mode fiber. Right: Output spectrum of the Raman continuum source. The output was 5.5W total, with 2.3 W in the spectral range from 1090 to 1370 nm. The output was filtered using a special WDM coupler to remove the pump wavelength and shape the spectrum.

Figure 8. (a) Rendering showing en face slice of human fingerpad created from cross-sectional images. Spiraling sweat ducts are seen along the fingerpad ridges (2.25 x 0.75 x 1.8 mm). (b) 3D rendering and segmentation of sweat ducts. Duct density as well as individual ducts can be assessed.



1.6 Femtosecond Nd:Glass Laser Sponsors National Institute of Health –R01-CA75289-06, R01-EY11289-18 National Science Foundation – ECS-01-19452, BES-0119494 Air Force Office of Scientific Research, Medical – F49620-01-1-0186, F49620-01-01-0084 Project Staff Aaron D. Aguirre, Dr. Robert Huber, Vikas Sharma, Professor James G. Fujimoto One of the key limitations in OCT has been the lack of compact, high performance, low coherence light sources with sufficient bandwidth and power to enable high resolution, real time imaging. High nonlinearity, air-silica microstructure fibers or tapered fibers can generate a broadband continuum that spans the visible to the near infrared using femtosecond pulses. These fibers have enhanced nonlinearity because of dispersion characteristics of these fibers, which shift the zero dispersion to shorter wavelengths, and the small core diameters, which provide tight mode confinement. High numerical aperture fibers also have been used with femtosecond Ti:Sapphire lasers to achieve bandwidths of up to 200 nm [32]. Continuum generation from a femtosecond Ti:Sapphire laser with air-silica microstructured photonic crystal fibers was demonstrated to achieve OCT image resolutions of 2.5 µm in the spectral region 1.2 µm to 1.5 µm [33], resolutions of 1.3 µm in the spectral region 800 nm to 1400 nm, [34] and resolution <1 µm in the spectral region of 550 nm to 950 nm.[35] While high resolution imaging was demonstrated, femtosecond Ti:Sapphire lasers are generally expensive and bulky. Previous studies also used galvanometer or voice coil scanned retroreflector delay lines which had limited axial scan speed, so that real time imaging was not performed. Ultrahigh resolution, real time OCT imaging is demonstrated using a commercially available, compact femtosecond Nd:Glass laser that is spectrally broadened in a high numerical aperture single mode fiber.[36] A reflective grating phase delay scanner enables broad bandwidth, high-speed group delay scanning. Ultrahigh resolution, high speed OCT imaging is achieved at 1 µm center wavelength with <5 µm axial resolution in free space (or ~4 µm in tissue) and 93 dB sensitivity. Figure 9 shows the spectrum (a) and point spread function (b). Axial scan speeds of ~11 m/s are demonstrated, which is sufficient for real time imaging. This system is compact and well suited for in vivo ultrahigh resolution OCT imaging studies. The 1 µm wavelength is important because it has improved image penetration depths when compared to 800 nm and ultrahigh resolution imaging requires less bandwidth than at 1.3 µm wavelengths.[37] Figure 10 shows images of the hamster cheek pouch with the Nd:Glass laser as the light source.

Figure 9. (a) Plot of the optical spectrum of the Nd:Glass laser broadened in the high numerical aperture (NA) single mode fiber. (b) Axial point spread function.


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Figure 10. In vivo ultrahigh resolution OCT images at 1 µm wavelength. Images were acquired at 4 frames per second and are 2.5 mm transverse x 1.2 mm axial with 11 µm x 4 µm resolution. (a,b) Images of the hamster cheek pouch showing squamous epithelium and connective tissue as well as a junction between two vessels. (c) Image of the nail bed showing the stratum corneum and the junction between the epidermis and the dermis. (d) Image of human volar finger pad showing stratum corneum and sweat ducts.

1.7 All-Fiber Femtosecond Fiber Laser Continuum Sponsors National Institute of Health –R01-CA75289-06, R01-EY11289-18 National Science Foundation – ECS-01-19452, BES-0119494 Air Force Office of Scientific Research, Medical – F49620-01-1-0186, F49620-01-01-0084 Project Staff Dr. Norihiko Nishizawa, Aaron D. Aguirre, Dr. Robert Huber, Vikas Sharma, Professor James G. Fujimoto Because optical scattering in tissues is reduced and the penetration depth is increased at longer wavelengths, the wavelength region around 1.4-1.6 µm is of interest for OCT imaging. This wavelength region is also attractive for the characterization of optical devices or spectroscopic applications measuring water absorption. Our group performed the first demonstration of an all-fiber scheme for real-time, high-resolution OCT imaging in the 1.4-1.7 µm wavelength region.[38] A low-noise supercontinuum with 38 mW of power and a 180-nm bandwidth is generated by use of a high-power, stretched-pulse, passively mode-locked fiber laser and highly nonlinear fiber. The spectrum of the fiber laser and the high-nonlinearity fiber are shown in Figure 11(a). Figure 11(b) shows the results of imaging human skin. In vivo imaging of human skin is demonstrated with the supercontinuum source and a high-speed OCT system. This light source is stable, compact, and self-starting and should be useful for a wide range of ultrahigh resolution, high-speed OCT imaging applications.



Figure 11. Optical spectrum on a linear scale of the output from the fiber laser and the supercontinuum generated in a high-nonlinearity fiber. The bandwidth can be broadened from 70 to 180 nm with a negligible increase in noise.

Figure 12. High-speed in vivo OCT image of human skin. The top surface is the transparent gel (g). The stratum corneum (sc), epidermis (e), sweat duct (s), and dermis (d) are clearly visible.

References [1] D. Huang, E.A. Swanson, C.P. Lin, J.S. Schuman, W.G. Stinson, W. Chang, M.R. Hee,

T. Flotte, K. Gregory, C.A. Puliafito, and J.G. Fujimoto, "Optical coherence tomography," Science 254(5035): 1178-1181 (1991).

[2] G.J. Tearney, M.E. Brezinski, B.E. Bouma, S.A. Boppart, C. Pitvis, J.F. Southern, and J.G. Fujimoto, "In vivo endoscopic optical biopsy with optical coherence tomography," Science 276(5321): 2037-2039 (1997).


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[3] S.A. Boppart, B.E. Bouma, C. Pitris, J.F. Southern, M.E. Brezinski, and J.G. Fujimoto, "In vivo cellular optical coherence tomography imaging," Nature medicine 4(7): 861-5 (1998).

[4] W. Drexler, U. Morgner, F.X. Kartner, C. Pitris, S.A. Boppart, X.D. Li, E.P. Ippen, and J.G. Fujimoto, "In vivo ultrahigh-resolution optical coherence tomography," Optics Letters 24(17): 1221-1223 (1999).

[5] U. Morgner, F.X. Kartner, S.H. Cho, Y. Chen, H.A. Haus, J.G. Fujimoto, E.P. Ippen, V. Scheuer, G. Angelow, and T. Tschudi, "Sub-two-cycle pulses from a Kerr-lens mode-locked Ti:sapphire laser," Optics Letters 24(6): 411-413 (1999).

[6] W. Drexler, U. Morgner, R.K. Ghanta, F.X. Kärtner, J.S. Schuman, and J.G. Fujimoto, "Ultrahigh-resolution ophthalmic optical coherence tomography," Nature Medicine 7(4): 502-507 (2001).

[7] W. Drexler, H. Sattmann, B. Hermann, T.H. Ko, M. Stur, A. Unterhuber, C. Scholda, O. Findl, M. Wirtitsch, J.G. Fujimoto, and A.F. Fercher, "Enhanced visualization of macular pathology with the use of ultrahigh-resolution optical coherence tomography," Archives of Ophthalmology 121(5): 695-706 (2003).

[8] T.H. Ko, J.G. Fujimoto, J.S. Duker, L.A. Paunescu, W. Drexler, C.R. Baumal, C.A. Puliafito, E. Reichel, A.H. Rogers, and J.S. Schuman, "Comparison of ultrahigh- and standard-resolution optical coherence tomography for imaging macular hole pathology and repair," Ophthalmology 111(11): 2033-2043 (2004).

[9] T.H. Ko, J.G. Fujimoto, J.S. Schuman, L.A. Paunescu, A.M. Kowalevicz, I. Hartl, W. Drexler, G. Wollstein, J.S. Duker, and H. Ishikawa, "Comparison of ultrahigh and standard resolution optical coherence tomography for imaging of macular pathology," Arch. Ophthalmol. (in press).

[10] G. Wollstein, L.A. Paunescu, T.H. Ko, J.G. Fujimoto, A. Kowalevicz, I. Hartl, S. Beaton, H. Ishikawa, C. Mattox, O. Singh, J. Duker, W. Drexler, and J.S. Schuman, "Ultrahigh-resolution optical coherence tomography in glaucoma," Ophthalmology 112(2): 229-37 (2005).

[11] T.H. Ko, D.C. Adler, J.G. Fujimoto, D. Mamedov, V. Prokhorov, V. Shidlovski, and S. Yakubovich, "Ultrahigh resolution optical coherence tomography imaging with a broadband superluminescent diode light source," Optics Express 12(10): 2112-2119 (2004).

[12] E. Ergun, B. Hermann, M. Wirtitsch, A. Unterhuber, T.H. Ko, H. Sattmann, C. Scholda, J.G. Fujimoto, M. Stur, and W. Drexler, "Assessment of Central Visual Function in Stargardt's Disease/Fundus Flavimaculatus with Ultrahigh-Resolution Optical Coherence Tomography," Invest Ophthalmol Vis Sci 46(1): 310-6 (2005).

[13] P.R. Herz, Y. Chen, A.D. Aguirre, J.C. Fujimoto, H. Mashimo, J. Schmitt, A. Koski, J. Goodnow, and C. Petersen, "Ultrahigh resolution optical biopsy with endoscopic optical coherence tomography," Optics Express 12(15) (2004).

[14] D.E. Spence, P.N. Kean, and W. Sibbett, "60-Fsec Pulse Generation from a Self-Mode-Locked Ti-Sapphire Laser," Optics Letters 16(1): 42-44 (1991).

[15] J.P. Zhou, G. Taft, C.P. Huang, M.M. Murnane, H.C. Kapteyn, and I.P. Christov, "Pulse Evolution in a Broad-Bandwidth Ti-Sapphire Laser," Optics Letters 19(15): 1149-1151 (1994).

[16] A.M. Kowalevicz, T.R. Schibli, F.X. Kartner, and J.G. Fujimoto, "Ultralow-threshold Kerr-lens mode-locked Ti:Al2O3 laser," Optics Letters 27(22): 2037-2039 (2002).

[17] B. Cense, N.A. Nassif, T.C. Chen, M.C. Pierce, S.-H. Yun, B.H. Park, B.E. Bouma, G.J. Tearney, and J.F. de Boer, "Ultrahigh-resolution high-speed retinal imaging using spectral-domain optical coherence tomography," Optics Express 12(11) (2004).

[18] A.T. Semenov, V.R. Shidlovski, D.A. Jackson, R. Willsch, and W. Ecke, "Spectral control in multisection AlGaAs SQW superluminescent diodes at 800 nm," Electronics Letters 32(3): 255-256 (1996).

[19] D.S. Mamedov, V.V. Prokhorov, and S.D. Yakubovich, "Superbroadband high-power superluminescent diode emitting at 920 nm," Quantum Electronics 33(6): 471-473 (2003).

[20] B.E. Bouma, G.J. Tearney, I.P. Bilinsky, B. Golubovic, and J.G. Fujimoto, "Self-phase-modulated Kerr-lens mode-locked Cr:forsterite laser source for optical coherence tomography," Optics Letters 21(22): 1839-1841 (1996).



[21] C. Chudoba, J.G. Fujimoto, E.P. Ippen, H.A. Haus, U. Morgner, F.X. Kärtner, V. Scheuer, G. Angelow, and T. Tschudi, "All-solid-state Cr:forsterite laser generating 14 fs pulses at 1.3 mm," Optics Letters 26: 292-294 (2001).

[22] F.X. Kärtner, N. Matuschek, T. Schibli, U. Keller, H.A. Haus, C. Heine, R. Morf, V. Scheuer, M. Tilsch, and T. Tschudi, "Design and fabrication of double-chirped mirrors," Opt. Lett. 22(13): 831-833 (1997).

[23] P.L. Hsiung, Y. Chen, T.H. Ko, J.G. Fujimoto, C.J.S. de Matos, S.V. Popov, J.R. Taylor, and V.P. Gapontsev, "Optical coherence tomography using a continuous-wave, high-power, Raman continuum light source," Optics Express 12(22): 5287-5295 (2004).

[24] K.L. Corwin, N.R. Newbury, J.M. Dudley, S. Coen, S.A. Diddams, B.R. Washburn, K. Weber, and R.S. Windeler, "Fundamental amplitude noise limitations to supercontinuum spectra generated in a microstructured fiber (vol B 77, pg 269, 2003)," Applied Physics B-Lasers and Optics 77(4): 467-468 (2003).

[25] L. Provino, J.M. Dudley, H. Maillotte, N. Grossard, R.S. Windeler, and B.J. Eggleton, "Compact broadband continuum source based on microchip laser pumped microstructured fibre," Electronics Letters 37(9): 558-560 (2001).

[26] A.V. Avdokhin, S.V. Popov, and J.R. Taylor, "Continuous-wave, high-power, in Raman continuum generation holey fibers," Optics Letters 28(15): 1353-1355 (2003).

[27] T.A. Birks, W.J. Wadsworth, and P.S. Russell, "Supercontinuum generation in tapered fibers," Optics Letters 25(19): 1415-1417 (2000).

[28] I. Hartl, X.D. Li, C. Chudoba, R.K. Hganta, T.H. Ko, J.G. Fujimoto, J.K. Ranka, and R.S. Windeler, "Ultrahigh-resolution optical coherence tomography using continuum generation in an air-silica microstructure optical fiber," Optics Letters 26(9): 608-610 (2001).

[29] Y. Wang, Y. Zhao, J.S. Nelson, Z. Chen, and R.S. Windeler, "Ultrahigh-resolution optical coherence tomography by broadband continuum generation from a photonic crystal fiber," Optics Letters 28(3): 182-4 (2003).

[30] B. Povazay, K. Bizheva, A. Unterhuber, B. Hermann, H. Sattmann, A.F. Fercher, W. Drexler, A. Apolonski, W.J. Wadsworth, J.C. Knight, P.S.J. Russell, M. Vetterlein, and E. Scherzer, "Submicrometer axial resolution optical coherence tomography," Optics Letters 27(20): 1800-2 (2002).

[31] S. Bourquin, A.D. Aguirre, I. Hartl, P. Hsiung, T.H. Ko, J.G. Fujimoto, T.A. Birks, W.J. Wadsworth, U. Bunting, and D. Kopf, "Ultrahigh resolution real time OCT imaging using a compact femtosecond Nd : Glass laser and nonlinear fiber," Optics Express 11(24): 3290-3297 (2003).

[32] Y. Wang, J.S. Nelson, Z. Chen, B.J. Reiser, R.S. Chuck, and R.S. Windeler, "Optimal wavelength for ultrahigh-resolution optical coherence tomography," Optics Express 11(12) (2003).

[33] N. Nishizawa, Y. Chen, P. Hsiung, E.P. Ippen, and J.G. Fujimoto, "Real-time, ultrahigh-resolution, optical coherence tomography with an all-fiber, femtosecond fiber laser continuum at 1.5 mu m," Optics Letters 29(24): 2846-2848 (2004).

2. Fourier Domain OCT Imaging 2.1 Introduction Conventional OCT systems perform measurements of the echo time delay of backscattered or backreflected light by using an interferometer with a mechanically scanned optical reference path. [1-3] Measurements of the echo delay and magnitude of light are performed by mechanically scanning the reference path length, so that light echoes with sequentially different delays are detected at different times as this reference path length is scanned. For this reason, these systems are known as “time domain” systems. Recently, dramatic advances in OCT technology have enabled OCT imaging with a ~15 to 50x increase in imaging speed over standard resolution OCT systems and ~100x over conventional ultrahigh resolution OCT systems. These novel


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detection techniques are known as Fourier domain detection because echo time delays of light are measured by Fourier transforming the interference spectrum of the light signal.[4, 5] Different echo time delays of light produce different frequencies of fringes in the interference spectrum. Fourier domain optical coherence tomography (FdOCT) offers significantly improved sensitivity and imaging speed compared to time domain OCT [6-8]. Fourier domain OCT detection can be performed in two ways; spectral OCT using a spectrometer with a multichannel analyzer [4, 5, 9-11] or swept source OCT using a rapidly tunable laser source [1, 12-15] Spectral and swept source OCT are especially promising for ultrahigh resolution imaging because they overcome imaging speed limitations of time domain OCT. Therefore, it is possible to use these techniques to form 3-dimensional maps of the macula and optic disk [16, 17]. This also enables cross-registration of three-dimensional data sets with fundus photographs, for more accurate diagnosis of disease and evaluation of treatment. In addition, Fourier domain OCT has the advantage of providing direct access to the spectral fringe pattern, enabling a wide range of novel applications. Fourier domain OCT can be used for absorption measurement [18], Doppler techniques can be used to image blood flow[19, 20], and the complex Fourier domain signal can be directly measured to double the axial measurement scan range [21, 22]. In addition spectral domain and swept source OCT are especially well suited for numerical dispersion compensation. Numerical dispersion compensation is especially powerful for applications such as ultrahigh resolution retinal imaging, because variations in eye length between different subjects can cause dispersion mismatch and therefore resolution loss. 2.2 Clinical Ophthalmic Imaging by Spectral OCT The application of ultrahigh resolution OCT in ophthalmic imaging has yielded axial resolutions of 3 um in the retina compared to 10 um for standard resolution OCT [23, 24]. Ultrahigh resolution greatly improves the ability to differentiate architectural morphology in the retina and visualize subtle changes associated with early disease. However, because light exposure levels in the eye are limited, the imaging speeds for ultrahigh resolution OCT that can be achieved using standard time domain detection techniques are very limited. Spectral OCT at standard and ultrahigh resolution has been applied for ophthalmic imaging in the anterior eye and retina and has demonstrated superior imaging speeds compared to time domain OCT [10, 11, 16, 17, 25-27]. Using these latest technological advances in the OCT field, we have recently designed and constructed a portable, high-speed ultrahigh resolution spectral OCT system which can achieve axial image resolution of ~3 µm and can perform high-speed video-rate OCT imaging in the ophthalmology clinic. This high-speed UHR-OCT prototype can be operated with either a femtosecond laser or a broadband SLD as imaging light source. Because femtosecond lasers can be difficult to operate in the clinical setting, the broadband SLD provides an easy to operate and relatively inexpensive light source for clinical UHR-OCT imaging. The combination of high image acquisition speed and ultrahigh OCT axial resolution promises to yield significant improvements in the clinical utility of high-speed ultrahigh resolution OCT systems over currently available prototype and commercial OCT systems. The imaging studies were performed using the high-speed, ultrahigh resolution OCT prototype in the ophthalmology clinic of the New England Eye Center of the Tufts-New England Medical Center (NEMC). The study was approved by the IRB committees of both MIT and NEMC and is compliant with the Health Insurance Portability and Accountability Act of 1996. OCT imaging was performed on more than 200 patients with macular diseases, including macular hole, diabetic retinopathy, age-related macular degeneration, epiretinal membrane, retinitis pigmentosa, and central serous chorioretinopathy.



Figure 1. Schematic diagram of the high speed spectral OCT instrument.

Figure 1 shows a schematic of the high-speed, UHR-OCT system using spectral/Fourier domain detection.[11] Light from a broad bandwidth light source is split between the sample and reference arms of an interferometer. Light in the reference arm is attenuated and reflected from a stationary mirror at a fixed delay. Light in the sample arm is directed though two galvanometer-actuated steering mirrors and relay imaged through the pupil onto the retina.[3] The transverse spot size on the retina is estimated to be 20 um, but depends on factors such as aberrations and refractive powers of different eyes. The galvanometer actuated mirrors can scan the OCT beam across the retina in any arbitrary pattern in order to perform cross-sectional imaging. Cross sectional OCT data was acquired by scanning the OCT beam on the retina under computer control. The incident light power on the eye is 750 µW, the same exposure used in commercial ophthalmic OCT systems and consistent with ANSI safety standards. The spectrum of the interferometer output is detected using a spectrometer consisting of a collimating lens, transmission grating, imaging lens, and CCD line scan camera. The interference spectrum data from the camera was transferred to computer system memory where it was rescaled from wavelength to frequency and Fourier transformed to generate axial measurements of the echo delay and magnitude of light from the retina. Examples of clinical imaging by high speed OCT instrument based on Spectral detection are shown in Figures 2,3. Figure 2 shows a case of a 69-year-old man with 20/80 vision in the right eye diagnosed with an epiretinal membrane (ERM) and a full-thickness macular hole upon clinical examination. The standard resolution StratusOCT image shows a full thickness macular hole. An epiretinal membrane is visible, and cystic structures appear in the inner and outer nuclear layers of the parafoveal region. In the high-speed UHR-OCT image, the ERM is more easily distinguished from the retina, and cystic structures are more clearly visualized.

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Figure 2. Macular hole along with epiretinal membrane: fundus photo (left top), cross-sectional image obtained by commercial instrument – Zeiss Stratus OCT (top right), cross-sectional image obtained by high speed spectral OCT instrument. The high quality image consists of 8100 axial scans (transverse pixels) with 1024 points (axial pixels) per scan and was measured in 0.35sec.

Figure 3. Chronic central serous choriorethinopathy: fundus photo (a), fluorescein angiogram (b), set of high quality cross-sectional images obtained by spectral OCT instrument (1-5). The images consist of 2048 axial scans (transverse pixels) with 1024 points (axial pixels) per scan. Each image was measured in 0.09sec.



In addition, the inner/outer photoreceptor junction (IS/OS) and external limiting membrane (ELM) appear normal adjacent to the hole, indicating photoreceptor integrity. However, they are not readily visualized in the Stratus OCT image. The reduction in signal from the IS/OS junction and ELM in the region of the hole may be attributed to the fact that the retina has lifted away from its normal anatomic position and backscattering is reduced, and does not necessarily indicate photoreceptor impairment. Figure 3 shows the case of a 44-year-old man with chronic central serous choriorethinopathy in the right eye, which had been recurring intermittently for the past several years. High-speed UHR-OCT images show serous fluid accumulation and elevation of the neurosensory retina above an optically clear space. Multiple focal elevations of retinal pigment epithelium (RPE) are seen in the cross-sectional images. The raster set of high pixel density images enables assessing the total volume of serous fluid accumulation. In all images the Bruch’s membrane is visible under the temporal RPE elevation. 2.3 Spectral OCT Retinal Imaging Of Small Animals Retinal imaging with increased sensitivity and speed offered by Fourier domain techniques can serve as valuable clinical tool for the early assessment of ophthalmic disease. It is also possible to track and monitor disease progression and evaluate an efficacy of medical treatment in various retinal pathologies. Systematic studies can be performed on small animal models of various retinal diseases. Recently we have designed and we constructed specialized high speed OCT instrument based on Fourier domain detection for imaging small animals like mice or rats. Figure 4 shows a cross-sectional image of the adult Long-Evans rat retina with 3-4 um axial resolution obtained by spectral OCT system. This figure shows that we are able to visualize all major intraretinal layers in small animals. Retinal imaging of small animals can be also useful as a tool in evaluation of new spectroscopic OCT and functional OCT imaging techniques.

Figure 4. Ultrahigh resolution in vivo OCT image of normal retina in the Long-Evans rat. Ultrahigh resolution OCT enables the layers of the retina to be visualized.

2.4 Imaging by Swept Source OCT Frequency swept light sources are especially important for imaging in the 1300 nm wavelength range, where low cost CCDs are not available. OCT with a frequency swept light source also has the advantage of enabling dual balanced detection and avoiding the need for high performance spectrometers and CCDs which are used in Fourier domain detection. For this reason there is considerable interest in the development of frequency swept light sources for OCT.


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Different laser designs for swept source OCT have been demonstrated. Early studies were performed with bulk optic designs using a Cr:Forsterite laser operating at ~1250 nm and an external cavity diode laser operating at ~800 nm at 10 Hz and 2 kHz sweep rate respectively.[12, 13] An all fiber ring laser employing a fiber Fabry-Perot tunable filter was demonstrated at sweep rates of 200 Hz.[7] Hybrid bulk optic and fiber cavity designs using a fiber coupled semiconductor amplifier as a gain medium and a diffraction grating for wavelength selection have been demonstrated. Frequency sweep rates of 500 Hz (1000 Hz effective sweep rate) and 15.7 kHz have been achieved using a galvanometer scanning mirror and a polygon rotating mirror, respectively. [14, 15]. Techniques for calibrating the instantaneous wavelength of the swept source have been demonstrated.[28, 29] We constructed the OCT system shown in Figure 5. It uses a dual balanced detection scheme for recording the fringe signal of the Michelson interferometer on one channel of a two channel, high speed analog to digital converter (ADC) operating up to 100 Msamples/s with 14 bit resolution. The photodiodes have a bandwidth of 80 MHz, a transimpedance gain of 50,000 V/A and an estimated responsivity of 0.8 A/W at 1300 nm. For real time recalibration of the instantaneous wavelength to frequency, a fraction of the laser output power is split off into a second, fixed fiber Fabry-Perot filter (FFPI) with a FSR of 50 GHz (0.28 nm).

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Figure 5. Top: Schematic of the amplified frequency swept laser source. The laser uses a fiber coupled SOA and a FFP-TF. A SOA boosts the laser output. Bottom: Schematic of the OCT system using Fourier domain / swept source detection.

A 125 MHz bandwidth photodetector records the comb like signal as the laser source sweeps in wavelength. Each peak in the recalibration signal corresponds to a frequency shift of 50 GHz with respect to the previous peak correlating to the transmission maxima of the FFPI. After electronic filtering, the recalibration signal is recorded on the second channel of the ADC synchronously with the sample signal from the OCT interferometer, allowing real time recalibration of the interference signal for equal spacing in frequency prior to the FFT algorithm. OCT imaging is performed by using a beam scanning system probe consisting of a 20 mm focal length fiber collimating lens and a 40 mm focal length objective lens using a 6 mm aperture XY galvanometer scanner which scans the beam post objective.



The application of the frequency swept light source for high speed OCT imaging is demonstrated in Figure 6. The images shown are acquired with axial scan rates of 20 and 27 kHz. The average power on the sample was 1.4 mW and 0.4 mW respectively. The sampling rates was 100 Msamples/s for the images at 20 kHz and 27 kHz line rates. The axial image resolution is 12 to 14 µm in air, corresponding to 9 to 11 µm in tissue. The images consist of 3,500 axial scans (transverse pixels) with 400 points (axial pixels) per scan. The number of axial pixels is determined by the frequency spacing of the recalibration signal and the spectral bandwidth of the source. In this case, the interference signal is zero padded up to a power of 2 to increase the final pixel density to 512 points (1024 points before FFT). In the measurements shown, the DC level was suppressed sufficiently by dual balancing and no additional background subtraction was applied. The images are generated using a postprocessing algorithm which interpolates between the sampling points and finds the maximum more accurately than the real time nearest neighbor check algorithm. A slightly lower background noise was achieved using this algorithm than the real time, “nearest neighbor check”, algorithm. The transverse spot size was ~20 µm.

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Figure 6. OCT images taken with swept source at different sweep rates. Top left: Nailfold imaged at 20 kHz line rate. Top right: Skin in the area of the palm imaged at 20 kHz line rate. Bottom left: Human skin in the area of the finger tip imaged at 20 kHz line rate. Bottom right: Human skin in the area of the finger tip imaged at 27 kHz line rate.

Figure 6 (top, left) shows an OCT image of human skin in vivo in the nailfold region, imaged at an axial scan rate of 20 kHz. Figure 6 (top, right) shows an OCT image of human skin on the palm imaged at 20 kHz. Figure 6 (bottom) shows images of the finger at axial scan rates of 20 kHz (left) and 27 kHz (right). These OCT images are comparable to those obtained using time domain OCT systems and features such as the stratum corneum (SC), the dermal epidermal junction (DEJ), the epidermis (ED), sweat ducts (SD) and the nail (N) can clearly be identified. The images in Figure 8 (middle) show a decrease in imaging quality between 20 kHz and 27 kHz, because of a decrease in laser power and increased ASE (as discussed in the previous section). This high scan rate corresponds to a frequency sweep which is slightly below the single roundtrip limit.


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References [1] D. Huang, E.A. Swanson, C.P. Lin, J.S. Schuman, W.G. Stinson, W. Chang, M.R. Hee, T.


[2] E.A. Swanson, J.A. Izatt, M.R. Hee, D. Huang, C.P. Lin, J.S. Schuman, C.A. Puliafito, and J.G. Fujimoto, "In vivo retinal imaging by optical coherence tomography," Optics Letters 18(21): 1864-1866 (1993).

[3] M.R. Hee, J.A. Izatt, E.A. Swanson, D. Huang, J.S. Schuman, C.P. Lin, C.A. Puliafito, and J.G. Fujimoto, "Optical coherence tomography of the human retina," Archives of Ophthalmology 113(3): 325-332 (1995).

[4] A.F. Fercher, C.K. Hitzenberger, G. Kamp, and S.Y. Elzaiat, "Measurement of Intraocular Distances by Backscattering Spectral Interferometry," Optics Communications 117(1-2): 43-48 (1995).

[5] G. Hausler and M.W. Linduer, ""Coherence radar" and "spectral radar"-new tools for dermatological diagnosis," Journal of Biomedical Optics 3(1): 21-31 (1998).

[6] R. Leitgeb, C.K. Hitzenberger, and A.F. Fercher, "Performance of Fourier domain vs. time domain optical coherence tomography," Optics Express 11(8): 889-894 (2003).

[7] M.A. Choma, M.V. Sarunic, C.H. Yang, and J.A. Izatt, "Sensitivity advantage of swept source and Fourier domain optical coherence tomography," Optics Express 11(18): 2183-2189 (2003).

[8] J.F. de Boer, B. Cense, B.H. Park, M.C. Pierce, G.J. Tearney, and B.E. Bouma, "Improved signal-to-noise ratio in spectral-domain compared with time-domain optical coherence tomography," Optics Letters 28(21): 2067-2069 (2003).

[9] M. Wojtkowski, R. Leitgeb, A. Kowalczyk, T. Bajraszewski, and A.F. Fercher, "In vivo human retinal imaging by Fourier domain optical coherence tomography," Journal of Biomedical Optics 7(3): 457-463 (2002).

[10] M. Wojtkowski, T. Bajraszewski, P. Targowski, and A. Kowalczyk, "Real-time in vivo imaging by high-speed spectral optical coherence tomography," Optics Letters 28(19): 1745-1747 (2003).

[11] M. Wojtkowski, V.J. Srinivasan, T.H. Ko, J.G. Fujimoto, A. Kowalczyk, and J.S. Duker, "Ultrahigh-resolution, high-speed, Fourier domain optical coherence tomography and methods for dispersion compensation," Optics Express 12(11): 2404-2422 (2004).

[12] S.R. Chinn, E.A. Swanson, and J.G. Fujimoto, "Optical coherence tomography using a frequency-tunable optical source," Optics Letters 22(5): 340-2 (1997).

[13] B. Golubovic, B.E. Bouma, G.J. Tearney, and J.G. Fujimoto, "Optical frequency-domain reflectometry using rapid wavelength tuning of a Cr4+:forsterite laser," Optics Letters 22(22): 1704-1706 (1997).

[14] S.H. Yun, C. Boudoux, M.C. Pierce, J.F. de Boer, G.J. Tearney, and B.E. Bouma, "Extended-cavity semiconductor wavelength-swept laser for biomedical imaging," IEEE Photonics Technology Letters 16(1): 293-5 (2004).

[15] S.H. Yun, C. Boudoux, G.J. Tearney, and B.E. Bouma, "High-speed wavelength-swept semiconductor laser with a polygon-scanner-based wavelength filter," Optics Letters 28(20): 1981-1983 (2003).

[16] M. Wojtkowski, T. Bajraszewski, I. Gorczynska, P. Targowski, A. Kowalczyk, W. Wasilewski, and C. Radzewicz, "Ophthalmic imaging by spectral optical coherence tomography," Am J Ophthalmol 138(3): 412-9 (2004).

[17] N.A. Nassif, B. Cense, B.H. Park, M.C. Pierce, S.H. Yun, B.E. Bouma, G.J. Tearney, T.C. Chen, and J.F. de Boer, "In vivo high-resolution video-rate spectral-domain optical coherence tomography of the human retina and optic nerve," Optics Express 12(3) (2004).

[18] R. Leitgeb, M. Wojtkowski, A. Kowalczyk, C.K. Hitzenberger, M. Sticker, and A.F. Fercher, "Spectral measurement of absorption by spectroscopic frequency-domain optical coherence tomography," Optics Letters 25(11): 820-2 (2000).

[19] B.R. White, M.C. Pierce, N. Nassif, B. Cense, B.H. Park, G.J. Tearney, B.E. Bouma, T.C. Chen, and J.F. de Boer, "In vivo dynamic human retinal blood flow imaging using ultra-



high-speed spectral domain optical Doppler tomography," Optics Express 11(25): 3490-3497 (2003).

[20] R.A. Leitgeb, L. Schmetterer, W. Drexler, A.F. Fercher, R.J. Zawadzki, and T. Bajraszewski, "Real-time assessment of retinal blood flow with ultrafast acquisition by color Doppler Fourier domain optical coherence tomography," Optics Express 11(23): 3116-3121 (2003).

[21] M. Wojtkowski, A. Kowalczyk, R. Leitgeb, and A.F. Fercher, "Full range complex spectral optical coherence tomography technique in eye imaging," Optics Letters 27(16): 1415-17 (2002).

[22] P. Targowski, M. Wojtkowski, A. Kowalczyk, T. Bajraszewski, M. Szkulmowski, and I. Gorczynska, "Complex spectral OCT in human eye imaging in vivo," Optics Communications 229(1-6): 79-84 (2004).


[24] T.H. Ko, J.G. Fujimoto, J.S. Duker, L.A. Paunescu, W. Drexler, C.R. Baumal, C.A. Puliafito, E. Reichel, A.H. Rogers, and J.S. Schuman, "Comparison of ultrahigh- and standard-resolution optical coherence tomography for imaging macular hole pathology and repair," Ophthalmology 111(11): 2033-2043 (2004).

[25] N. Nassif, B. Cense, B.H. Park, S.H. Yun, T.C. Chen, B.E. Bouma, G.J. Tearney, and J.F. de Boer, "In vivo human retinal imaging by ultrahigh-speed spectral domain optical coherence tomography," Opt Lett 29(5): 480-2 (2004).


[27] R.A. Leitgeb, W. Drexler, A. Unterhuber, B. Hermann, T. Bajraszewski, T. Le, A. Stingl, and A.F. Fercher, "Ultrahigh resolution Fourier domain optical coherence tomography," Optics Express 12(10): 2156-2165 (2004).

[28] M.A. Choma, C.H. Yang, and J.A. Izatt, "Instantaneous quadrature low-coherence interferometry with 3 x 3 fiber-optic couplers," Optics Letters 28(22): 2162-2164 (2003).

[29] J. Zhang, W.G. Jung, J.S. Nelson, and Z.P. Chen, "Full range polarization-sensitive Fourier domain optical coherence tomography," Optics Express 12(24): 6033-6039 (2004).

3. OCT Technology: MEMS Microscanning Technologies for OCT Sponsors National Institute of Health – NIH-2-R01 EY11289-16, NIH-5-R01-CA75289-06 MFEL F49620-01-1-0186 National Science Foundation – BES-01-19494 Project Staff Aaron D. Aguirre, Paul R. Herz, Dr. Yu Chen, Professor James G. Fujimoto (MIT) Wibool Piyawattanametha, Li Fan, Professsor Ming C. Wu (UC Berkeley) Development of scanning fiber-optic catheters has enabled in vivo endoscopic OCT imaging [7], and results to date have shown the ability to distinguish tissue architectural layers and to differentiate normal from certain pathologic conditions within the human gastrointestinal tract [64-66]. A need for further development of compact, robust scanning devices for endoscopic applications has led to the development of micro-electromechanical systems (MEMS) scanning mirrors for confocal microscopy and for optical coherence tomography [67, 68]. MEMS-based catheters with distal beam scanning can image with higher speed, precision, and repeatability than conventional linear scanning catheters in which the entire catheter is mechanically translated with respect to the tissue. These advantages will enable three dimensional endoscopic OCT


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imaging, which promises to enhance visualization of mucosal architecture and glandular morphology in the upper and lower gastrointestinal tracts. We have developed a two-dimensional MEMS scanning endoscope for OCT imaging. The MEMS scanning mirror is capable of both DC and resonant operation and can produce arbitrary beam scan patterns, enabling both cross-sectional and three-dimensional imaging. The MEMS optical scanner used in the OCT endoscope is shown in Figure 1(a). The mirror is actuated by angular vertical comb (AVC) actuators, which provide larger scan angles for the same comb dimensions compared to standard vertical comb actuators [69]. The device is realized by combining a foundry surface-micromachining process (MUMPS) with a 3-mask deep-reactive-ion-etching (DRIE) post process. Surface micromachining provides versatile mechanical design and electrical interconnect while bulk micromachining offers flat micromirrors and high-force actuators. The scanner utilizes torsion beams and a gimbal-mounting configuration to scan the MEMS mirror on two axes. The DC mechanical scanning angles are illustrated in Figure 1(b). The scanner achieves +/- 5 degrees at 90 Vdc and +/- 5 degrees at 120 Vdc for the inner X and outer Y axes, respectively. In addition, the optical scanner also has excellent frequency response with resonances in the hundreds of Hz for both axes. Resonant operation of the mirror will enable high-speed raster scanning for en face microscopy.

Figure 1. (a) Scanning Electron Micrograph (SEM) of the two-axis MEMS optical scanner. Linear actuation force is generated by angled vertical comb (AVC) actuators. (b) DC scanner characteristics.

The two-axis MEMS scanner is integrated into a compact aluminum housing. Figure 2 compares the conventional OCT catheter (a) with the MEMS scanning device (b). The conventional catheter consists of an optical fiber fused to a graded-index (GRIN) lens, which images the fiber spot to a spot on the tissue. The beam is directed orthogonal to the axis of the catheter by a microprism. The entire device is encased by a transparent plastic sheath and scanning is performed by rotating or translating the device from the proximal end. Typical device outer diameter is < 2 mm. In the MEMS scanning endoscope, light is first collimated by a GRIN fiber collimator fused to the optical fiber. A small achromatic lens focuses light onto tissue. The MEMS optical scanner is mounted at 45 degrees to the optical axis of the lens and redirects the beam out of the endoscope in a similar side-view fashion as used in the conventional catheter. The MEMS scanner performs post-objective scanning, which eliminates the off-axis optical aberrations generated when the beam is scanned on the objective lens. The outer diameter of the prototype MEMS catheter used in this study was ~ 5 mm, and with future improvements in design it can be reduced to less than 3 mm in diameter to be compatible with the working channel of standard endoscopes used in clinical gastrointestinal endoscopy.

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Figure 2. (a) Conventional OCT catheter consisting of an optical fiber, GRIN lens, and microprism encased in a transparent plastic sheath. (b) MEMS scanning OCT endoscope with a fused GRIN fiber collimator, achromatic objective lens, and 2-axis MEMS optical scanner housed in an aluminum package with outer diameter ~ 5 mm.

The MEMS OCT endoscope was incorporated into an ultrahigh resolution OCT imaging system at 1.06 um center wavelength [36]. The system diagram is shown in Figure 3. The system images with axial resolution of < 5 um set by the light source wavelength and bandwidth and transverse resolution of ~ 12 um set by the focusing optics of the MEMS probe. Preliminary results were obtained with the MEMS scanning endoscope in a benchtop imaging experiment. Figure 3 presents sample cross-sectional images of human skin in vivo obtained with the endoscope. Distinction can be made between the stratum corneum (sc), the epidermis (e), and the dermis (d). In addition, sweat duct structure can be visualized in the stratum corneum. The images were acquired at 4 frames per second by scanning the inner mirror while holding the gimbal axis stationary. Images consist of 500 x 1000 pixels (transverse x axial).

Figure 3. Cross-sectional OCT images of human skin in vivo acquired with the MEMS scanning endoscope. The stratum corneum (sc), epidermis (e), and dermis (d) as well as sweat duct structure in the stratum corneum can be visualized.

The two-axis scanning capability of the MEMS endoscope can also be utilized to acquire three-dimensional volume scans. Figure 4 presents an OCT volume scan of lime pulp to demonstrate this concept. To acquire the volume, the inner mirror axis was scanned at 4 Hz while the outer gimbal axis was stepped incrementally after each cross-sectional image was taken. The set of

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cross-sectional images at different transverse locations was then rendered into a three-dimensional reconstruction of the sample. The volume dimensions of the reconstruction are (X, Y, Z) 0.92 mm x 0.2 mm x 1.26 mm.

Figure 4. Three-dimensional OCT reconstruction of lime pulp. Volume dimensions are 0.92 mm x 0.2 mm x 1.26 mm.

Future extension of this work will focus on watertight packaging and further reduction in size to be compatible with standard working channels in clinical endoscopes. Combined with high speed Fourier Domain OCT methods [47, 70], 2D scanning MEMS devices have exciting promise for real-time three-dimensional endoscopic OCT imaging. In addition, using high-speed resonant scanning, these devices also promise to enable endoscopic optical coherence microscopy and multiphoton microscopy [71, 72]. References [1] G. J. Tearney, M. E. Brezinski, B. E. Bouma, S. A. Boppart, C. Pitvis, J. F. Southern, and

J. G. Fujimoto, "In vivo endoscopic optical biopsy with optical coherence tomography," Science, vol. 276, pp. 2037-2039, 1997.

[2] B. E. Bouma, G. J. Tearney, C. C. Compton, and N. S. Nishioka, "High-resolution imaging of the human esophagus and stomach in vivo using optical coherence tomography," Gastrointestinal endoscopy, vol. 51(4) Pt 1, pp. 467-74, 2000.

[3] S. Jäckle, N. Gladkova, F. Feldchtein, A. Terentieva, B. Brand, G. Gelikonov, V. Gelikonov, A. Sergeev, A. Fritscher-Ravens, J. Freund, U. Seitz, S. Schröder, and N. Soehendra, "In vivo endoscopic optical coherence tomography of esophagitis, Barrett's esophagus, and adenocarcinoma of the esophagus," Endoscopy, vol. 32, pp. 750-5, 2000.

[4] M. V. Sivak, Jr., K. Kobayashi, J. A. Izatt, A. M. Rollins, R. Ung-Runyawee, A. Chak, R. C. Wong, G. A. Isenberg, and J. Willis, "High-resolution endoscopic imaging of the GI tract using optical coherence tomography," Gastrointestinal endoscopy, vol. 51(4) Pt 1, pp. 474-9, 2000.

[5] D. L. Dickensheets and G. S. Kino, "Micromachined scanning confocal optical microscope," Optics Letters, vol. 21, pp. 764-6, 1996.

[6] Y. Pan, H. Xie, and G. K. Fedder, "Endoscopic optical coherence tomography based on a microelectromechanical mirror," Optics Letters, vol. 26, pp. 1966-8, 2001.

[7] P. R. Patterson, D. Hah, H. Nguyen, H. Toshiyoshi, R.-m. Chao, and M. C. Wu, "A scanning micromirror with angular comb drive actuation," presented at Technical Digest.



MEMS 2002 IEEE International Conference. Fifteenth IEEE International Conference on Micro Electro Mechanical Systems, 20-24 Jan. 2002, Las Vegas, NV, USA, 2002.

[8] S. Bourquin, A. D. Aguirre, I. Hartl, P. Hsiung, T. H. Ko, J. G. Fujimoto, T. A. Birks, W. J. Wadsworth, U. Bunting, and D. Kopf, "Ultrahigh resolution real time OCT imaging using a compact femtosecond Nd : Glass laser and nonlinear fiber," Optics Express, vol. 11, pp. 3290-3297, 2003.

[9] M. Wojtkowski, T. Bajraszewski, P. Targowski, and A. Kowalczyk, "Real-time in vivo imaging by high-speed spectral optical coherence tomography," Optics Letters, vol. 28, pp. 1745-1747, 2003.

[10] S. H. Yun, G. J. Tearney, J. F. de Boer, N. Iftimia, and B. E. Bouma, "High-speed optical frequency-domain imaging," Optics Express, vol. 11, pp. 2953-2963, 2003.

[11] F. Helmchen, M. S. Fee, D. W. Tank, and W. Denk, "A miniature head-mounted two-photon microscope: High-resolution brain imaging in freely moving animals," Neuron, vol. 31, pp. 903-912, 2001.

[12] A. D. Aguirre, P. Hsiung, T. H. Ko, I. Hartl, and J. G. Fujimoto, "High-resolution optical coherence microscopy for high-speed, in vivo cellular imaging," Optics Letters, vol. 28, pp. 2064-6, 2003.

4. Optical Coherence Microscopy Sponsors National Institute of Health – NIH-5-R01-CA75289-06 MFEL – F49620-01-1-0186 National Science Foundation – BES-0119494 Project Staff Aaron D. Aguirre, Professor James G. Fujimoto Ultrahigh resolution Optical Coherence Tomography (OCT) can achieve 1-2 um axial resolution in tissue but is limited in transverse resolution in order to maintain a sufficient depth of field over the range of the cross-sectional image [9]. Figure 1 illustrates this depth of field restriction on traditional OCT. Low numerical aperture (NA) focusing must be maintained to prevent loss of signal and maintain relatively uniform transverse resolution at large optical delays away from the focus. At high NA, a relatively small confocal range will limit the realizable depth scan in the OCT image.

Figure 1. Depth of field limitation in OCT is determined by the numerical aperture of the focusing optics. High NA focusing results in loss of signal at large optical delays away from the focus. Low NA is used in standard OCT to maintain relatively uniform transverse resolution without signal loss over the range of the axial depth scan.

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The relatively low lateral resolution achievable with cross-sectional OCT is generally insufficient for imaging of cellular features and therefore limits the utility of OCT in applications requiring cellular level diagnostics. Confocal microscopy, on the other hand, provides high lateral and axial resolution to image cellular features, but it can image only relatively surface tissues due to rapid loss of contrast in highly scattering tissue. To extend the imaging power of optical coherence tomography to very high transverse resolution, we are developing a technology known as optical coherence microscopy (OCM), which combines optical coherence tomography with confocal microscopy. Figure 2 compares the image penetration depth and resolution of OCM with that of OCT and confocal microscopy. OCM can provide enhanced penetration depth compared to standard confocal microscopy while dramatically improving the resolution over typical cross-sectional OCT imaging methods.

Figure 2. Transverse resolution and image penetration in optical coherence microscopy (OCM). OCM can dramatically enhance image penetration compared to confocal microscopy alone while significantly improving transverse resolution in OCT to enable cellular level imaging.

Optical coherence microscopy overcomes the depth of field limitation present in traditional OCT imaging by imaging in the en face plane rather than the cross-sectional plane. To image en face, the optical path length of the reference arm is matched exactly to the focus of the sample arm microscope while scanning a transverse raster pattern on the tissue. This eliminates the need for path length scanning to generate an axial depth map and allows the use of high NA lenses to provide very small spot sizes. Figure 3 compares the scan planes for OCT and OCM. Because of the high NA focusing used in OCM, the field of view in the images is necessarily smaller than in OCT. OCT images architectural features over a field of view of 3-6 mm square while OCM can image down to the cellular level but with reduced field of view of 100 – 500 um square. OCM has the unique advantage of using two distinct optical sectioning techniques – confocal gating and coherence gating. The confocal point spread function is entirely determined by the sample arm optics, and in particular, the numerical aperture of the final objective lens. The coherence gate, however, is determined by the light source bandwidth. The degree of confocal rejection of unfocused scattered light can be varied by changing the numerical aperture of the objective lens while the amount of coherence gated sectioning can be varied by changing the bandwidth of the light source. The multiplicative effect of the two sectioning methods strengthens the overall optical sectioning power, allowing increased rejection of unwanted, out of focus scattered light. Studies from our group as well as others have demonstrated that combined confocal and coherence gating can provide improved imaging depth compared to confocal alone [73-75]. The addition of high sensitivity coherence gated detection to confocal detection extends

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the imaging depth in scattering media to the shot noise quantum limit, providing a factor of 2-3 increase over standard confocal microscopy.

Figure 3. Scan planes for OCT and OCM. OCT scans a cross sectional image plane and produces images of millimeter dimensions. OCM scans an en face image plane and achieves very high transverse resolution over smaller fields of view spanning hundreds of micrometers.

The use of multiple optical sectioning techniques also allows considerable flexibility in system design for achieving high-resolution cellular images. Broad bandwidth light sources as used in ultrahigh resolution OCT can provide thin optical sectioning via coherence gating, and the confocal sectioning can be relaxed to facilitate development of miniaturized imaging devices. Figure 4 compares the confocal axial and transverse imaging resolution as a function of the numerical aperture of the probe optics to demonstrate this operating limit for OCM. The axial section thickness degrades much more quickly than the transverse resolution, and there exists a region where the transverse resolution is sufficient for cellular imaging but the axial resolution is not. Addition of a short coherence gate to provide tissue sectioning can therefore make cellular imaging possible with much lower NA than is sufficient for confocal microscopy alone. This operating regime for OCM imaging has very important clinical implications, since it promises to allow cellular imaging with small diameter probes compatible with standard endoscopic and laparoscopic procedures. Recently, we verified the ability to use combination coherence gating and confocal gating to image cellular features when confocal micrsocopy alone would be inadequate [72]. We developed a high speed OCM system capable of imaging with broadband femtosecond laser sources as used in ultrahigh resolution OCT. The system used a modelocked Ti:Sapphire laser and a novel broadband optical phase modulator to provide high resolution imaging at 800 nm wavelength. Figure 5 presents key results. A short coherence gate of ~3 um was combined with a confocal gate of ~30 um. While this confocal section thickness is nearly 6 times that of standard histology, a high transverse resolution of < 2 um can still be maintained. Together the combined gating effects are sufficient for high-resolution in vivo imaging of cellular features in various tissues, demonstrated in Xenopus laevis tadpole and in human skin. In both Xenopus tadpole and in human skin, cellular features were be clearly visualized.


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Figure 4. Numerical aperture requirement for OCM compared to confocal microscopy. OCM can image with high transverse resolution at much lower numerical aperture than confocal microscopy because it does not depend on high axial resolution for optical sectioning.

Demonstration of this operating limit for OCM marks an important advance because it will facilitate the development of miniaturized OCM imaging probes. Elimination of the need for high NA confocal sectioning will significantly relax the design constraints for catheter and rigid endoscope design. Future work will focus on combining OCM technology with the development of miniaturized MEMS scanners for endoscopic microscopy.

Figure 5. OCM can perform cellular imaging with lower NA than confocal microscopy. (Left) A short coherence gate is used to compensate a longer confocal gate, which will relax probe design constraints for endoscopic imaging. (Right) High transverse resolution cellular imaging can be achieved despite weaker confocal sectioning. In vivo OCM cellular imaging was performed in Xenopus laevis tadpole (A) and human skin (B). Nuclei (N), cell membranes (CM) are visible in the tadpole images, and epidermal cells (EC) can be seen in the skin images. Images were acquired at 800 nm wavelength at 4 frames per second.

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References [1] W. Drexler, U. Morgner, F. X. Kartner, C. Pitris, S. A. Boppart, X. D. Li, E. P. Ippen, and J.

G. Fujimoto, "In vivo ultrahigh-resolution optical coherence tomography," Optics Letters, vol. 24, pp. 1221-1223, 1999.

[2] J. A. Izatt, M. R. Hee, G. M. Owen, E. A. Swanson, and J. G. Fujimoto, "Optical coherence microscopy in scattering media," Optics Letters, vol. 19, pp. 590-2, 1994.

[3] J. A. Izatt, M. D. Kulkarni, H.-W. Wang, K. Kobayashi, and M. V. Sivak, Jr., "Optical coherence tomography and microscopy in gastrointestinal tissues," IEEE Journal of Selected Topics in Quantum Electronics, vol. 2, pp. 1017-28, 1996.

[4] M. Kempe, W. Rudolph, and E. Welsch, "Comparative study of confocal and heterodyne microscopy for imaging through scattering media," Journal of the Optical Society of America a-Optics Image Science and Vision, vol. 13, pp. 46-52, 1996.

[5] A. D. Aguirre, P. Hsiung, T. H. Ko, I. Hartl, and J. G. Fujimoto, "High-resolution optical coherence microscopy for high-speed, in vivo cellular imaging," Optics Letters, vol. 28, pp. 2064-6, 2003.

5. Pathology Laboratory Imaging Studies using Ultrahigh Resolution OCT OCT technology is rapidly evolving and performance has improved significantly in the past years. Recent advances in OCT technology have enabled the ultrahigh resolution OCT with the axial resolution < 5 µm. The improved imaging resolution would improve the image quality and therefore potentially reveal more information related to the tissue physiological status. Ex vivo investigation is often a necessary step in the process of validating new technologies and devices for clinical applications. Ex vivo imaging has the advantage of allowing careful control of imaging parameters difficult to perform in vivo, enabling accurate registration of images with histology and allowing feasibility for future in vivo studies to be established. Imaging in the pathology laboratory allows rapid access to fresh surgical specimens, reducing possible artifacts from tissue degradation. The capability of performing ultrahigh resolution imaging in the pathology lab setting enables access to tissues that were previously inaccessible and is especially important as a research tool. The purpose of this study is to provide a basis for interpreting future in vivo endoscopic and laparoscopic OCT studies using ultrahigh resolution imaging. For the pathology lab study, we used two portable broadband light sources. One is a novel compact solid-state Cr4+:Forsterite laser operating at a center wavelength of 1260 nm. Tissue scattering and absorption are minimized by using light in the longer wavelength (1200 to 1300 nm) range, thus permitting OCT imaging to be performed at 1-2 mm depths in most tissues. Using this laser light source, it was possible to achieve an axial resolution of ~4.5 µm in tissue, a factor of two to three times higher than conventional OCT systems previously used to image scattering tissues in vivo. The second system was performed a femtosecond Nd:Glass laser light source at a center wavelength of 1090 nm, yielding a 3.5 µm axial resolution in tissue. This laser is commercially available (High Q Laser Productions, Hohenems, Austria) and is compact, robust, and turn-key, making it suitable for in vivo clinical studies. Although shorter wavelengths are attenuated more rapidly by tissue scattering, they enable finer axial resolutions for a given bandwidth as well as smaller spot sizes. The transverse image resolution was determined by the focused spot size of the optical beam, which is related to the numerical aperture of the focusing lens and the wavelength of the light source, as in conventional microscopy. The transverse spot size diameter for those two OCT systems used in the pathology lab were approximately 11 µm and 6 µm, respectively. Imaging was performed by using a surgical imaging probe (measuring 1 cm in diameter and 15 cm in length) that was mounted on a precision stage to provide controlled displacements of the light beam orientation and position. Imaging was performed at 1-2 frames per second, enabling real-time visualization and the ability to adjust the imaging plane. Imaging


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was performed without contact with the tissue in order to prevent alterations in scattering and brightness due to the pressure of the probe on the tissue. The first phase of this study is to investigate the effects of different tissue preservation solutions on tissue morphology detected by OCT over time. This part of study was performed at MIT. Pathological studies were performed on freshly excised surgical specimens in the pathology laboratory of Beth Israel Deaconess Medical Center. Normal and diseased tissues from surgical resections were imaged within two hours of excision. Since the OCT source light was in the near infrared and invisible to the naked eye, tissue registration was performed with a visible green light guiding beam that was coincident with the OCT imaging beam. Specimens were marked at the beginning and end of each OCT scan with microinjections of ink to designate the orientation of the OCT imaging plane. Tissue specimens were then placed in 10% buffered formalin, routinely processed, and paraffin embedded. Multiple serial tissue sections of 5 µm thickness were obtained and stained with hematoxylin and eosin. Samples were sectioned in the same plane as the OCT imaging plane. Digital two-dimensional OCT images and histology sections corresponding to the imaged areas were compared. Minor discrepancies between histology and OCT imaging can be attributed to tissue fixation, processing, and sectioning artifacts. 5.1 Tissue Preservation Study Sponsors National Institutes of Health – R01-CA75289-05 National Science Foundation – ECS-01-19452 Air Force Office of Scientific Research – F49620-01-1-0186 Project Staff Pei-Lin Hsiung, Dr. Yu Chen, Dr. Prashant R. Nambiar, Professor James G. Fujimoto (MIT) Since the pathology laboratory studies involve the imaging of specimens ex vivo, it is important to understand the time-dependent tissue degradation in various preservation solutions first. The objectives of this study were to investigate the effect of formalin fixation on imaging using ultrahigh resolution OCT, and to test the effect of two commonly used laboratory culture media on ex vivo preservation of tissue optical scattering characteristics over time. Standard formalin fixation followed by paraffin-embedding leads to shrinkage and processing artifacts which complicate OCT image interpretation. Imaging at sequential time-points during the fixation process allows one-to-one correlation of tissue structures with histology, enabling accurate interpretation of in vivo tissue architectural features. As it is not always possible in human studies to image tissue specimens immediately after excision, the effect of post-excision imaging time on optical scattering characteristics was also investigated. Imaging was performed on the same tissue specimens in vivo and up to a maximum time of 10-18 hours after excision, enabling direct comparison of in vivo versus ex vivo imaging. The hamster cheek pouch was chosen for this study because it consists of representative tissue components, is easily accessible, and is a commonly used system for the study of tumorigenesis. When dosed with carcinogen, histological features in this model have been shown to correspond closely with premalignant and malignant lesions in human oral mucosa.[76] The tissue constituents in the hamster cheek pouch are representative of the composition of human epithelial tissues such as oral mucosa, esophagus, and cervix. In this study, three solutions were tested to assess their effects on tissue optical scattering characteristics. First 10% neutral-buffered formalin was investigated to determine the effect of the fixation process on imaging using OCT. Formalin, consisting of 4% formaldehyde, is the most common fixative used for routine paraffin-embedded sections. Formaldehyde acts by forming cross-linkages between tissue proteins, fixing cellular constituents in their in vivo positions and thus preventing autolytic processes. Although formalin penetrates tissue well, the rate of tissue penetration is relatively slow.[77] The second solution tested was isotonic phosphate-buffered



saline (PBS). PBS is a common solution used in many laboratory procedures, often serving as medium for irrigating, transporting and diluting fluid. Isotonic PBS maintains intra- and extra-cellular osmotic balance and a buffering system keeps the medium within the physiological pH range. The third solution tested was Dulbecco’s Modified Eagle’s Media (DMEM), a commonly used synthetic cell culture media. DMEM includes higher concentrations of amino acids and vitamins than basal salt solutions as well as additional supplementary components so that the medium more closely approximates the protein composition of mammalian cells. Autolysis begins immediately after tissue harvest. Tissue culture media was therefore expected to slow the autolytic processes, thereby slowing changes which may influence tissue contrast. Ultrahigh resolution OCT imaging was performed using a state-of-the-art broadband Ti:Al2O3 laser operating at a center wavelength of 800 nm.[9] Imaging was performed with an axial resolution of 2 µm in free space, corresponding to ~1.5 µm in tissue. To achieve high transverse resolution, an imaging probe consisting of two specially-designed achromatic doublets with 10 mm focal length was used, yielding a transverse resolution of ~5 µm in air. The system sensitivity was 103 dB with 3 mW of optical power on the sample. The imaging probe was mounted on two precision computer-controlled micrometer stages to control imaging in two lateral dimensions. All images acquired were 1 mm x 1 mm in dimension and consisted of 1000 axial x 1000 transverse pixels. This enabled an imaging sample density of 1 pixel/µm in both axial and transverse dimensions. Separate animals were used for each solution and temperature investigated in this study in order to enable registered imaging in vivo and ex vivo over the entire time course. For each solution, imaging was performed in vivo, at 15 min, 30 min, 1 hr, 1.5 hours, 2 hours, 2.5 hours, 3 hours, 3.5 hours, 4 hours, 5 hours, 6 hours, 7 hours, 8 hours, and 10 hours. Imaging formalin solution was additionally performed at 12, 14, 16, and 18 hours in order to insure complete fixation. After completion of the imaging time course, the OCT imaging plane was marked with two microinjections of ink, placed in 10% neutral-buffered formalin, routinely processed and paraffin embedded. Standard Hematoxylin-Eosin as well as Masson’s Trichrome stains were used to visualize the imaged tissues for light microscopy. OCT images and corresponding histology sections were compared. Minor discrepancies between histology and OCT images can be attributed to residual tissue fixation, processing and sectioning artifacts. Figure 1 illustrates a sequence of representative OCT images acquired in vivo and over 16 hours in 10% neutral-buffered formalin. During the fixation process, change in contrast was visible within the squamous epithelial layer itself, with an increase in scattering signal in the lower 1/3-1/2 of the epithelium nearer the basal boundary. Decrease in scattering within the lamina propria was apparent, whereas scattering within muscle bundles appeared to increase. Overall contrast between tissue architectural features appeared to increase during fixation over 18 hours and corresponded in time with significantly noticeable tissue shrinkage. Shrinkage was visible in the deeper muscle bundle and connective tissue layers by 1 hour after immersion in formalin and increase in scattering was significantly noticeable by 2 hours. The squamous epithelial layer, loose connective tissue, and muscle layers exhibited approximately 50% overall shrinkage. Little or no shrinkage was observed in adipose tissue, however, boundaries between adipocytes sharpened over the 18 hour time course. Shrinkage and distortion of the surrounding connective tissue is most likely responsible for visible architectural changes in the adipose tissue layers.


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Figure 1. Sequence of representative OCT images acquired in vivo and ex vivo over 16 hours in 10% neutral-buffered formalin: A, In vivo; B, 30 minutes; C, 1 hour; D, 2 hours; E, 4 hours; F, 8 hours; G, 12 hours; H, 16 hours. Shrinkage was visible in all tissue layers. e: epithelium; lp: lamina propria; m1, m2: muscle layers; c: connective tissue; a: adipose tissue; v: blood vessel. Tick marks represent 250 µm. (Wavelength 800 nm, Resolution 1.5 µm axial x 5 µm transverse)

Figure 2 illustrates OCT images acquired in vivo and over 6 hours after excision and immersion in isotonic phosphate-buffered saline at room temperature. Appearance of tissue architectural features in vivo was consistent with the in vivo appearance of the formalin time course specimen discussed previously. However, post-excision optical scattering intensity appeared to decrease monotonically after immersion in PBS. Decrease in overall signal intensity was visible in the image within 30 minutes after tissue excision and was significant by 2 hours. Loss of scattering signal appeared first in the deeper muscle and connective tissue layer (m2) and was consistently noticeable within 3-4 hours. Swelling in the superficial muscle (m1) and thick connective tissue layer (c) was visible. Significant overall change in the epithelial layer was not observed during immersion in isotonic PBS at room temperature over the 10 hour time course, although shrinkage is visible in the histology due to formalin fixation.

Figure 2. Sequence of representative OCT images acquired in vivo and over 6 hours in isotonic phosphate-buffered saline solution at room temperature: A, In vivo; B, 30 minutes; C, 1 hour; D, 2 hours; E, 3 hours; F, 4 hours; G, 6 hours; H, Masson’s trichrome. Loss of scattering signal was significant within 2 hours and appeared to decrease monotonically after tissue excision. Swelling in the muscle (m1 and m2) and thick connective tissue layer (c) was visible. e: epithelium; lp: lamina propria; v: blood vessel. Tick marks represent 250 µm. (Wavelength 800 nm, Resolution 1.5 µm axial x 5 µm transverse)



Figure 3. Sequence of representative OCT images acquired in vivo and over 10 hours in DMEM solution at room temperature: A, In vivo; B, 30 minutes; C, 2 hours; D, 4 hours; E, 6 hours; F, 8 hours; G, 10 hours; H, Masson’s trichrome. Optical scattering characteristics were maintained close to in vivo conditions up to 4-6 hours after excision and remained similar over the entire time course. e: epithelium; lp: lamina propria; m1, m2: muscle layers; c: connective tissue; v: blood vessel. Tick marks represent 250 µm. (Wavelength 800 nm, Resolution 1.5 µm axial x 5 µm transverse)

Figure 3 illustrates OCT images acquired in vivo and over 10 hours after excision and immersion in DMEM at room temperature. Visualization of tissue architectural morphology in vivo was also consistent with in vivo imaging for both formalin and PBS specimens discussed previously. Scattering signal intensity was maintained close to in vivo conditions during immersion in DMEM up to 4-6 hours after excision. Slight decrease of scattering signal visible at the 6-8 hour time point also appeared first in the deep muscle and connective tissue layer (m2), but was not significantly noticeable until 8-10 hours after tissue excision. No significant overall swelling in any of the tissue layers was observed during immersion in DMEM at room temperature. Significant change in the epithelial layer was not observed during immersion in DMEM at room temperature over the 10 hour time course, although shrinkage is again visible in the histology due to formalin fixation. Preservation in DMEM maintained the most similar image appearance and overall signal intensity compared with in vivo imaging. In conclusion, ultrahigh resolution OCT consistently identified the normal keratinized squamous epithelial layer, lamina propria, collagen and muscle bundles, blood vessels, and adipose tissue in vivo and ex vivo. Formalin fixation can result in significant shrinkage of all tissue layers, resulting in tissue architectural distortion. Imaging through the fixation process allowed one-to-one correlation of hamster cheek pouch structures with histology, enabling accurate interpretation of in vivo tissue architectural features. Change in optical scattering was significant within the first 2 hours post-excision after preservation with isotonic PBS. DMEM best preserved in vivo tissue scattering characteristics over the 10 hour imaging time course. While neoplastic lesions were not evaluated in this study, it is expected that similar benefits of imaging using DMEM or other tissue culture media would apply. These results suggest that a tissue culture environment is preferable for preparation of recently excised tissue specimens for optical imaging. These results can also be used as a baseline to aid interpretation of tumor progression studies in the hamster cheek pouch model using optical methods. References [1] D.G. MacDonald, "Comparison of epithelial dysplasia in hamster cheek pouch

carcinogenesis and human oral mucosa," J Oral Pathol, 10: 186-91 (1981). [2] R.D. Start, C.M. Layton, S.S. Cross, and J.H. Smith, "Reassessment of the rate of fixative

diffusion," J Clin Pathol, 45: 1120-1 (1992). [3] W. Drexler, U. Morgner, F.X. Kartner, C. Pitris, S.A. Boppart, X.D. Li, E.P. Ippen, and J.G.

Fujimoto, "In vivo ultrahigh-resolution optical coherence tomography," Optics Letters, 24: 1221-1223 (1999).

5.2. Imaging of Gastrointestinal Pathology


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Sponsors U.S. Army Medical Research Material Command – DAMD 17-01-1-156 National Institutes of Health – R01-CA75289-05 National Science Foundation – ECS-01-19452 Air Force Office of Scientific Research – F49620-01-1-0186 Poduska Family Foundation Fund for Innovative Research in Cancer Philanthropy of Mr. Gerhard Andlinger Project Staff Pei-Lin Hsiung, Aaron D. Aguirre, Dr. Yu Chen, Tony H. Ko, Dr. Stephane Bourquin, Karl Schneider, Professor James G. Fujimoto (MIT); Dr. Liron Pantanowitz, Dr. Darshan Phatak, Professor James L. Connolly (BIDMC) Colorectal cancer (CRC) is a common and lethal disease in the United States. Approximately 145,000 new cases of colorectal cancer are diagnosed each year in the United States, and approximately one in three people who develop CRC die of this disease. CRC is the second leading cause of cancer death, accounting for 10-11% of cancer deaths overall.[78] The majority of colon and rectal cancers are endoluminal adenocarcinomas that arise from the mucosa. Therefore, there is significant interest in the development of diagnostic techniques such as optical coherence tomography that could detect early stage colorectal diseases. OCT has been applied in imaging of lower GI tissues ex vivo and in vivo.[66, 74, 79-87] The purpose of this study is to investigate new ultrahigh resolution OCT technology and to correlate imaging with histology. In this study, normal, inflammatory, and neoplastic tissues of the large and small intestine were imaged ex vivo. Imaging was performed using portable, ultrahigh resolution OCT systems in the clinical pathology laboratory. Ultrahigh resolution OCT imaging at ~1.3 µm wavelength was performed with 4.5 µm axial resolution and 11 µm spot size using a Cr:Forsterite laser light source. This study evaluates improved resolution using the standard OCT imaging wavelength of 1.3 µm. Complementary studies on a smaller set of specimens were performed at ~1.1 µm wavelength with a 3.5 µm axial resolution and 6 µm spot size using a Nd:Glass laser light source.[36] The Nd:Glass laser has the important advantage of being compact and commercially available and is therefore more suitable for future clinical endoscopic imaging studies. This system performs imaging at wavelengths shorter than the wavelength typically used in OCT. The transverse spot size was chosen to be small in order to improve transverse resolution. Finally, three-dimensional OCT imaging was performed to investigate the utility of 3D visualization of tissue architectural morphology. A total of 65 sites from 23 surgical specimens from adult patients were imaged. Both normal and pathologic sites from each specimen were imaged. The majority of specimens were imaged with the 1.3 µm wavelength OCT system. Normal intestinal sites imaged in this first phase of the study included duodenum (n = 2), jejunum (n = 5), and colon (n = 2). Diseased tissue sites included inflammatory bowel disease (n = 9; ulcerative colitis n = 8; Crohn’s disease n = 1), and neoplasia (n = 6), with one tubulovillous adenoma involving the duodenum, one submucosal lipoma, three sites of invasive adenocarcinoma involving the colon, and one squamous cell carcinoma in the anorectal region. A smaller set of specimens were imaged with the 1.1 µm wavelength OCT system. Sites imaged in the second phase of the study included normal colon (n = 11), ulcerative colitis (n = 2), and five sites of adenoma or invasive adenocarcinoma. Specimens with ulcerative colitis of varying severity, ranging from minimal to severe inflammation and ulceration were imaged. Grossly normal as well as ulcerated and scarred lesions (n = 14) were also imaged. Figure 1(A) shows normal colon specimens imaged with OCT. Ultrahigh resolution OCT images of normal colon produced distinct images of the mucosa and submucosa characteristic of normal colonic microstructure. OCT clearly visualized the full thickness of the colon mucosa in almost all



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specimens. The submucosa appeared as a lighter and less optically scattering layer. The muscularis mucosa appeared as a scattering band in the OCT image separating the mucosa and submusoca. Enhanced optical signal intensity beneath individual crypt structures was observed and may be the result of increased transmission of the light through the crypts. Figure 1(B) shows an image of Crohn’s disease of the colon. Abundant irregular branching glands are apparent in the OCT images and histology. The dense inflammatory infiltrate in the lamina propria was again highly scattering. Figure 1(C) shows an OCT image and corresponding histology of adenocarcinoma. OCT image revealed complete loss of normal mucosal architecture and invasion of the submuscosa. Highly scattering and irregular invasive glands were visible in OCT images of adenocarcinoma. OCT images also show areas with cysts and ulceration with overlying fibrinopurulent exudates.

Figure 1. (A) Ultrahigh resolution OCT image of the normal colon (Wavelength 1.26 µm, Resolution 4.5 µm axial x 11 µm transverse). Mucosa (M) is clearly delineated from underlying submucosa (SM) by a scattering band that corresponds to the thin muscularis mucosa (arrows). The submucosa is shown as a loose and less optically scattering layer. Insert: Magnified view of the OCT image. Bottom: Histology, H&E, 40x. (B) Ultrahigh resolution OCT image of chronic colitis due to Crohn’s disease of the colon showing branching irregular glands and increased scattering due to dense inflammatory infiltrate in the lamina propria. Insert: Magnified view of the OCT image. Bottom: Histology, H&E, 40x. (C) Ultrahigh resolution OCT image of well-differentiated adenocarcinoma. Highly irregular invasive glands are visible in a desmoplastic stroma. No clear boundary between mucosa and submucosa is evident in this case. Bottom: Histology, H&E, 40x. All OCT images are taken with central wavelength of 1.26 µm and resolution of 4.5 µm (axial) x 11 µm (transverse).

Figure 2 shows representative 3D OCT volume renderings of the normal colon and a polypoid adenoma of the colon, respectively. Rendered 3D OCT data can be viewed from a virtual surface perspective, yielding a view similar to that of magnification endoscopy. 3D OCT data can also provide a subsurface view. In normal mucosa, crypts were sometimes difficult to identify within individual cross-sectional OCT images. However using 3D OCT, crypts can be visualized based on their shape and distribution in the en face plane. Folds in the epithelium which can appear similar to crypts in an individual cross-sectional OCT image can be readily identified and differentiated from crypts in 3D OCT. Comparison of the 3D OCT rendering of normal colon versus polypoid adenoma shows striking differences in both the surface views as well as the cutaway views. Normal colon exhibits a well organized distribution of crypts which are uniform in


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size and spacing in the en face plane. In contrast, polypoid adenoma exhibits irregular glandular structure.

Figure 2. (A) Volume rendering of normal colon mucosa. Rendered cutaway view showing individual crypts (C) and folds (F) in the en face plane. Folds in the mucosa could be clearly distinguished from crypts (C). Distinctive surface pit patterns in normal mucosa could be seen. (B) Volume rendering of a polypoid adenoma of the colon. Rendered cutaway view shows irregular glandular structures in adenoma (G, arrows) which were distinct from the uniform crypt distribution in normal colon mucosa. The rendered volume is approximately 1 mm x 1.2 mm x 1.3 mm in dimensions. The OCT images are acquired with the ultrahigh resolution system with the central wavelength of 1.09 µm and resolution of 3.5 µm (axial) x 6 µm (transverse).

In conclusion, this study demonstrated ultrahigh resolution OCT imaging of lower GI pathology and correlation with histology. Performing the study in the pathology laboratory setting enabled rapid access to tissue specimens, reducing degradation effects and enabling accurate registration of OCT imaging with histology. Ultrahigh resolution imaging demonstrated clear delineation of the epithelium from the lamina propria and visualization of individual crypts and glandular structure, but individual cells were not visible. Further investigations using ultrahigh resolution imaging in clinical endoscopic studies will be necessary to evaluate the impact of improved resolution. Three-dimensional OCT imaging was demonstrated and enables improved differentiation of en face and three-dimensional structure of glandular organization. These studies can be used as a baseline for development of new OCT technology and could aid the interpretation of future ultrahigh resolution endoscopic imaging studies. References [1] A. Jemal, T. Murray, E. Ward, A. Samuels, R.C. Tiwari, A. Ghafoor, E.J. Feuer, and M.J.

Thun, "Cancer statistics, 2005," CA Cancer J Clin, 55: 10-30 (2005). [2] J.A. Izatt, M.D. Kulkarni, H.-W. Wang, K. Kobayashi, and M.V. Sivak, Jr., "Optical

coherence tomography and microscopy in gastrointestinal tissues," IEEE Journal of Selected Topics in Quantum Electronics, 2: 1017-28 (1996).

[3] G.J. Tearney, M.E. Brezinski, J.F. Southern, B.E. Bouma, S.A. Boppart, and J.G. Fujimoto, "Optical biopsy in human gastrointestinal tissue using optical coherence tomography," The American journal of gastroenterology, 92: 1800-4 (1997).

[4] K. Kobayashi, J.A. Izatt, M.D. Kulkarni, J. Willis, and M.V. Sivak, Jr., "High-resolution cross-sectional imaging of the gastrointestinal tract using optical coherence tomography: preliminary results," Gastrointestinal endoscopy, 47: 515-23 (1998).

[5] C. Pitris, C. Jesser, S.A. Boppart, D. Stamper, M.E. Brezinski, and J.G. Fujimoto, "Feasibility of optical coherence tomography for high-resolution imaging of human gastrointestinal tract malignancies," Journal of gastroenterology, 35: 87-92 (2000).

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[6] M.V. Sivak, Jr., K. Kobayashi, J.A. Izatt, A.M. Rollins, R. Ung-Runyawee, A. Chak, R.C. Wong, G.A. Isenberg, and J. Willis, "High-resolution endoscopic imaging of the GI tract using optical coherence tomography," Gastrointestinal endoscopy, 51(4) Pt 1: 474-9 (2000).

[7] S. Jäckle, N. Gladkova, F. Feldchtein, A. Terentieva, B. Brand, G. Gelikonov, V. Gelikonov, A. Sergeev, A. Fritscher-Ravens, J. Freund, U. Seitz, S. Soehendra, and N. Schrödern, "In vivo endoscopic optical coherence tomography of the human gastrointestinal tract--toward optical biopsy," Endoscopy, 32: 743-9 (2000).

[8] A. Das, M.V. Sivak, Jr., A. Chak, R.C. Wong, V. Westphal, A.M. Rollins, J. Willis, G. Isenberg, and J.A. Izatt, "High-resolution endoscopic imaging of the GI tract: a comparative study of optical coherence tomography versus high-frequency catheter probe EUS," Gastrointestinal endoscopy, 54: 219-24 (2001).

[9] P.R. Pfau, M.V. Sivak, Jr., A. Chak, M. Kinnard, R.C. Wong, G.A. Isenberg, J.A. Izatt, A. Rollins, and V. Westphal, "Criteria for the diagnosis of dysplasia by endoscopic optical coherence tomography," Gastrointest Endosc, 58: 196-202 (2003).

[10] B. Shen, G. Zuccaro, T.L. Gramlich, N. Gladkova, B.A. Lashner, C.P. Delaney, J.T. Connor, F.H. Remzi, M. Kareta, C.L. Bevins, F. Feldchtein, S.A. Strong, M.L. Bambrick, P. Trolli, and V.W. Fazio, "Ex vivo histology-correlated optical coherence tomography in the detection of transmural inflammation in Crohn's disease," Clin Gastroenterol Hepatol, 2: 754-60 (2004).

[11] B. Shen, G. Zuccaro, T.L. Gramlich, N. Gladkova, P. Trolli, M. Kareta, C.P. Delaney, J.T. Connor, B.A. Lashner, C.L. Bevins, F. Feldchtein, F.H. Remzi, M.L. Bambrick, and V.W. Fazio, "In vivo colonoscopic optical coherence tomography for transmural inflammation in inflammatory bowel disease," Clin Gastroenterol Hepatol, 2: 1080-7 (2004).

[12] V. Westphal, A.M. Rollins, J. Willis, M.V. Sivak, and J.A. Izatt, "Correlation of endoscopic optical coherence tomography with histology in the lower-GI tract," Gastrointest Endosc, 61: 537-46 (2005).

[13] S. Bourquin, A.D. Aguirre, I. Hartl, P. Hsiung, T.H. Ko, J.G. Fujimoto, T.A. Birks, W.J. Wadsworth, U. Bunting, and D. Kopf, "Ultrahigh resolution real time OCT imaging using a compact femtosecond Nd : Glass laser and nonlinear fiber," Optics Express, 11: 3290-3297 (2003).

5.3 Imaging of Breast Pathology Sponsors U.S. Army Medical Research Material Command – DAMD 17-01-1-156 National Institutes of Health – R01-CA75289-05 National Science Foundation – ECS-01-19452 Air Force Office of Scientific Research – F49620-01-1-0186 Poduska Family Foundation Fund for Innovative Research in Cancer Philanthropy of Mr. Gerhard Andlinger Project Staff Pei-Lin Hsiung, Dr. Yu Chen, Aaron D. Aguirre, Professor James G. Fujimoto (MIT); Dr. Darshan Phatak, Professor James L. Connolly (BIDMC) Breast cancer is also a common and lethal disease. Approximately 212,930 new cases of invasive breast cancer are expected to occur in the United States in 2005, accounting for 32% of all cancer cases in women. In addition to invasive breast cancer, an estimated 58,490 new cases of in situ cancer are expected to occur in women in 2005.[78] Despite these high incidence rates, breast cancer mortality has declined in the last decade due to both earlier detection and improved treatment methods. Currently 63% of breast cancers are diagnosed at a localized stage, which is attributed to increased use of mammography for screening as well as awareness of breast cancer symptoms in the population. However, a significant percentage of early cancers are still missed,


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suggesting that early detection methods can still be improved. Conventional needle-based biopsy techniques suffer from false-negative rates resulting from sampling limitations. In addition, surgical methods for therapeutic treatment critically depend on accurate methods for assessment of microscopic resection margins. A technique capable of performing high resolution, real time, subsurface imaging like OCT could permit guidance of biopsy and intraoperative monitoring of surgical procedures, offering immediate information to the clinician and likely improving patient outcome. As an emerging medical imaging modality, OCT has been demonstrated to visualize normal and pathologic microstructure within the cardiovascular system, gastrointestinal tract, upper respiratory tract, gynecologic system, urologic tract, and pancreatobiliary tree.[64, 66, 82, 88-94] However, no studies have been reported to date on the ability of OCT to visualize normal and pathologic microstructure in human breast tissue. The purpose of this study was to investigate new ultrahigh resolution OCT technology for imaging normal, benign, and malignant lesions in the human breast. In this study, ultrahigh resolution OCT imaging was performed with 3.5 µm axial resolution and 6 µm spot size at ~1.1 µm wavelength.[36] Ex vivo imaging was performed in order to enable registration to be carefully controlled, allowing accurate correlation of OCT images with histologic micrographs. Three-dimensional OCT imaging was also performed to investigate the utility of 3D visualization for breast architectural morphology. The goal of this study was to determine which features characteristic of normal and pathologic breast tissue can be visualized using OCT. 35 cases from adult patients, consisting of a total of 142 separate specimens, were each imaged within several hours of excision. Fresh specimens were selected based on the presence of pathology upon radiographic findings or gross examination, prompt arrival in the pathology laboratory, and large specimen size, allowing normal and pathologic tissue to be collected for the study without interfering with routine diagnostic procedures. In selected cases, specimens used for clinical diagnostic purposes were imaged immediately prior to placement in cassettes for standard processing. For selected specimens with suspicion of in situ pathology, additional mammography was performed on the excised specimens to guide OCT imaging to areas of microcalcifications which have increased likelihood of pathology. Collected specimens consisted of both normal and pathologic sites from each case. Benign breast tissue imaged consisted of fibroadipose tissue including mammary ducts and lobules (n = 37), fibrocystic changes (n = 8), fibroadenoma (n = 7), ductal hyperplasia (n = 6), and isolated cases of lipoma and neurofibroma. Malignant lesions imaged included ductal carcinoma-in-situ (DCIS, n = 6), infiltrating ductal carcinoma (n = 12), infiltrating lobular carcinoma (n = 8), and infiltrating carcinoma with ductal and lobular features (n = 5). 3 infiltrating ductal carcinoma specimens included regions of DCIS. 9 specimens of DCIS or cancer included microcalcifications. Imaging of normal breast tissue produced distinct images of fibrous stroma and adipose tissue characteristic of normal breast. Figure 1(A) shows ultrahigh resolution OCT images of normal fibroadipose tissue and corresponding histology. Normal fibrous stroma appears heterogeneous and highly scattering and can clearly be distinguished from surrounding adipose tissue. Fibrous stroma exhibited variations in scattering intensity which appeared to correspond to local variations in collagen fiber density visible in histology. Adipocytes appeared round or oval in shape, with individual well-circumscribed borders. Adipose tissue exhibited decreased signal attenuation relative to the surrounding tissue at the same depth. Consequently, regions directly beneath individual adipocytes appear to have higher signal intensity than equivalent regions at the same depth. Figure 1(B) shows an ultrahigh resolution OCT image of a terminal duct lobular unit. Normal glands were visible as round structures up to 100 µm in cross-section. Individual glands appeared less scattering than the surrounding fibrous stroma, and were clearly distinguishable from adipocytes. However, contrast between glandular structures and fibrous stroma was sometimes low, making the size, shape, and extent of individual glands difficult to determine.



Figure 1. (A) OCT image of normal fibroadipose tissue. Fibrous stroma (F) appears heterogeneous and highly scattering while individual adipocytes (A) appear low scattering, with individual well-circumscribed scattering borders. Bottom: Histology, H&E, 40x. (B) OCT images of a terminal duct lobular unit (indicated by the box). Individual glands as elliptical regions of lower scattering within the higher scattering intralobular stroma. A duct is clearly visible to the right of the lobule. Bottom: Histology, H&E, 40x.

Figure 2(A) shows a representative OCT image of fibrocystic changes. Individual cysts within fibrous tissue appeared well-circumscribed and contained scattering material. Epithelium lining cysts was visible as a thin, ~30-40 µm band lining cysts. Figure 2(B) shows representative OCT images of fibroadenoma. Fibroadenoma specimens imaged appeared more uniformly scattering than the heterogeneous fibrous stroma associated with normal fibroadipose tissue. Regions of high scattering suggestive of boundaries between stromal nodules were observed. Compressed ducts with epithelium were visible as finger-like projections of low scattering. OCT images of DCIS within ducts are shown in Figure 3. Tumor cells within a large duct also appeared uniformly lower scattering relative to the surrounding thick collagenous stroma. This study presents the first ultrahigh resolution OCT images of normal and neoplastic lesions of the human breast. OCT imaging of normal, benign, and malignant breast tissue was performed and correlated with histology. Imaging of normal mammary ducts and lobules was demonstrated, and benign changes and infiltrative carcinoma could be visualized. Further investigation using larger numbers of specimens and blinded studies will be required in order to establish sensitivities and specificities to determine the usefulness of OCT for identifying breast cancers and suitability for pursuing in vivo studies. These studies can be used as a baseline for development of new imaging technology for research in breast cancer and could aid the interpretation of future in vivo OCT imaging studies.

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Figure 2. (A) OCT image of benign fibrocystic changes. Individual cysts within fibrous stroma appeared well-circumscribed and irregularly filled with scattering material. Epithelium is occasionally visible as a uniform scattering band lining cysts. Right: Histology, H&E, 40x. (B) OCT image of benign fibroadenoma. Stromal nodules appeared more uniformly scattering than normal fibrous tissue and compressed ducts with epithelial component (arrows) were visible as finger-like regions of low scattering. Right: Histology, H&E, 40x.

Figure 3. (a, b) OCT image of ductal carcinoma-in-situ within a duct in the setting of infiltrating cancer. The duct lumen is completely occluded with low scattering tumor cells. Irregular regions of increased scattering may correspond to collagenous stroma protruding into the duct lumen. Diffusely infilitrating tumor cells are present in the surrounding highly scattering desmoplastic stroma. (c, d) Histology, H&E, 40x and 100x, respectively.

References [1] A. Jemal, T. Murray, E. Ward, A. Samuels, R.C. Tiwari, A. Ghafoor, E.J. Feuer, and M.J.

Thun, "Cancer statistics, 2005," CA Cancer J Clin, 55: 10-30 (2005). [2] M.E. Brezinski, G.J. Tearney, B.E. Bouma, S.A. Boppart, M.R. Hee, E.A. Swanson, J.F.

Southern, and J.G. Fujimoto, "High-Resolution Imaging of Plaque Morphology with Optical Coherence Tomography," Circulation, 92: 103-103 (1995).

[3] C. Pitris, M.E. Brezinski, B.E. Bouma, G.J. Tearney, J.F. Southern, and J.G. Fujimoto, "High resolution imaging of the upper respiratory tract with optical coherence tomography: a feasibility study," American journal of respiratory and critical care medicine, 157(5) Pt 1: 1640-4 (1998).

[4] J.G. Fujimoto, S.A. Boppart, G.J. Tearney, B.E. Bouma, C. Pitris, and M.E. Brezinski, "High resolution in vivo intra-arterial imaging with optical coherence tomography," Heart, 82: 128-33 (1999).



[5] C. Pitris, A. Goodman, S.A. Boppart, J.J. Libus, J.G. Fujimoto, and M.E. Brezinski, "High-resolution imaging of gynecologic neoplasms using optical coherence tomography," Obstetrics and gynecology, 93: 135-9 (1999).

[6] C.A. Jesser, S.A. Boppart, C. Pitris, D.L. Stamper, G.P. Nielsen, M.E. Brezinski, and J.G. Fujimoto, "High resolution imaging of transitional cell carcinoma with optical coherence tomography: feasibility for the evaluation of bladder pathology," British Journal of Radiology, 72: 1170-6 (1999).

[7] B.E. Bouma, G.J. Tearney, C.C. Compton, and N.S. Nishioka, "High-resolution imaging of the human esophagus and stomach in vivo using optical coherence tomography," Gastrointestinal endoscopy, 51(4) Pt 1: 467-74 (2000).

[8] S. Jäckle, N. Gladkova, F. Feldchtein, A. Terentieva, B. Brand, G. Gelikonov, V. Gelikonov, A. Sergeev, A. Fritscher-Ravens, J. Freund, U. Seitz, S. Soehendra, and N. Schrödern, "In vivo endoscopic optical coherence tomography of the human gastrointestinal tract--toward optical biopsy," Endoscopy, 32: 743-9 (2000).


[10] X.D. Li, S.A. Boppart, J. Van Dam, H. Mashimo, M. Mutinga, W. Drexler, M. Klein, C. Pitris, M.L. Krinsky, M.E. Brezinski, and J.G. Fujimoto, "Optical coherence tomography: advanced technology for the endoscopic imaging of Barrett's esophagus," Endoscopy, 32: 921-30 (2000).

[11] U. Seitz, J. Freund, S. Jaeckle, F. Feldchtein, S. Bohnacker, F. Thonke, N. Gladkova, B. Brand, S. Schröder, and N. Soehendra, "First in vivo optical coherence tomography in the human bile duct," Endoscopy, 33: 1018-21 (2001).

[12] S. Bourquin, A.D. Aguirre, I. Hartl, P. Hsiung, T.H. Ko, J.G. Fujimoto, T.A. Birks, W.J. Wadsworth, U. Bunting, and D. Kopf, "Ultrahigh resolution real time OCT imaging using a compact femtosecond Nd : Glass laser and nonlinear fiber," Optics Express, 11: 3290-3297 (2003).

6. Endoscopic OCT Studies OCT enables the visualization of tissue microstructures in vivo and in situ, with resolutions of 1-15 µm, approaching to that of histopathology level. Like conventional biopsy and histology, OCT can provide three-dimensional cross-sectional images that may allow differentiation of normal from diseased tissues. However, unlike biopsy, OCT can be performed in real-time and without the need of tissue excision, thus increasing its possible application areas in where the excision biopsy is impossible or undesirable, and potentially allowing the use of OCT in guidance of biopsy or surgery. We are investigating the use of OCT in imaging a variety of clinically relevant tissue types and pathologies, for application in clinical diagnosis and treatment, and for basic biomedical research. Integration of high resolution OCT into portable systems with various delivering devices offers promise to expand the use of OCT in in vivo and ex vivo imaging in clinic and pathology lab. Endoscopic techniques provide unique approach for evaluating diseases present over the epithelial surface inside human body. Endoscopic OCT imaging holds great promises for early detection, screening, diagnosis, and monitoring the treatment of diseases. 6.1 Endoscopic Imaging in Animals – Enabling Technologies Sponsors


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National Institute of Health – NIH-5-R01-CA75289-06, NIH-2-R01-EY11289-18 National Science Foundation – ECS-0119452, BES-0119494 Air Force Office of Scientific Research – F49620-01-01-0084, F49620-01-01-0186 Project Staff Dr. Yu Chen, Paul Herz, Aaron D. Aguirre, Pei-Lin Hsiung, Professor James G. Fujimoto (MIT) Amanda Koski, Dr. Joseph Schmitt (LightLab Imaging) The development of endoscopic imaging technologies has a profound impact on both fundamental biomedical research and clinical patient care. These technologies include advanced imaging techniques themselves as well as sophisticated diagnostic probes that enable the imaging of internal body organs that would otherwise be inaccessible to various imaging modalities. Such devices are critical for the implementation and realization of OCT as a diagnostic imaging modality, since OCT image penetration is limited to the first several millimeters of a tissue surface. Imaging in animal models allows in vivo and ex vivo imaging to be performed on the same tissue specimen, enabling direct investigation of ex vivo handing protocol on image characteristics. Therefore it would be an important step in terms of evaluation of novel technology as well as understanding of the mechanism of diseases pathogenesis. In this section, we will summarize two types of catheter design, the proximally actuated imaging catheter and the distally actuated imaging catheter, and demonstrate the imaging capabilities on animal model. 1) Proximally Actuated Catheter Probe for Endoscopic Imaging A key technology for OCT imaging of internal organ systems is the catheter/endoscope which is capable of delivering, focusing, scanning, and collecting a single spatial mode optical beam. We have developed, tested, and applied a prototype transverse scanning OCT catheter which is suitable for luminal imaging.[95] The OCT catheter consists of an optical coupling element at its proximal end, a single mode fiber running the length of the catheter, and optical focusing and beam directing elements at the distal end. The catheter is designed to scan the beam in a circumferential or longitudinal pattern in order to cross-sectionally image through the lumen (or other biological structure) into which it is inserted. A single-mode optical fiber lies within a flexible speedometer cable that could be either actuated transversely or rotated with a motor drive unit at the proximal end. The distal end contains a GRIN lens to focus the imaging beam and an angled microprism or micromirror to direct the beam at around 90 degree angle with respect to the axis of the catheter. The entire assembly is housed in a transparent plastic sheath which can be disinfected prior to use and is discarded after use. Ultrahigh resolution endoscopic OCT imaging system using a compact, broadband Cr:Forsterite laser as the light source is used in this study. Prismless operation within the laser cavity was achieved by using double-chirped mirrors for intracavity dispersion compensation. Combined with a dispersion-shifted highly non-linear fiber, a spectral bandwidth of >180 nm at center wavelength of 1250 nm and an output power of 50 mW was generated, yielding an axial resolution of < 4 µm in tissue. The compact Cr:Forsterite laser has a footprint of 40 cm by 60 cm, and is portable to clinical settings. The feasibility of endoscopic OCT imaging was demonstrated on the animal model. The gastrointestinal tracts of New Zealand White rabbit were imaged by using both the linear scanning and the rotary catheters. All imaging procedures were performed at MIT facilities with protocols approved by the MIT Committee on Animal Care (CAC). The rabbit was anesthetized, and the OCT imaging catheter was manually introduced into the upper gastrointestinal (esophagus) and lower gastrointestinal (colon) tracts respectively. Figure 1 illustrates the tip of the 1.8 mm diameter catheter used in the study. Figure 2(a) shows an in vivo OCT image of the rabbit esophagus taken with the linear scanning catheter. The corresponding histology is shown in Figure 2(b). The structure of the esophagus is delineated clearly, with well-organized layers of the squamous epithelium, lamina propria, muscularis



mucosa, submucosa, and inner and outer muscular layers. The tracheal hyaline cartilage (hc) is also visible through the esophageal wall, demonstrating the ability of the endoscopic OCT system to image deeply within the tissue. In addition, the structural details of the tracheal mucosa and trachealis muscle are visible. The corresponding histology section shows very good correlation with the architecture seen in the OCT image. Trichrome staining was used in the histological section to enhance delineation of cartilage rings in the trachea.

Figure 1. Photograph of the tip of catheter designed for endoscopic imaging with the US dime. The diameter is 1.8 mm. The red dot is the aiming beam indicating the imaging region.

In Figure 3, an in vivo OCT image and a corresponding histology of the rabbit colon are shown. The OCT image shows a highly scattering and multistructured layer at the surface that correlates with the colonic mucosa. Individual crypts can be seen in the OCT image and they correlate well with the histology. The crypt features within the mucosa are visualized at greater detail in the 3x magnification images at the right side. Crypt boundaries within the lamina exhibit high scattering intensity in the OCT image, thereby increasing the contrast between individual crypts. This allows the colonic crypt structure to be visualized clearly. The capability to resolve crypt features within the colon is important in the clinical diagnosis of conditions such as inflammatory bowel disease and colon cancer.

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Figure 2. (a) In vivo OCT image and (b) histology of rabbit esophagus and trachea viewed intraluminally from the esophagus. Tracheal hyaline cartilage (hc) between the tracheal mucosa and trachealis muscle is well defined. The image demonstrates the ability of the endoscopic OCT system to image deeply within the tissue. Trichrome stain was used to highlight cartilage and muscle layers.

Figure 3. In vivo endoscopic OCT image of rabbit colon (a) with corresponding histology (b). Delineation of colon upper mucosa (um), muscular mucosa (mm), and muscularis propria (me) is possible. Enlarged images show the capability to visualize crypt structures in finer details.

Figure 4 illustrates a pullback imaging technique that is useful for surveying luminal structures using rotational catheter design. To generate this image, the optical fiber assembly was pulled back within the catheter sheath at a constant rate of 0.5 mm/sec while imaging in a rotational scan mode at 4 rotations per second. Maintaining the external catheter sheath stationary during the pullback scan minimized any motion artifacts or image distortions that would have been

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caused by movement of the outer sheath surface sheath relative to the colonic wall. Figure 4(a) shows one rotational scan in the pullback sequence. The location of this scan is indicated by the solid line in the full pullback transverse cross section (Figure 4(b)). The full cross-sectional image of the colon over a 35 mm scan range was acquired in 70 seconds. The orientation of the cross-sectional imaging plane is seen on the rotational scan image (Figure 4(a)) as the dashed line in the figure. The rotational and cross-sectional pullback images allow simultaneous views of the colon from two aspects. The longitudinal scan format can be valuable in cases where the extent and dimensions of atypical tissue structures must be identified. With the ultrahigh resolution capability of the system, delineation of the colonic mucosa and muscularis was possible. The full rotational scan data sets acquired provided sufficient information for reconstruction of three-dimensional images of the gastrointestinal tract. At this imaging speed, the separation of the successive image planes and corresponding pixel spacing is 125 µm in the pullback direction. With increased imaging speeds, higher pixel density three-dimensional image reconstruction will be possible.

Figure 4. (a) Rotational scan and (b) longitudinal pullback cross-section of rabbit colon in vivo. Architectural detail is distinguished over the large scan field seen in lower image. The location of the rotational image (a) in the pullback sequence is indicated by the solid line in (b). The cross-section imaging plane of the pullback scan is shown as a dashed line in the rotational image (a). The far right area in the pullback cross-section is the animal rectum.

In summary, we have developed a ultrahigh resolution endoscopic OCT imaging system and have obtained the ultrahigh resolution OCT images of the rabbit gastrointestinal tract with minimally invasive catheter devices. Using a compact, broadband Cr:Forsterite laser, we demonstrated endoscopic OCT imaging in vivo at the highest resolution attained to date. Identification of clinically relevant tissue layers was possible at high resolution and is in agreement with histological findings. These results demonstrate the feasibility of using ultrahigh resolution endoscopic OCT for the identification of clinically significant architectural features in the gastrointestinal tract. 2) Distally Actuated Micromotor Catheter Probe for Endoscopic Imaging

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The common catheter probe design for endoscopic OCT imaging uses a mechanical cable, an optical fiber, and a lens assembly housed in a transparent plastic sheath. The cable within the sheath is either rotated or translated in a push-pull motion to produce a rotary or linear scanning motion of the optics in order to generate a transverse or longitudinal OCT image. However, because the distal optics is actuated from the proximal end of the long probe, this design can result in a nonuniform scanning motion that introduces artifacts in the OCT image. Imaging speeds and duty cycles are limited. In addition, for rotary scan designs, a proximally located rotary optical coupling is required. Forward imaging designs avoid these problems, but scan ranges can be limited. Recent developments in novel MEMS scanning technologies promise distal beam scanning, which would improve the speed and reduce image distortion in endoscopic imaging.[68, 96] Despite these advances, previous endoscope-catheter probe designs could not provide focus adjustment during OCT imaging. For this reason, relatively large spot sizes were necessary in order to preserve a sufficient depth of field to enable OCT imaging in intraluminal structures. This limits the transverse image resolution. We have developed a new, high resolution, micromotor endoscope-catheter with adjustable focus capability. The mechanical scanning and microoptic components are located at the distal end of the probe, thereby eliminating the need for proximally actuated rotating or translating elements. Distal actuation provides better uniformity of beam scanning with reduced image distortion artifacts and an improved range of image speeds. Standard OCT imaging probes usually require long depth of fields and, therefore, the minimum transverse focused beam sizes are limited. With the ability to independently adjust the optical beam focal position and use higher numerical aperture optics, transverse image resolution can be improved. In addition, an adjustable focus device can enable C-mode OCT imaging (acquiring images with several focal planes and fusing them together) to overcome depth of field limitations.[9] The micromotor imaging probe also has a larger field of view than conventional OCT endoscope-catheter devices, thereby allowing larger diameter lumens to be more readily visualized.

Figure 5 shows a schematic and photograph of the micromotor endoscope-catheter probe assembly. The probe consists of a distally actuated micromotor and an optical assembly for the beam focusing. The micromotor is a commercially available subsystem (manufactured by Micro Precision Systems AG), which consists of a three-phase brushless DC motor. It has a planetary gearhead design and is configured with a 125:1 gear reduction ratio. A vertical Hall sensor built into the motor is used to accurately control and stabilize the rotation speed with a 5V operating voltage. The optical assembly consists of a GRIN lens collimator and an achromatic focusing lens that move independently of the probe sheath (in a longitudinal direction) in order to adjust the focus position within the tissue. The transverse resolution of the focused optical beam was measured at ~8 µm (2ω0 spot diameter) by using a knife-edge beam profile technique. The corresponding depth of field or confocal parameter is ~80 µm at a 1250 nm center wavelength. An aluminum-coated rod mirror is mounted onto the motor shaft to direct the focused light onto the tissue and to enable a rotational OCT scan. The motor speed can be adjusted, thus enabling rotation speeds from 1 Hz to 100 Hz. The motor and distal optics fit within a 4.8 mm OD stainless-steel housing, which is enclosed within a 5 mm OD transparent plastic sheath. Control wires to actuate the motor are fed through the tubing to the proximal end of the probe. The fiber collimator and focusing lens are attached to a mechanical speedometer cable, which enables the distal focus to be adjusted by translating the fiber and lens assembly with respect to the fixed scanning motor at the distal end.

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Figure 5. Schematic and photograph of the micromotor catheter assembly and components. a) Wireframe cutaway with fiber collimator, focusing lens, rod mirror, and micromotor enclosed in housing. b) Stainless steel housing and catheter components photographed before assembly. The micromotor is shown with the rod lens attached.

The OCT interferometer, which consists of a broadband optical circulator and 90/10 fiber optic beam splitter, transmits 90% of the light into the sample arm and 10% into the reference arm. Dual-channel detection acquired by using polarization diversity is used before digital signal demodulation. A rapid scanning delay line is used in the system reference arm with 2000 axial scans per second. This corresponds to 1000 axial scans per OCT image at a 2 Hz image frame rate. For a probe diameter of 5 mm, this gives a transverse pixel spacing of 15 µm/pixel along the probe circumference. A broadband femtosecond Cr:Forsterite laser and nonlinear optical fiber are used as the light source. The measured axial image resolution in air is 5 µm, which corresponds to a resolution of ~3.7 µm in tissue. The imaging scan depth is 4 mm and the axial pixel spacing is 2 µm/pixel. The sensitivity is 92 dB at an incident power of 12 mW. In order to demonstrate the micromotor endoscope-catheter probe, in vivo OCT imaging was performed on an anesthetized New Zealand white rabbit. Animal handling was performed in accordance with protocols approved by the MIT Committee on Animal Care. Imaging was performed on the colon, because the columnar epithelial structure of colonic mucosa provided well-defined tissue morphology for validation of the adjustable focus probe operation. To minimize animal discomfort and to reduce the risk of damaging the colonic mucosa, a sterile bacteriostatic lubricant was used during catheter insertion. Multiple locations within the colon were imaged. At each imaging position, OCT scans were taken at several focus depths to demonstrate the adjustable focusing capability of the probe. The beam focal depth was changed by translating the collimator and focusing lens assembly with respect to the fixed motor assembly. Figure 6 shows an OCT image obtained with the micromotor probe inserted 3 cm inside the colon. The raw image information was converted to polar coordinate display form. Since the axial scans are radially oriented, a bilinear interpolation was performed at larger radii to account for the wider separation of the axial scans. The probe sheath radius is 2.5 mm, while the OCT scan depth extends an additional 2 mm beyond the sheath. The OCT image was able to clearly delineate the upper mucosa (um), muscularis mucosae (mm), and submucosa (sm) regions within the colon. Adjustment of the optical beam focus in real time was also achieved.

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Figure 6. OCT image of rabbit colon in vivo. Imaging was performed with <4 µm axial resolution. Visualization of colonic upper mucosa (um), muscularis mucosae (mm), and submucosa (sm) layers is clear within the colon with >1 mm image penetration. Shadow regions at the 11 o'clock, 4 o'clock, and 7 o'clock positions indicate areas of housing support struts that occlude approximately 20% of the scan field.

This micromotor imaging probe design has several attractive features. Adjustable focus enables smaller focused spot sizes and a shorter depth of field than what was previously feasible. The resulting improvement in image resolution could improve the performance of ultrahigh resolution OCT imaging by enabling a better visualization of tissue pathology. With additional engineering, it should be possible to perform dynamic focusing, where the focal position is adjusted as a function of rotary beam position. This would enable focus tracking in cases where the endoscope-catheter probe was not centered in the lumen. The use of a micromotor for distal actuation reduces the motion nonuniformity that could occur with previous proximally actuated endoscope catheter probes. A wider range of imaging speeds can be achieved with improved duty cycle. Finally, although this prototype imaging probe was 5 mm in diameter, a reduction in probe size is possible with smaller micromotors.[97] In conclusion, a new micromotor endoscope-catheter imaging probe for OCT has been developed. This device enables focus adjustment and ultrahigh resolution endoscopic imaging. In vivo endoscopic imaging was demonstrated in an animal model and ultrahigh resolution imaging with <4 µm axial resolution was demonstrated. This new design promises to improve imaging performance in future endoscopic OCT imaging studies. References [1] G.J. Tearney, S.A. Boppart, B.E. Bouma, M.E. Brezinski, N.J. Weissman, J.F. Southern,

and J.G. Fujimoto, "Scanning single-mode fiber optic catheter-endoscope for optical coherence tomography," Optics Letters, 21: 543-5 (1996).

[2] Y. Pan, H. Xie, and G.K. Fedder, "Endoscopic optical coherence tomography based on a microelectromechanical mirror," Optics Letters, 26: 1966-8 (2001).

[3] J.M. Zara, S. Yazdanfar, K.D. Rao, J.A. Izatt, and S.W. Smith, "Electrostatic micromachine scanning mirror for optical coherence tomography," Optics letters, 28: 628-30 (2003).



[4] W. Drexler, U. Morgner, F.X. Kartner, C. Pitris, S.A. Boppart, X.D. Li, E.P. Ippen, and J.G. Fujimoto, "In vivo ultrahigh-resolution optical coherence tomography," Optics Letters, 24: 1221-1223 (1999).

[5] P.H. Tran, D.S. Mukai, M. Brenner, and Z. Chen, "In vivo endoscopic optical coherence tomography by use of a rotational microelectromechanical system probe," Opt Lett, 29: 1236-8 (2004).

6.2 Endoscopic Imaging of Barrett's Esophagus and Cancer Sponsors National Institute of Health – R01-CA75289-06, R01-EY11289-18 National Science Foundation – ECS-0119452, BES-0119494 U.S. Air Force Office of Scientific Research – F49620-01-01-0084, F49620-01-01-0186 Cancer Research and Prevention Foundation Project Staff Dr. Yu Chen, Paul Herz, Pei-Lin Hsiung, Aaron D. Aguirre, Professor James G. Fujimoto (MIT); Dr. Hiroshi Mashimo, Dr. Saleem Desai, Dr. Marcos Pedrosa (VA Medical Center); Amanda Koski, Dr. Joseph Schmitt (LightLab Imaging) Optical coherence tomography (OCT) is an emerging biomedical imaging technology which can generate high resolution, cross-sectional images of biological tissues in situ and in real time[1, 2, 98]. OCT can function as a type of optical biopsy to enable imaging of tissue microstructure with resolution approaching that of standard excision biopsy, without the need of excising the tissue specimen[3, 5, 99]. One promising application of optical biopsy using OCT is the endoscopic imaging of gastrointestinal (GI) tract[64, 66]. Gastrointestinal (GI) endoscopy has received increased attention due to the prevalence of esophageal, stomach, and colonic cancers. In the early stages of development, these cancers manifest as cellular and architectural changes in the epithelium or mucosal layers located inside the wall of the gastrointestinal tract. In contrast to conventional video endoscopy, which can only visualize the surface alterations, OCT can detect changes in tissue morphology beneath the tissue surface. Therefore, endoscopic imaging with high resolution OCT could potentially improve the detection, visualization, and diagnosis of gastrointestinal diseases. The first in vivo endoscopic OCT (EOCT) imaging study was performed in a rabbit model, demonstrating the capability of distinguishing architectural layers within the GI tract[7]. Preliminary in vivo EOCT imaging in humans subjects showed the ability to differentiate normal from tumor lesions[100]. Several other studies with EOCT have indicated that it has the potential to distinguish morphological changes associated with clinically significant disease pathologies[64-66, 93, 101]. However, currently almost all clinical studies have been performed using standard OCT with 10-15 µm resolution. Ultrahigh resolution OCT could enhance the imaging performance for the identification of early neoplastic changes, and could improve the sensitivity of biopsy by reducing false negative rates from sampling errors. In order to achieve the high powers and short coherence lengths necessary for high resolution, high speed imaging, a Cr:Forsterite laser is used as the light source[18]. This laser generates short pulses in the 1200 - 1350 nm wavelength regime. The Cr:Forsterite laser has a coherence length of 15 µm, however, because it generates short pulses, the output bandwidth can be increased by using nonlinear effects in optical fibers to yield coherence lengths of 5 µm or less. Figure 1 shows the output spectrum and the corresponding point spread function, in comparison with the standard superluminescent diode (SLD) light source for current clinical OCT systems, which gives about 10-20 µm axial resolution.


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Figure 1. (A) The output spectrum from Cr:Forsterite laser source and compared with standard SLD source. (B) The corresponding point spread function using those two sources.

Figure 2 shows the schematic of the OCT imaging system. Light was coupled into a broadband optical circulator and a 90/10 fiber optic coupler, which transmits 90% of the incident light to the sample arm. A rapid scanning delay line in the reference arm provided real-time imaging up to 3200 axial scans per second. Polarization controllers were used in both the sample and reference arms to optimize signal strength and interference. To match the optical dispersion within the system, dispersion-compensating glass (DCG) was inserted in the reference arm, and an air-gap coupling (AGC) was used in the sample arm. The use of the air coupling allowed precise dispersion compensation and enabled broadband operation for high resolution imaging performance. The OCT signal was split into two orthogonal polarization dependent channels by a polarizing beam splitter (PBS), and the two detector outputs were digitally demodulated. A polarization diversity signal was calculated using the square root of the sum of the squared signal intensities from the two channels. Due to the bandwidth limitations in the optical circulator and other optical components in the sample and reference arms, the back-coupled spectrum on the detector has a bandwidth of 150 nm, which corresponds to a theoretical axial resolution of 4.6 µm in air. The width of the measured axial point spread function is 5 µm, which is a two- to three-fold improvement over standard OCT systems. This corresponds to a ~4 µm resolution in the tissue. With improvements in the bandwidth support of the optical components, higher axial resolutions could be achieved. The system sensitivity was 102 dB at 4 Hz frame rate with up to 15 mW power on the sample.

Figure 2. Schematic of an endoscopic OCT imaging system using a broadband Cr:Forsterite light source. Accurate dispersion matching is achieved through the use of an air-gap coupling (AGC) and dispersion-compensating prisms (DCG), which enables broadband, high resolution operation.

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Figure 3(A) shows the photo of the OCT system in the GI clinic of VA Medical Center. The light was delivered through a 1.8-mm-diameter imaging catheter with a 15 µm focal spot size and a 0.5 mm focal distance from the outer wall of the catheter sheath. The catheter was driven by a magnetic actuator in order to create a linear cross-sectional imaging scan through the tissue surface. A 4 Hz frame rate was implemented in the in vivo imaging to reduce the movement artifacts. In the clinical imaging on human patients, the OCT catheter is inserted through the accessory channel of standard endoscope (Figure 3(B)) and imaging location is identified by the use of visible aiming beam (Figure 3(C)). Pinch biopsy is performed after OCT imaging for histological processing.

Figure 3. (A) Photo of the portable OCT system in the GI clinic. (B) Passage of the catheter through the accessory port of standard endoscope. (C) The endoscopic view of the catheter imaging area inside the human esophagus.

Thirty patients with known Barrett’s esophagus were imaged using the OCT system. Barrett’s esophagus is the replacement of the squamous epithelium of the distal esophagus with columnar epithelium[102]. Several well designed studies have demonstrated that Barrett’s is associated with a 30-40x increased risk of developing adenocarcinoma. Figure 4 shows an OCT image of normal esophagus with squamous epithelium. The OCT image illustrates the relatively homogeneous epithelium (ep), the high-backscattering region (appears brighter) of the lamina propria (lp), The low-backscattering muscularis mucosa (mm), the high-backscattering submucosa (sm), and the low-backscattering and thick muscularis propria (mp).

Figure 4. Ultrahigh resolution OCT image of normal esophagus. Layered structure of the esophagus is well defined with a homogeneous epithelium (ep) above a highly backscattering lamina propria (lp). Deeper layers of the muscularis mucosa (mm), submucosa (sm), and muscularis propria (mp) are also visualized.

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Images were also acquired from regions of Barrett’s epithelium in the same patient. Figure 5 shows an OCT image of an abnormal esophageal region. The OCT image shows clear differences in the tissue architecture of the Barrett's region when compared to squamous epithelium of normal esophagus. The uniformly layered structure has been disrupted by the presence of multiple columnar and gland-like structures and is highly heterogeneous.

Figure 5. Ultrahigh resolution OCT image of Barrett's esophagus. Architectural disruption and glandular features are highly distinct and can be differentiated well form OCT images of normal esophagus.

The ability of OCT to differentiate normal and pathologic tissue has been demonstrated by our group and others [64, 81, 93, 101, 103, 104]. The image resolution of OCT enabled the visualization of architectural morphology features such as the normal layered structure of the epithelium versus glandular and columnar structures associated with Barrett’s esophagus. In the context of surveillance of patients with Barrett’s for high grade dysplasia and adenocarcinoma, the most intriguing application of OCT would be to direct excisional biopsy to reduce sampling errors. One can envision new OCT imaging probes which integrate OCT imaging with pinch biopsy to provide a real time “first look” at pathology prior to excision and processing of a specimen. If successful, this technology could be used to help guide excision biopsy to reduce sampling errors and false negative rates. This could improve sensitivity of diagnosis, reduce the cost of surveillance and provide enhanced diagnosis and treatment decisions. References [1] D. Huang, E.A. Swanson, C.P. Lin, J.S. Schuman, W.G. Stinson, W. Chang, M.R. Hee,

T. Flotte, K. Gregory, C.A. Puliafito, and J.G. Fujimoto, "Optical coherence tomography," Science, 254: 1178-1181 (1991).

[2] J.M. Schmitt, "Optical coherence tomography (OCT): a review," IEEE Journal of Selected Topics in Quantum Electronics, 5: 1205-15 (1999).

[3] J.G. Fujimoto, "Optical coherence tomography for ultrahigh resolution in vivo imaging," Nature Biotechnology, 21: 1361-1367 (2003).

[4] J.G. Fujimoto, M.E. Brezinski, G.J. Tearney, S.A. Boppart, B. Bouma, M.R. Hee, J.F. Southern, and E.A. Swanson, "Optical biopsy and imaging using optical coherence tomography," Nature Medicine, 1: 970-972 (1995).

[5] M.E. Brezinski, G.J. Tearney, B.E. Bouma, J.A. Izatt, M.R. Hee, E.A. Swanson, J.F. Southern, and J.G. Fujimoto, "Optical coherence tomography for optical biopsy. Properties and demonstration of vascular pathology," Circulation, 93: 1206-13 (1996).

[6] J.G. Fujimoto, C. Pitris, S.A. Boppart, and M.E. Brezinski, "Optical coherence tomography: an emerging technology for biomedical imaging and optical biopsy," Neoplasia, 2: 9-25 (2000).

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[8] B.E. Bouma, G.J. Tearney, C.C. Compton, and N.S. Nishioka, "High-resolution imaging of the human esophagus and stomach in vivo using optical coherence tomography," Gastrointestinal endoscopy, 51(4) Pt 1: 467-74 (2000).

[9] G.J. Tearney, M.E. Brezinski, B.E. Bouma, S.A. Boppart, C. Pitvis, J.F. Southern, and J.G. Fujimoto, "In vivo endoscopic optical biopsy with optical coherence tomography," Science, 276: 2037-2039 (1997).

[10] A.M. Sergeev, V.M. Gelikonov, G.V. Gelikonov, F.I. Feldchtein, R.V. Kuranov, N.D. Gladkova, N.M. Shakhova, L.B. Suopova, A.V. Shakhov, I.A. Kuznetzova, A.N. Denisenko, V.V. Pochinko, Y.P. Chumakov, and O.S. Streltzova, "In vivo endoscopic OCT imaging of precancer and cancer states of human mucosa," Optics Express, 1 (1997).

[11] X.D. Li, S.A. Boppart, J. Van Dam, H. Mashimo, M. Mutinga, W. Drexler, M. Klein, C. Pitris, M.L. Krinsky, M.E. Brezinski, and J.G. Fujimoto, "Optical coherence tomography: advanced technology for the endoscopic imaging of Barrett's esophagus," Endoscopy, 32: 921-30 (2000).

[12] G. Zuccaro, N. Gladkova, J. Vargo, F. Feldchtein, E. Zagaynova, D. Conwell, G. Falk, J. Goldblum, J. Dumot, J. Ponsky, G. Gelikonov, B. Davros, E. Donchenko, and J. Richter, "Optical coherence tomography of the esophagus and proximal stomach in health and disease," The American journal of gastroenterology, 96: 2633-9 (2001).

[13] S. Jäckle, N. Gladkova, F. Feldchtein, A. Terentieva, B. Brand, G. Gelikonov, V. Gelikonov, A. Sergeev, A. Fritscher-Ravens, J. Freund, U. Seitz, S. Schröder, and N. Soehendra, "In vivo endoscopic optical coherence tomography of esophagitis, Barrett's esophagus, and adenocarcinoma of the esophagus," Endoscopy, 32: 750-5 (2000).

[14] P.R. Herz, Y. Chen, A.D. Aguirre, J.C. Fujimoto, H. Mashimo, J. Schmitt, A. Koski, J. Goodnow, and C. Petersen, "Ultrahigh resolution optical biopsy with endoscopic optical coherence tomography," Optics Express, 12 (2004).

[15] R.W. Phillips and R.K. Wong, "Barrett's esophagus. Natural history, incidence, etiology, and complications," Gastroenterol Clin North Am, 20: 791-816 (1991).

[16] C. Pitris, C. Jesser, S.A. Boppart, D. Stamper, M.E. Brezinski, and J.G. Fujimoto, "Feasibility of optical coherence tomography for high-resolution imaging of human gastrointestinal tract malignancies," Journal of gastroenterology, 35: 87-92 (2000).

[17] S. Brand, J.M. Poneros, B.E. Bouma, G.J. Tearney, C.C. Compton, and N.S. Nishioka, "Optical coherence tomography in the gastrointestinal tract," Endoscopy, 32: 796-803 (2000).

[18] J.M. Poneros, S. Brand, B.E. Bouma, G.J. Tearney, C.C. Compton, and N.S. Nishioka, "Diagnosis of specialized intestinal metaplasia by optical coherence tomography," Gastroenterology, 120: 7-12 (2001).

7. Functional Brain Imaging with OCT Sponsors Air Force Office of Scientific Research – MFEL F49620-01-1-0186 Project Staff Aaron D. Aguirre, Dr. Yu Chen, Professor James G. Fujimoto (MIT); Professor David Boas, Lana Ruvinskaya, Dr. Anna Devor (Harvard Medical School) Brain neural response to cognitive and sensory stimuli is currently an active area of research. Several imaging techniques have been applied to probe neural response to external stimuli in a noninvasive manner. The most useful techniques of functional brain imaging currently fall into two


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main categories: hemodynamic techniques, such as positron emission tomography (PET), single-photon emission computed tomography (SPECT), and functional magnetic resonance imaging (fMRI), are particularly useful for visualizing the spatial location of neural activity but with limited temporal resolution on the order of 1 second; electrophysiological techniques, such as event related brain potentials (also known as event-triggered EEG, or ERPs) and magnetoencepholography (MEG), are particularly useful for determining the temporal sequencing of neural activity on millisecond timescale but with limited spatial resolution [105]. One of the major challenges in this field is to obtain high spatial resolution maps of neuronal activity on a millisecond time scale to gain a better understanding of neural connectivity and processing networks. In addition, techniques which lead to a better understanding of the vascular response to neuronal activity will enable quantitative neuronal interpretation of hemodynamic measures afforded by high spatial resolution fMRI. Optical techniques in general are uniquely sensitive to neuronal and vascular responses to brain activation, both invasively [106-108] and non-invasively [109], and are currently contributing to unraveling the neuro-vascular relationship [110]. Depth resolution is a limitation of the optical methods primarily used to date. Two-photon microscopy provides depth resolution and is providing access to neuronal and vascular signals during brain activation. Optical coherence tomography can potentially give better depth penetration and complementary signal contrast. Very few studies exist to date using OCT imaging to study functional activation in neuronal tissues. Maheswari et al. showed depth resolved stimulus specific profiles of slow processes during functional activation in the cat visual cortex [111]. They attributed these signals to variation in scattering due to localized structural changes such as capillary dilation and cell swelling. Lazebnik et al. subsequently performed a simple measurement of scattering changes corresponding to fast and slow signals triggered by action potential propagation in the sea slug abdominal ganglion [112]. Other studies have looked at precise measurement of nerve axon displacement using low coherence interferometry methods [113]. We have extended previous work by demonstrating for the first time OCT imaging of scattering changes due to functional activation in the rat somatosensory cortex. This is an important model system for neurophysiology and these results suggest that OCT could have an important role in future studies of the functional neurovascular response. Simultaneous optical coherence tomography and video microscopy imaging were performed using the system shown in Figure 1. The OCT system used here is a standard time domain OCT system with dual-balanced detection and log demodulation. This system operates at 1060 nm center wavelength using an Nd:Glass laser and results in axial resolution of ~18 um in air (13 um in tissue). The beam is positioned on the sample using a XY scanning pair of galvanometers. A single 60 mm focal length lens is used after the collimating lens to provide a lateral spot size of ~40 um. The focusing OCT beam is redirected onto the sample using a hot mirror which allows visible light to pass while reflecting infrared. This configuration has the advantage of allowing easy interface with the existing video microscopy system. The video microscopy technique illuminates the cortex at 570 nm and images scattered light onto a two-dimensional CCD. Changes in cortical reflectivity are measured in response to functional activation. Changes in reflectivity at this wavelength are primarily due to change in the total hemoglobin concentration. The video microscopy method is used in this experiment to localize and characterize the functional activation region for placement of the OCT beam. Figure 1 also shows the stimulation protocol. Forepaw stimulation is performed using 20 second stimulation blocks. Each block consists of 1 second of pre-stimulus period followed by 4 seconds of stimulation with ~1.8 mA pulses at 3 Hz. A 15 second post-stimulus period was then provided to allow full recovery of the excitable tissues to baseline. The stimulus block is repeated 60 times during data acquisition over a 20 min period. Block average signals are then computed to reduce the effects of physiologic noise in the measurements. For data processing, digitized OCT images are converted from log to linear.



Figure 1. Simultaneous OCT and video microscopy system for imaging of functional brain activation.

Experiments were conducted according to protocols approved by the Committee on Animal Care at the Massachusetts Institute of Technology and the Animal Care Committee at the Massachusetts General Hospital. Rats were anesthetized with urethane and immobilized in a sterotaxis before beginning the cranial preparation. An area of skull overlying the primary somatosensory cortex on the contralateral side of the stimulated forepaw was thinned with a dental burr until transparent. Thickness was measured with OCT to be 100-200 um. Sample OCT imaging results are shown in Figure 2. The results shown are averages over 60 repetitions of the 20 second stimulus block protocol. The mean image during the prestimulus period is shown in Figure 2A and provides a structural map of the thinned skull and cortex. The skull (S) appears patchy with a highly scattering layer thought to be the dura separating it from the subarachnoid space and the cortex (C). Large surface cortical vessels (V) can be seen clearly. Figure 2B characterizes the physiologic noise during the 20 min data collection run by plotting the standard deviation divided by mean intensity at every pixel. As expected, the immobilized skull shows very low standard deviation while the cortex shows higher variation. The large surface vessels and the shadowed regions below them show particularly high variation due to flow.

Figure 2. Sample OCT imaging results of functional activation in the rat brain. A) Mean image during prestimulus period. B) Image of standard deviation normalized to the mean. C) OCT functional activation image. D) Functional signal timecourse in region of interest.


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The OCT functional signal is shown in Figure 2C. The signal is computed by dividing each time point in the stimulus and post-stimulus periods by the mean of the pre-stimulus period. The signal is then averaged over all 60 blocks in the data collection run. The time window chosen for plotting Figure 2C corresponds to mean of the stimulus period. Several local regions of interest are chosen from the images for time course analysis. These ROI’s appear as boxes of different colors in the images and the corresponding time courses are plotted in Figure 2D. The time courses seem to stay at baseline during the prestimulus period, begin rising ~ 1 second after the onset of stimulus, and continue to rise until 1-2 seconds after the cessation of the stimulus. The time course returns to baseline after reaching its maximum. Variations seen in the baseline likely result from physiologic variation in baseline and appear exaggerated due to the normalization to the prestimulus mean. The response timecourse varies in amplitude from region to region. For the most part, the shape of the response is uniform with slight variation in time to peak response. The shape of the response is similar to the response observed with the video microscopy suggesting that the OCT signal change may arise from changes in total hemoglobin concentration. Further work is necessary to confirm the origin of this functional signal. References [1] D. Huang, E.A. Swanson, C.P. Lin, J.S. Schuman, W.G. Stinson, W. Chang, M.R. Hee, T.


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[86] B. Shen, G. Zuccaro, T.L. Gramlich, N. Gladkova, P. Trolli, M. Kareta, C.P. Delaney, J.T. Connor, B.A. Lashner, C.L. Bevins, F. Feldchtein, F.H. Remzi, M.L. Bambrick, and V.W. Fazio, "In vivo colonoscopic optical coherence tomography for transmural inflammation in inflammatory bowel disease," Clin Gastroenterol Hepatol 2(12): 1080-7 (2004).

[87] V. Westphal, A.M. Rollins, J. Willis, M.V. Sivak, and J.A. Izatt, "Correlation of endoscopic optical coherence tomography with histology in the lower-GI tract," Gastrointest Endosc 61(4): 537-46 (2005).

[88] M.E. Brezinski, G.J. Tearney, B.E. Bouma, S.A. Boppart, M.R. Hee, E.A. Swanson, J.F. Southern, and J.G. Fujimoto, "High-Resolution Imaging of Plaque Morphology with Optical Coherence Tomography," Circulation 92(8): 103-103 (1995).

[89] C. Pitris, M.E. Brezinski, B.E. Bouma, G.J. Tearney, J.F. Southern, and J.G. Fujimoto, "High resolution imaging of the upper respiratory tract with optical coherence tomography: a feasibility study," American journal of respiratory and critical care medicine 157(5) Pt 1: 1640-4 (1998).

[90] J.G. Fujimoto, S.A. Boppart, G.J. Tearney, B.E. Bouma, C. Pitris, and M.E. Brezinski, "High resolution in vivo intra-arterial imaging with optical coherence tomography," Heart 82(2): 128-33 (1999).

[91] C. Pitris, A. Goodman, S.A. Boppart, J.J. Libus, J.G. Fujimoto, and M.E. Brezinski, "High-resolution imaging of gynecologic neoplasms using optical coherence tomography," Obstetrics and gynecology 93(1): 135-9 (1999).

[92] C.A. Jesser, S.A. Boppart, C. Pitris, D.L. Stamper, G.P. Nielsen, M.E. Brezinski, and J.G. Fujimoto, "High resolution imaging of transitional cell carcinoma with optical coherence tomography: feasibility for the evaluation of bladder pathology," British Journal of Radiology 72(864): 1170-6 (1999).

[93] X.D. Li, S.A. Boppart, J. Van Dam, H. Mashimo, M. Mutinga, W. Drexler, M. Klein, C. Pitris, M.L. Krinsky, M.E. Brezinski, and J.G. Fujimoto, "Optical coherence tomography: advanced technology for the endoscopic imaging of Barrett's esophagus," Endoscopy 32(12): 921-30 (2000).

[94] U. Seitz, J. Freund, S. Jaeckle, F. Feldchtein, S. Bohnacker, F. Thonke, N. Gladkova, B. Brand, S. Schröder, and N. Soehendra, "First in vivo optical coherence tomography in the human bile duct," Endoscopy 33(12): 1018-21 (2001).

[95] G.J. Tearney, S.A. Boppart, B.E. Bouma, M.E. Brezinski, N.J. Weissman, J.F. Southern, and J.G. Fujimoto, "Scanning single-mode fiber optic catheter-endoscope for optical coherence tomography," Optics Letters 21(7): 543-5 (1996).

[96] J.M. Zara, S. Yazdanfar, K.D. Rao, J.A. Izatt, and S.W. Smith, "Electrostatic micromachine scanning mirror for optical coherence tomography," Optics letters 28(8): 628-30 (2003).

[97] P.H. Tran, D.S. Mukai, M. Brenner, and Z. Chen, "In vivo endoscopic optical coherence tomography by use of a rotational microelectromechanical system probe," Opt Lett 29(11): 1236-8 (2004).

[98] J.M. Schmitt, "Optical coherence tomography (OCT): a review," IEEE Journal of Selected Topics in Quantum Electronics 5(4): 1205-15 (1999).

[99] M.E. Brezinski, G.J. Tearney, B.E. Bouma, J.A. Izatt, M.R. Hee, E.A. Swanson, J.F. Southern, and J.G. Fujimoto, "Optical coherence tomography for optical biopsy. Properties and demonstration of vascular pathology," Circulation 93(6): 1206-13 (1996).

[100] A.M. Sergeev, V.M. Gelikonov, G.V. Gelikonov, F.I. Feldchtein, R.V. Kuranov, N.D. Gladkova, N.M. Shakhova, L.B. Suopova, A.V. Shakhov, I.A. Kuznetzova, A.N. Denisenko, V.V. Pochinko, Y.P. Chumakov, and O.S. Streltzova, "In vivo endoscopic OCT imaging of precancer and cancer states of human mucosa," Optics Express 1(13) (1997).

[101] G. Zuccaro, N. Gladkova, J. Vargo, F. Feldchtein, E. Zagaynova, D. Conwell, G. Falk, J. Goldblum, J. Dumot, J. Ponsky, G. Gelikonov, B. Davros, E. Donchenko, and J. Richter, "Optical coherence tomography of the esophagus and proximal stomach in health and disease," The American journal of gastroenterology 96(9): 2633-9 (2001).

[102] R.W. Phillips and R.K. Wong, "Barrett's esophagus. Natural history, incidence, etiology, and complications," Gastroenterol Clin North Am 20(4): 791-816 (1991).


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[103] S. Brand, J.M. Poneros, B.E. Bouma, G.J. Tearney, C.C. Compton, and N.S. Nishioka, "Optical coherence tomography in the gastrointestinal tract," Endoscopy 32(10): 796-803 (2000).

[104] J.M. Poneros, S. Brand, B.E. Bouma, G.J. Tearney, C.C. Compton, and N.S. Nishioka, "Diagnosis of specialized intestinal metaplasia by optical coherence tomography," Gastroenterology 120(1): 7-12 (2001).

[105] G. Gratton and M. Fabiani, "Shedding light on brain function: the event-related optical signal," Trends Cogn Sci 5(8): 357-363 (2001).

[106] H.S. Orbach, L.B. Cohen, and A. Grinvald, "Optical mapping of electrical activity in rat somatosensory and visual cortex," J Neurosci 5(7): 1886-95 (1985).

[107] A. Grinvald, R.D. Frostig, E. Lieke, and R. Hildesheim, "Optical imaging of neuronal activity," Physiol Rev 68(4): 1285-366 (1988).

[108] E.E. Lieke, R.D. Frostig, A. Arieli, D.Y. Ts'o, R. Hildesheim, and A. Grinvald, "Optical imaging of cortical activity: real-time imaging using extrinsic dye-signals and high resolution imaging based on slow intrinsic-signals," Annu Rev Physiol 51: 543-59 (1989).

[109] A. Villringer and B. Chance, "Non-invasive optical spectroscopy and imaging of human brain function," Trends Neurosci 20(10): 435-42 (1997).

[110] A. Devor, I. Ulbert, A.K. Dunn, S.N. Narayanan, S.R. Jones, M.L. Andermann, D.A. Boas, and A.M. Dale, "Coupling of the cortical hemodynamic response to cortical and thalamic neuronal activity," Proc Natl Acad Sci U S A 102(10): 3822-7 (2005).

[111] R.U. Maheswari, H. Takaoka, H. Kadono, R. Homma, and M. Tanifuji, "Novel functional imaging technique from brain surface with optical coherence tomography enabling visualization of depth resolved functional structure in vivo," Journal of neuroscience methods 124(1): 83-92 (2003).

[112] M. Lazebnik, D.L. Marks, K. Potgieter, R. Gillette, and S.A. Boppart, "Functional optical coherence tomography for detecting neural activity through scattering changes," Optics Letters 28(14): 1218-20 (2003).

[113] C. Fang-Yen, M.C. Chu, H.S. Seung, R.R. Dasari, and M.S. Feld, "Noncontact measurement of nerve displacement during action potential with a dual-beam low-coherence interferometer," Optics Letters 29(17): 2028-2030 (2004).

Publications Journal Articles, Published V. Guedes, J.S. Schuman, E. Hertzmark, G. Wollstein, A. Correnti, R. Mancini, D. Lederer, S. Voskanian, L. Velazquez, H.M. Pakter, T. Pedut-Kloizman, J.G. Fujimoto, and C. Mattox, “Optical coherence tomography measurement of macular and nerve fiber layer thickness in normal and glaucomatous human eyes,” Ophthalmology 110, 177-189, January 2003. P.L. Hsiung, X. Li, C. Chudoba, I. Hartl, T.H. Ko, and J.G. Fujimoto, “High-speed path-length scanning with a multiple-pass cavity cell delay line,” Appl. Opt. 42, 640-648, February 2003. S.D. Martin, N.A. Patel, S.B. Adams, Jr., M.J. Roberts, S. Plummer, D.L. Stamper, M.E. Brezinski, and J.G. Fujimoto, “New technology for assessing microstructural components of tendons and ligaments,” Intl. Orthop. 27, 184-189, February 2003. N.A. Patel, X. Li, D.L. Stamper, J.G. Fujimoto, and M.E. Brezinski, “Guidance of aortic ablation using optical coherence tomography,” Intl. J. Cardiovasc. Imaging 19, 171-178, April 2003. J.S. Schuman, G. Wollstein, T. Farra, E. Hertzmark, A. Aydin, J.G. Fujimoto, and L.A. Paunescu, “Comparison of optic nerve head measurements obtained by optical coherence tomography and confocal scanning laser ophthalmoscopy,” Am. J. Ophthalmol. 135, 504-512, April 2003.



W. Drexler, H. Sattmann, B. Hermann, T.H. Ko, M. Stur, A. Unterhuber, C. Scholda, O. Findl, M. Wirtitsch, J.G. Fujimoto, and A.F. Fercher, “Enhanced visualization of macular pathology in the use of ultrahigh-resolution optical coherence tomography,” Arch. Ophthalmol. 121, 695-706, May 2003. D.E. Lederer, J.S. Schuman, E. Hertzmark, J. Heltzer, L.J. Velazques, J.G. Fujimoto, and C. Mattox, “Analysis of macular volume in normal and glaucomatous eyes using optical coherence tomography,” Am. J. Ophthalmol., 135, 838-843, June 2003. A. Ayedin, G. Wollstein, L.L. Price, J.G. Fujimoto, and J.S. Schuman, “Optical coherence tomography assessment of retinal nerve fiber layer thickness changes after glaucoma surgery,” Ophthalmology 110, 1506-1511, August 2003. J.G. Fujimoto, “Optical coherence tomography: Principles and applications,” Rev. of Laser Engin., 31, 635-642, October 2003. M.J. Roberts, S.B. Adams, Jr., N.A. Patel, D.L. Stamper, M.S. Westmore, S.D. Martin, J.G. Fujimoto, and M.E. Brezinski, “A new approach for assessing early osteoarthritis in the rat,” Anal. Bioanal. Chem. 377, 1003-1006, October 2003. A.D. Aguirre, P. Hsiung, T.H. Ko, I. Hartl, and J.G. Fujimoto, “High-resolution optical coherence microscopy for high-speed, in vivo cellular imaging,” Opt. Lett. 28, 2064-2066, November 2003. J.G. Fujimoto, “Optical coherence tomography for ultrahigh resolution in vivo imaging,” Nat. Biotechnol., 21, 1361-1367, November 2003. E. Lai, G. Wollstein, L.L. Price, L.A. Paunescu, P.C. Stark, J.G. Fujimoto, and J.S. Schuman, “Optical coherence tomography disc assessment in optic nerves with peripapillary atrophy,” Ophthalmic Surg. Lasers Imaging 34, 498-504, November/December 2003. S. Bourquin, A.D. Aguirre, I. Hartl, P. Hsiung, T.H. Ko, and J.G. Fujimoto, “Compact broadband light source for ultrahigh resolution optical coherence tomography imaging using a femtosecond Nd:Glass laser and a nonlinear fiber,” Opt. Exp., 11, 3290-3297, December 2003. L. Pantanowitz, P.L. Hsiung, T.H. Ko, K. Schneider, P.R. Herz, J.G. Fujimoto, S. Raza, and J.L. Connolly, “High-resolution imaging of the thyroid gland using optical coherence tomography,” Head and Neck 26: 425-434, April 2004. J. Rogowska, N.A. Patel, J.G. Fujimoto, and M.E. Brezinski, “Optical coherence tomographic elastography technique for measuring deformation and strain of atherosclerotic tissues,” Heart 90: 556-562, May2004. T.H. Ko, D.C. Adler, J.G. Fujimoto, D. Mamedov, V. Prokhorov, V. Shidlovski, and S. Yakubovich, “Ultrahigh resolution optical coherence tomography imaging with broadband superluminescent diode light source,” Optics Express 12: 2112-2119, May 2004. M. Wojtkowski, V.J. Srinivasan, T.H. Ko, J.G. Fujimoto, J.S. Duker, and A. Kowalevicz, “Ultrahigh resolution, high speed, Fourier domain optical coherence tomography and methods for dispersion compensation,” Optics Express12: 2404-2422, May 2004. Journal Articles, Accepted for Publication J.S. Schuman, T. Pedut-Kloizman, E. Hertzmark, M.R. Hee, L. Pieroth, N. Wang, W. C. Kiernan, J.G. Fujimoto, W. Drexler, and R. Ghanta, “Quantitation of nerve fiber layer thickness loss over time in the glaucomatous monkey model using optical coherence tomography,” Arch. Ophthalmol., forthcoming.


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W.P. Roach, C.P. Cain, C.D. DiCarlo, R. Birngruber, J.G. Fujimoto, and C.A. Toth, “The retinal response of macacca mulatta to picosecond laser pulses of varying energy and spot-size,” J. Biomed. Opt., forthcoming. T.H. Ko, J.G. Fujimoto, J.S. Schuman, L.A. Paunescu, A.M. Kowalevicz, I. Hartl, W. Drexler, G. Wollstein, J.S. Duker, and H. Ishikawa, “Comparison of ultrahigh and standard resolution optical coherence tomography for imaging of macular pathology,” Arch. Ophthalmol., forthcoming. T.H. Ko, J.G. Fujimoto, J.S. Duker, L.A. Paunescu, W. Drexler, C.R. Baumal, C. Mattox, E. Reichel, A.H. Rogers, and J.S. Schuman, “Comparison of ultrahigh and standard optical coherence tomography for imaging macular hole pathology and repair,” Ophthalmology, forthcoming. Journal Articles, Submitted for Publication X.D. Li, H. Gold, N. Weissman, K. Saunders, C. Pitris, J.G. Fujimoto, and M.E. Brezinski, “Optical coherence tomography guided stent deployment: a comparison with IVUS in an in vivo rabbit model,” submitted to J. Am. Coll. Cardiol. L.A. Paunescu, J.S. Schuman, L.L. Price, P.C. Stark, S. Beaton, H. Ishikawa, G. Wollstein, and J.G. Fujimoto, “Reproducibility of nerve fiber thickness, macular thickness and optic nerve head measurements using StratusOCT optical coherence tomography,” submitted to Invest. Ophthalmol. and Vis. Sci. N. Patel, J. Zoeller, D.L. Stamper, J.G. Fujimoto, and M.E. Brezinski, “Monitoring osteoarthritis in the rat model using optical coherence tomography, submitted to IEEE Trans. Med. Imaging. P.L. Hsiung, L. Pantanowitz, T.H. Ko, S. Bourquin, J.L. Connolly, and J.G. Fujimoto, “Ultra-high-resolution imaging of lower GI pathology using optical coherence tomography, submitted to Gastrointestinal Endoscopy. E. Ergun, B. Hermann, M. Wirtisch, A. Unterhuber, T. Ko, H. Sattmann, C. Scholda, J.G. Fujimoto, A.F. Fercher, M. Stur, and W. Drexler, “Assessment of central visual function in Stargardt’s Disease/Fundus flavimaculatus with ultrahigh resolution optical coherence tomography,” submitted to Invest. Ophthal. and Vis. Sci. M.G. Wirtitsch, E. Ergun, B. Hermann, A. Unterhuber, M. Stur, C. Scholda, H. Sattmann, T.H. Ko, J.G. Fujimoto, A.F. Fercher, and W. Drexler, “Ultrahigh resolution optical coherence tomography in macular dystrophy,” submitted to Arch. Ophthalmol. X. Li, S. Martin, C. Pitris, R. Ghanta, M. Harman, D.L. Stamper, J.G. Fujimoto, and M.E. Brezinski, “High-resolution optical coherence tomographic imaging of osteoarthritis cartilage during open knee surgery,” submitted to Arthritis Research and Therapy. P.R. Herz, Y. Chen, A. Aguirre, P. Hsiung, K. Schneider, J.G., Fujimoto, H. Mashimo, J. Schmitt, A. Koski, J. Goodnow, and C. Petersen, “High resolution optical biopsy with endoscopic optical coherence tomography,” submitted to Optics Express. P. Hsiung, Y. Chen, T.H. Ko, and J.G. Fujimoto, “Continuous-wave, high-power, Raman continuum light source for optical coherence tomography, submitted to Opt. Lett. A. Chan, J.S. Duker, T.H. Ko, J.G. Fujimoto, and J.S. Schuman, “Normal macular thickness measurements in healthy eyes using Stratus optical coherence tomography (OCT3),” submitted to Arch. Ophthal.



A. Chan, J.S. Duker, J.S. Schuman, and J.G. Fujimoto, “Stage zero macular holes: observations by optical coherence tomography,” submitted to Ophthal. S.G. Schuman, E. Hertzmark, J.G. Fujimoto, and J.S. Schuman, “Wavelength independence and interdevice variability of optical coherence tomography,” submitted to Ophthal. Lasers Surg. and Imaging. Book/Chapters in Books J.G. Fujimoto, M. Brezinski, W. Drexler, I. Hartl, F. Kärtner, X. Li, and M. Morgner, “Optical coherence tomography,” in Ultrafast Lasers, Eds. M.E. Fermann, A. Galvanauskas, and G. Sucha (New York: Marcel Dekker, 2003). J.G. Fujimoto, “Optical coherence tomography,” In Encyclopedia of Optical Engineering, Ed., R. Driggers (New York: Marcel Dekker, 2003). J.G. Fujimoto and M.E. Brezinski, “Optical coherence tomography imaging,” in Biomedical Photonics Handbook, Ed. T. Vo-Dinh, (Boca Raton, Florida: CRC Press, 2003). J.S. Schuman, Z.Y. Williams, J.G. Fujimoto, and L.A. Paunescu, “Imaging in ophthalmology,” in Lasers in Ophthalmology—Basic, Diagnostic and Surgical Aspects, Eds. F. Fankhauser and S. Kwasniewska (The Hague, The Netherlands: Kugler Publications, 2003). J.S Schuman, C.A Puliafito, and J.G. Fujimoto, Optical Coherence Tomography of Ocular Diseases, 2nd edition (Thorofare, New Jersey: SLACK, Inc., 2004). Meeting Papers S. Bourquin, I. Hartl, A. D. Aguirre, P. Hsiung, T. H. Ko, J. G. Fujimoto, T. A. Birks, W. J. Wadsworth, U. Buenting, D. Kopf, “Portable broadband light sources using a femtosecond Nd:Glass laser and nonlinear fiber for ultrahigh resolution OCT imaging,” SPIE International Biomedical Optics Symposium, BIOS’2003, San Jose, California, January 25-31, 2003, paper 4956-03. P. Hsiung, K. Schneider, P. R. Herz, S. Bourquin, T. H. Ko, X. Li, J. G. Fujimoto, M. E. Weinstein, M. S. Hirsch, A. D'Amico, J. Richie, “Ultrahigh resolution OCT of prostate pathology in the clinic using a portable Cr:forsterite laser,” SPIE International Biomedical Optics Symposium, BIOS’2003, San Jose, California, January 25-31, 2003, paper 4956-16. P. R. Herz, S. Martin, P. Hsiung, T. H. Ko, X. Li, N. A. Patel, J. G. Fujimoto, M. E. Brezinski, “Arthroscopic optical coherence tomography for minimally invasive assessment of osteoarthritic cartilage,” SPIE International Biomedical Optics Symposium, BIOS’2003, San Jose, California, January 25-31, 2003, paper 4958-16. W. Drexler, H. Sattmann, B. Hermann, T. H. Ko, A. Unterhuber, M. Stur, C. Scholda, O. Findl, J. G. Fujimoto, and A. F. Fercher, “Evaluation of clinical feasibility of ultrahigh resolution ophthalmic optical coherence tomography,” SPIE International Biomedical Optics Symposium, BIOS’2003, San Jose, California, January 25-31, 2003, paper 4951-32. T.H. Ko, W. Drexler, L.A. Paunescu, I. Hartl, R.K. Ghanta, J.S. Schuman, and J.G. Fujimoto, “Ultrahigh resolution optical coherence tomography for enhanced visualization of retinal pathology, SPIE International Biomedical Optics Symposium, BIOS’2003, San Jose, California, January 25-31, 2003, paper 4656-12.


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L.A. Paunescu, T.H. Ko, W. Drexler, H. Ishikawa, I. Hartl, R.K. Ghanta, J.G. Fujimoto, and J.S. Schuman, “Optic nerve head imaging: Comparison between ultrahigh resolution OCT, standard OCT, and HRT,” SPIE International Biomedical Optics Symposium, BIOS’2003, San Jose, California, January 25-31, 2003, paper 4951-35. J.G. Fujimoto, “Optical coherence tomography: a new method for in vivo imaging,” Keystone Symposia on Molecular and Cellular Biology, Optical Imaging: Applications to Biology and Medicine, Taos, NM, February 11-16, 2003, invited talk. J.G. Fujimoto, “Optical coherence tomography,” SPIE International Symposium on Medical Imaging, San Diego, California, February 14-20, 2003, invited plenary talk. H. Ishikawa, G. Wollstein, L.A. Paunescu, J.G. Fujimoto, and J.S. Schuman, “Evaluation of OCT 3 border detection algorithms, American Glaucoma Society Annual Meeting, San Francisco, California, March 7-9, 2003. J.S. Schuman, T.H. Ko, J.G. Fujimoto, C. Mattox, L. Paunescu, H. Ishikawa, G.Wollstein, W. Drexler, A.M. Kowalevicz, I. Hartl, “Comparative study of glaucoma using ultrahigh resolution and standard optical coherence tomography,” Association for Research in Vision and Ophthalmology, ARVO, Fort Lauderdale, Florida, May 4-9, 2003, paper 976. W. Drexler, B. Hermann, A. Unterhuber, M. Wirtitsch, M. Stur, E. Ergun, E.M. Anger, T. Ko, P.K. Ahnelt, and J.G. Fujimoto, “Investigation of photoreceptor layer impairment in macular pathologies using ultrahigh resolution ophthalmic optical coherence tomography” Association for Research in Vision and Ophthalmology, ARVO, Fort Lauderdale, Florida, May 4-9, 2003, board B511. M.G. Wirtitsch, E. Ergun, B. Hermann, A. Unterhuber, H. Sattmann, M. Stur, T. Ko, O. Findl, J.G. Fujimoto, and W. Drexler, “Ultrahigh resolution optical coherence tomography in macular dystrophy,” Association for Research in Vision and Ophthalmology, ARVO, Fort Lauderdale, Florida, May 4-9, 2003, board B513. L.A. Paunescu, L.L. Price, P.C. Stark, S. Beaton, H. Ishikawa, G. Wollstein, J.G. Fujimoto, and J.S. Schuman, “OCT 3 measurement reproducibility of nerve fiber layer thickness, macula thickness map and optic nerve head analysis on normal subjects,” Association for Research in Vision and Ophthalmology, ARVO, Fort Lauderdale, Florida, May 4-9, 2003, board B12. T.H. Ko, J.S. Schuman, J.G. Fujimoto, L.A. Paunescu, A.M. Kowalevicz, I. Hartl, W. Drexler, J. Duker, G. Wollstein, and H. Ishikawa, “Comparison of ultrahigh resolution and standard optical coherence tomography for imaging retinal pathologies,” Association for Research in Vision and Ophthalmology, ARVO, Fort Lauderdale, Florida, May 4-9, 2003, paper 3277. H. Ishikawa, R. Hariprasad, G. Wollstein, L.A. Paunescu, L.L. Price, P.C. Stark, J.G. Fujimoto, and J.S. Schuman, “New quality assessment parameters for OCT 3,” Association for Research in Vision and Ophthalmology, ARVO, Fort Lauderdale, Florida, May 4-9, 2003, board B61. G. Wollstein, J.S. Schuman, L.L. Price, A. Aydin, L.A. Paunescu, P.C. Stark, J.G. Fujimoto, and H. Ishikawa, “Relationship between optical coherence tomography (OCT) macular and peripapillary measurements and automated visual fields,” Association for Research in Vision and Ophthalmology, ARVO, Fort Lauderdale, Florida, May 4-9, 2003, board B72. C.D. Scholda, M. Wirtitsch, B. Hermann, A. Unterhuber, H. Sattmann, M. Stur, T. Ko, O. Findl, J.G. Fujimoto, and W. Drexler, “Macular hole imaging using ultrahigh resolution optical coherence tomography,” Association for Research in Vision and Ophthalmology, ARVO, Fort Lauderdale, Florida, May 4-9, 2003, board B339.



S. Beaton, H. Ishikawa, G. Wollstein, L.A. Paunescu, J.G. Fujimoto, and J.S. Schuman, “Alternate color scheme for OCT images: A pilot study,” Association for Research in Vision and Ophthalmology, ARVO, Fort Lauderdale, Florida, May 4-9, 2003, board B94. A.J. Correnti, H. Ishikawa, G. Wollstein, L.A. Paunescu, L.L. Price, P.C. Stark, J.G. Fujimoto, and J.S. Schuman, “Evaluation of two OCT 3 border detection algorithms,” Association for Research in Vision and Ophthalmology, ARVO, Fort Lauderdale, Florida, May 4-9, 2003, board B97. E. Ergun, M. Wirtitsch, B.Hermann, A. Unterhuber, H. Sattmann, M. Stur, C. Scholda, J. Fujimoto, A. Fercher, and W. Drexler, “Ultrahigh resolution optical coherence tomography in Stargardt’s disease,” Association for Research in Vision and Ophthalmology, ARVO, Fort Lauderdale, Florida, May 4-9, 2003, board B767. J.G. Fujimoto, T.H. Ko, J.S. Schuman, J. Duker, L.A. Paunescu, A.M. Kowalevicz, I. Hartl, W. Drexler, H. Ishikawa, and G. Wollstein, “Comparative study of macular holes using ultrahigh resolution and standard optical coherence tomography,” Association for Research in Vision and Ophthalmology, ARVO, Fort Lauderdale, Florida, May 4-9, 2003, board B520. S. Bourquin, P. Hsiung, I. Hartl, A. Aguirre, T.H. Ko, J.G. Fujimoto, T.A. Birks, W.J. Wadsworth, U. Buenting, and D. Kopf, “Compact ultrahigh resolution OCT system for real-time in vivo imaging,” Conference on Lasers and Electro-Optics, CLEO, June 1-6, 2003, Baltimore, Maryland, paper CMU1. P.R. Herz, P. Hsiung, S. Bourquin, T.H. Ko, K. Schneider, H. Ishikawa, and J.G. Fujimoto, “Micro-motor endoscope with adjustable focus for ultrahigh resolution OCT,” Conference on Lasers and Electro-Optics, CLEO, June 1-6, 2003, Baltimore, Maryland, paper CMU2. S.B. Adams, Jr., M.J. Roberts, N.A. Patel, S. Plummer, J. Rogowska, D.L. Stamper, J.G. Fujimoto, and M.E. Brezinski, “The use of polarization sensitive optical coherence tomography and elastography to assess connective tissue,” Conference on Lasers and Electro-Optics, CLEO, June 1-6, 2003, Baltimore, Maryland, paper CTuM48. A.D. Aguirre, P. Hsiung, T.H. Ko, S. Bourquin, K. Schneider, I. Hartl, and J.G. Fujimoto, “Alternate scanning modalities for high resolution optical coherence tomography,” Conference on Lasers and Electro-Optics, CLEO, June 1-6, 2003, Baltimore, Maryland, paper CTuM49. P. Hsiung, A.D. Aguirre, T.H. Ko, S. Bourquin, K. Schneider, I. Hartl, and J.G. Fujimoto, “Transverse priority scanning and microscopy for high resolution optical coherence tomography,” European Conference on Biomedical Optics, Munich, Germany, June 22-25, 2003, 5140-20. T.H. Ko, A.M. Kowalevicz, Jr., A. Paunescu, J.S. Schuman, W. Drexler, I. Hartl, F. Kaertner, and J.G. Fujjimoto, “Ultrahigh resolution OCT imaging of retinal pathology in the ophthalmology clinic,” European Conference on Biomedical Optics, Munich, Germany, June 22-25, 2003, 5140-26. B. Hermann, A. Unterhuber, H. Sattmann, M. Stur, M. Wirtisch, M. Glosman, C. Schubert, C. Scholda, O. Findl, T.H. Ko, P. Ahnelt, and J. Fujimoto, “Investigation of photoreceptor layer impairment in macular pathologies, European Conference on Biomedical Optics, Munich, Germany, June 22-25, 2003, 5140-27. P.R. Herz, S. Adams, Jr., J. Roberts, S. Bourquin, P. Hsiung, T. Ko, N. Patel, D. Stamper, S. Martin, K. Schneider, J.G. Fujimoto, and M. Brezinski, “Imaging of cartilage degeneration progression in vivo,” European Conference on Biomedical Optics, Munich, Germany, June 22-25, 2003, 5140-30.


34-69

S. Bourquin, A.D. Aguirre, I. Hartl, P. Hsiung, T.H. Ko, and J.G. Fujimoto, “Real-time ultrahigh resolution OCT systems for in vivo imaging,” European Conference on Biomedical Optics, Munich, Germany, June 22-25, 2003, 5140-31. T. H. Ko, A. Paunescu, I. Hartl, A. Kowalevicz, J. Duker, J. Schuman, W. Drexler, A. Rogers, C. Baumal, E. Reichel, L.P. Aiello, S. Bursell, J. Fujimoto, “Comparing ultrahigh resolution optical coherence tomography and StratusOCT imaging of retinal pathologies,” SPIE International Biomedical Optics Symposium, BIOS’2004, San Jose, California, January 24-29, 2004, paper 5314-22. W. Drexler, B. Hermann, A. Unterhuber, H. Sattmann, T.H. Ko, M. Wirtitsch, M. Stur, C. Scholda, E. Ergun, E. Anger, P. Ahnelt, J.G. Fujimoto, and A.F. Fercher, “Quantification of photoreceptor layer thickness in different macular pathologies using ultrahigh resolution optical coherence tomography,” SPIE International Biomedical Optics Symposium, BIOS’2004, San Jose, California, January 24-29, 2004, paper 5314-27. T.H. Ko, A. Paunescu, A. Kowalevicz, I. Hartl, J. Schuman, J. Duker, W. Drexler, L.P. Aiello, S. Bursell, J. Fujimoto, “Ultrahigh resolution optical coherence tomography for imaging retinal pathology,” SPIE International Biomedical Optics Symposium, BIOS’2004, San Jose, California, January 24-29, 2004, paper 5316-02. J.G. Fujimoto, Y. Chen, D. Adler, A. Aguirre, P. Herz, P. Hsiung, T. Ko, and M. Wojtkowski, “Ultrahigh resolution imaging using optical coherence tomography,” OSA Biomedical Optics Topical Meetings, Miami Beach, Florida, April 14-17, 2004, paper FD1. B.K. Courtney, A.L. Robertson, Jr., F.P. Chan, Y. Honda, P.G. Yock, P.J. Fitzgerald, P.R. Herz, T.H. Ko, J.G. Fujimoto, D.L. Stamper, S.B. Adams, M.J. Roberts, and M.E. Brezinski, “Attentuation measurements of optical coherence tomographic signals for the assessment of vascular tissue composition,” OSA Biomedical Optics Topical Meetings, Miami Beach, Florida, April 14-17, 2004, paper FD3. D.C. Adler, T.H. Ko, J.G. Fujimoto, D. Mamedov, V. Prokhorov, V. Shidlovski, and S. Yakubovich, “Ultrahigh resolution optical coherence tomography using broadband superluminescent diodes,” OSA Biomedical Optics Topical Meetings, Miami Beach, Florida, April 14-17, 2004, paper SE2. Y. Chen, P.R. Herz, P. Hsiung, K. Schneider, J.G. Fujimoto, K. Madden, J. Schmitt, A. Koski, J. Goodnow, and C. Petersen, “Endoscopic imaging in rabit gastrointestinal tract with high resolution optical coherence tomography,” OSA Biomedical Optics Topical Meetings, Miami Beach, Florida, April 14-17, 2004, paper SE4. P. Hsiung, Y. Chen, N. Nishizawa, P.R. Herz, T.H. Ko, J.G. Fujimoto, C.J.S. de Matos, S.V. Popov, J.R. Taylor, J. Broeng, V.P. Gapontsev, “All-fiber, continuous wave, Raman continuum sources for optical coherence tomography,” OSA Biomedical Optics Topical Meetings, Miami Beach, Florida, April 14-17, 2004, paper SE7. T.H. Ko, A.M. Kowalevicz, M.S. Wojtkowski, D. Adler, J.S. Duker, J.S. Schuman, W. Drexler, V. Shidlovski, S. Yakubovich, and J.G. Fujimoto, “Advances in ultrahigh resolution optical coherence tomography imaging of retinal pathology,” 2nd Annual International Society for Imaging in the Eye Meeting, ISIE, Ft. Lauderdale, Florida, April 23-24, 2004, poster B647. M. Wojtkowski, V.J. Srinivasan, T.H. Ko, J.S. Duker, J.S. Schuman, and A. Kowalczyk, “Video rate ultrahigh resolution optical coherence tomography for retinal imaging,” 2nd Annual International Society for Imaging in the Eye Meeting, ISIE, Ft. Lauderdale, Florida, April 23-24, 2004.



A. Chan, J.S. Duker, and J.G. Fujimoto, “Stage Zero Macular Holes: Observations by Optical Coherence Tomography,” Association for Research in Vision and Ophthalmology, ARVO, Fort Lauderdale, Florida, April 25-29, 2004, poster 1967/B778. B. Hermann, A. Unterhuber, M. Wirtitsch, M. Stur, E. Ergun, C. Scholda, T. Ko, J.S. Schuman, J.G. Fujimoto, and W. Drexler, “Quantification Of Photoreceptor Layer Thickness Using Ultrahigh Resolution Optical Coherence Tomography, Association for Research in Vision and Ophthalmology, ARVO, Fort Lauderdale, Florida, April 25-29, 2004, poster 2373/B8. J.G. Fujimoto, T.H. Ko, D.C. Adler, D. Mamedov, V. Prokhorov, V. Shidlovski, S .Yakubovich, M.D. Wojtkowski, J.S. Duker, and J.S. Schuman, “New Technology for Ultrahigh Resolution Optical Coherence Tomography Imaging Using Diode Light Sources,” Association for Research in Vision and Ophthalmology, ARVO, Fort Lauderdale, Florida, April 25-29, 2004, poster 3002/B637. M. Wojtkowski, T.H. Ko, J.G. Fujimoto, T. Bajraszewski, I. Gorczynska, P. Targowski, A. Kowalczyk, J.S. Schuman, and J.S. Duker, “Ultrahigh Speed, Ultrahigh Resolution Optical Coherence Tomography Using Spectral Domain Detection,” Association for Research in Vision and Ophthalmology, ARVO, Fort Lauderdale, Florida, April 25-29, 2004, poster 3009/B644. T.H. Ko, J.G. Fujimoto, J.S. Duker, L.A. Paunescu, A. Chan, W. Drexler, A.H. Rogers, C.R. Baumal, E. Reichel, and J.S. Schuman, “Ultrahigh Resolution Optical Coherence Tomography Imaging of Architectural Morphology in Retinal Diseases,” Association for Research in Vision and Ophthalmology, ARVO, Fort Lauderdale, Florida, April 25-29, 2004, poster 3012/B647. H. Ishikawa, G. Wollstein, M. Aoyama, D. Stein, S. Beaton, J.G. Fujimoto, and J.S. Schuman, “Stratus OCT Image Quality Assessment,” Association for Research in Vision and Ophthalmology, ARVO, Fort Lauderdale, Florida, April 25-29, 2004, poster 3317/B952. G. Wollstein, H. Ishikawa, S. Beaton, D.M. Stein, M. Aoyama, J.G. Fujimoto, and J.S. Schuman, “Pointwise relationship between OCT nerve fiber layer thickness and visual field threshold level,” Association for Research in Vision and Ophthalmology, ARVO, Fort Lauderdale, Florida, April 25-29, 2004, poster 3329/B964. C. De Matos, S.V. Popov, J.R. Taylor, P. Hsiung, Y. Chen, N. Nishizawa, P.R. Herz, T.H. Ko, J.G. Fujimoto, J. Broeng, and V. P. Gapontsev, “Continuous wave, all-fibre broad-band sources for optical coherence tomography,” Conference on Lasers and Electro-Optics, CLEO, San Francisco, California, May 16-21, 2004, paper CWJ1. W. Piyawattanametha, L. Fan, S. Hsu, M. Fujino, M.C. Wu, P.R. Herz, A.D. Aguirre, Y. Chen, and J.G. Fujimoto, “Two-dimensional endoscopic MEMS scanner for high resolution optical coherence tomography,” Conference on Lasers and Electro-Optics, CLEO, San Francisco, California, May 16-21, 2004, paper CWS2. P.R. Herz, Y. Chen, K. Schneider, P. Hsiung, J.G. Fujimoto, K.A. Madden, J.M. Schmitt, A. Koski, J. Goodnow, and C.L. Petersen, ”Ultrahigh resolution optical coherence tomography for in vivo endoscopic imaging of the rabbit,” Conference on Lasers and Electro-Optics, CLEO, San Francisco, California, May 16-21, 2004, paper CWS3. T.-M. Liu, S.-W. Chu, S.-P. Tai, C.-K. Sun, F.X. Kärtner, and J.G. Fujimoto, “Compact self-started femtosecond Cr:forsterite laser for non-linear optical microscopy,” Conference on Lasers and Electro-Optics, CLEO, San Francisco, California, May 16-21, 2004, paper CThFF4.