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Picosecond Electronic Time-Gated Imaging of Bones in Tissues Manuel E. Zevallos, S. K. Gayen, Bidyut Baran Das, Mohammad Alrubaiee, and R. R. Alfano, Senior Member, IEEE

Abstract— Two-dimensional in vivo optical images of metacarpal bones of a human palm and in vitro images of turkey and chicken bones embedded in tissues were obtained in the nearinfrared region using femtosecond pulse transillumination and picosecond electronic time-sliced detection technique. A small hole drilled in chicken bone and embedded in chicken breast tissue was imaged using early arriving light. Time-gated fluorescence images of an interior marrow region of a bone injected with a fluorescent dye were recorded. The techniques have potential for monitoring bone fracture, bone diseases such as, osteoporosis and arthritis, and diseases that originate in or affect bone marrow. Index Terms— Biomedical optical imaging, bone diseases, bones, fluorescence imaging, fluorescence lifetime, imaging of objects in turbid media, near-infrared imaging, time-gated imaging, transillumination imaging.

I. INTRODUCTION

T

HE RECENT surge in interest for developing optical biomedical imaging and diagnostic modalities [1]–[30]1 have encompassed lesions in many organs of the human body, such as, breast, brain, prostrate, gynecological tract, but there is a lack of work on imaging of bone tissues [7], [8], [23]–[26]. Noninvasive optical imaging of bone abnormalities found in cases of bone tumors, osteoporosis, arthritis, and other diseases would be a highly desirable tool for the biomedical community. More than half of all cancer patients at the time of death show metastatic bone lesions [31]. New bone is formed in a disorganized fashion in osteoblastic bone lesions while bone is destroyed causing fractures in osteolytic lesions. These lesions commonly occur in vertebral column, ribs, pelvis as well as arm and leg bones. Osteosarcoma, a malignancy of immature bone, is more common in children. In case of osteoporosis, bone becomes thinner and porous making it more fragile. Change in bone density at the onset of osteoporosis would alter the scattering properties of bone which can be detected by optical measurements. In osteoarthritis, there is a deterioration of the articular cartilage which Manuscript received January 27, 1999. This work was supported by the New York State Science and Technology Foundation, by the NAA IRA Program DOE through the Center of Excellence Program, by the Heat program of the City University of New York, and by the Mediscience Technology Corporation. The authors are with the Institute for Ultrafast Spectroscopy and Lasers, New York State Center for Advanced Technology for Ultrafast Photonic Materials and Applications, Department of Physics, City College of the City University of New York, New York, NY 10031 USA. Publisher Item Identifier S 1077-260X(99)07534-6. 1 There has been a profusion of work on optical imaging and diagnostics since the early 1980’s. For a collection of representative papers, see [1].

might have been caused due to orthopedic abnormalities or injuries. In rheumatoid arthritis, inflammation and thickening of the synovial membranes (the sacs that hold the fluid that lubricates the joints) cause irreversible damage to the joint capsule and the articular (joint) cartilage as these structures are replaced by scar tissue. Acute inflammation with swollen and pus-filled joints occurs in septic arthritis [31]. Successful optical modalities for monitoring injuries, growths, and other abnormalities in bone and bone marrow can be potentially helpful in the early detection of osteoporosis, arthritis, bone tumor, leukemia, malignant bone marrow tumor, melanoma, and benign eosinophilic granuloma and other diseases. In this paper, we present the results of two different optical imaging experiments in an attempt to explore the potential of light-based methods in detecting and diagnosing lesions that affect bones or have manifestations in bone marrow. First experiment involved two-dimensional (2-D) time-sliced transillumination imaging of bones in tissues. Specific systems investigated include metacarpal bones of human palm in vivo, and turkey and chicken bones in vitro. Although there is a paucity of data on the optical parameters of bone some preliminary measurements suggest it to be several times more highly scattering than muscular tissues [32].2 This difference in scattering characteristics provides a basis for time-gated transillumination measurements since higher scattering would entail a greater loss of transmitted photons at the early time casting a shadow of the bone. One would expect the shadow to diminish at later times as it would be washed out by a profusion of multiscattered photons. The second experiment was performed to image a small interior region of bone marrow inside a chicken bone using an external fluorescent marker. Some fluorescing dyes like hematoporphyrin derivative (HPD) are known to bind to tumor cells longer than normal surrounding tissues. This method has led to investigation of extrinsic fluorescence mapping as a tool for malignant tissue imaging [27]. Here, we have attempted to explore the feasibility of fluorescence mapping of bone marrow as the surrounding wall of highly scattering bone poses a formidable challenge to generate fluorescence inside bone marrow and detect the highly weakened signal outside. Imaging of bone marrow was performed after injecting fluorescing Coumarin 440 (Exciton Chemicals) dye into a chicken bone as an external marker. Though we are not aware of Coumarin 440 dye binding to any diseased cell we have 2 Transport

mean-free path for a skull bone was found to be around 0.2 mm.

1077–260X/99$10.00  1999 IEEE

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(b) Fig. 1. (a) A schematic diagram of the experimental arrangement for time-sliced imaging measurements. (b) The alterations in the experimental setup shown in Fig. 1(a) required for the fluorescence imaging measurements. The second harmonic beam was incident on a 5-mm-thick chicken bone that had been injected with Coumarin 440 dye.

used it as an extrinsic marker for its high quantum efficiency only to test the viability of this optical modality for bone marrow imaging. II. EXPERIMENTAL METHODS A schematic diagram of the experimental setup used in time-sliced imaging are displayed in Fig. 1(a). Ultrashort laser pulses of about 150 fs duration at 800 nm, at an 82 MHz repetition rate from a self-modelocked Ti:sapphire laser (Spectra Physics Tsunami) were amplified with a regenerative

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amplifier. The average amplified beam power used in the transillumination imaging experiment was approximately 200 mW at a repetition rate of 1 kHz. The beam was expanded and the central part of it was selected using an aperture to illuminate the sample. This was done in order to make the spatial intensity distribution of the beam smoother and to reduce the intensity difference between the center and the periphery of the beam. Light transmitted through the sample was recorded using an ultrafast electronic-gated imaging camera system (La Vision, Pico-Star) consisting of a time-gated image intensifier unit fiber optically coupled to a charge-coupled-device (CCD) camera. The transmitted light from the sample was collected by a camera lens and directed onto the image intensifier. The imaging system provided an electronic gate pulse whose fullwidth at half-maximum (FWHM) duration could be adjusted to a minimum of approximately 80 ps, and whose position could be varied in steps of 25 ps over a range of 20 ns. The transillumination signal recorded by the system at a particular gate position was a convolution of the transmitted light pulse with the gate pulse centered on the gate position. Time-sliced transillumination images of several bone samples were recorded. These include: 1) a turkey drum with a bone of approximately 12-mm diameter compressed between two glass plates to a uniform thickness of 60 mm over a cross-sectional area of about 20-mm diameter; 2) a chicken bone of approximately 8.5-mm diameter embedded in a 38mm-thick chicken tissue; and 3) a third sample of chicken wing to resolve two narrow bones of about 3-mm thickness with a separation of about 7 mm, embedded in a 32-mm-thick chicken breast tissue. In order to find out the potential of this time-sliced imaging method to detect fractures in bones a small hole of 1.5-mm diameter was drilled on a 5.5-mm-thick and 8.5-mm-wide chicken thigh bone. The bone was embedded in a 20-mm-thick chicken tissue held between two glass plates with the hole facing toward the incident beam. The sample used for in vivo transillumination imaging was a 27-mm-thick human palm. The changes in the experimental arrangement of Fig. 1(a) required for fluorescence imaging are indicated in Fig. 1(b). The 400-nm second-harmonic (SH) radiation of the amplified Ti:sapphire fundamental beam at 800 nm was used for exciting the dye fluorescence. A phase-matched beta barium borate (BBO) crystal was used to generate the SH beam. A narrow band filter was used to cut off the fundamental frequency and transmit the 400-nm light, which was loosely focused using a 30-cm focal-length lens to a spot size of about 2 mm on a 5-mm-thick chicken bone. The 31-mm-long bone had a taper shape with a 14-mm base and a 6-mm top end. A solution of Coumarin 440 dye in methanol (concentration 180 mg/l) was injected into the bone marrow. The dye was excited by a fraction of the 400-nm beam that entered into the bone. The fluorescence signal that exited in the forward direction was collected using a camera lens through a bandpass filter centered at 450 nm to cut off the 400-nm excitation light and directed to the image intensifier of the Pico Star camera. The images were recorded by varying the gate position in steps of 25 ps over a 12-ns range, and a temporal profile of fluorescence decay was constructed from this series of measurements.

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(c) Fig. 2. 2-D images of a 12-mm-diameter bone in a 60 mm thick turkey drum at 75- and 1200-ps time gates. Spatial intensity profiles were obtained by integrating the intensity over a 20-pixel-wide horizontal region shown by dashed boxes in (a) and (b). (b)

III. RESULTS

Fig. 3. (a) 2-D image of a chicken wing with two bones of about 3-mm thickness with a separation of about 7 mm embedded in 32 mm tissue. (b) a spatial intensity profile integrated over a 20-pixel-wide region highlighted by the dashed box in (a).

A. Transillumination Imaging of Bones Transillumination images of the turkey drum for two different time gates centered at 75 and 1200 ps are displayed in Fig. 2(a) and (b), respectively. These time gates were selected relative to the zero-position which was taken to be the time of arrival of light pulses after traversing a cell of water with a thickness equal to that of the sample. The spatial distribution of transmitted light intensity shows the location of the bone clearly (the darker part) when the early arriving, image-bearing photons are chosen by the time-gate centered at 75 ps [Fig. 2(a)]. In addition, some structural features of the muscle such as joining of two layers can be observed in the image. At later times the image deteriorates as more diffuse photons contribute to noise and overwhelm the imagebearing photons. The spatial intensity profiles integrated over a horizontal area with a vertical width of about 20 pixels are shown in Fig. 2(c) to give a measure of the intensity contrast near the bone. The FWHM of the dip found in the spatial intensity profile at the 75 ps-gate is about 22 mm, compared to the actual width of the bone of about 12 mm. This difference is expected since the image is formed by collecting mostly snake photons with hardly any coherent (ballistic) light which would have given a much sharper image. To have a measure of the image quality, image contrast, , was calculated as defined below: (1) where maximum and minimum intensities were calculated from the spatial intensity profiles. Contrast for the first image (75-ps gate position) was found to be about 0.74 and that was

greatly reduced for the second image. This can be attributed to much higher scattering in bone, which causes a loss of forward-scattered light at the early time resulting in the central dip in the intensity profile. But, as we go to later time slices, more scattered photons are detected washing out the imagebearing signal and the intensity dip almost disappears. The smaller dip to the left of the bone is at the boundary of two overlapping layers of muscle. Similar time-sliced 2-D transillumination image was also obtained for an 8.5-mm-thick chicken bone embedded in a 38mm-thick chicken tissue. In order to examine if two closely situated thinner bones can be discerned by this technique similar 2-D images were obtained for a chicken wing with two 3-mm-thick bones at a 7-mm separation inside a 32-mmthick chicken tissue. The image for the gate position at 25 ps and a spatial intensity profile are displayed in Fig. 3(a) and (b), respectively. The profile shows two dips clearly resolving the two thin bones even though the contrast has been reduced to around 0.4. 2-D transillumination images of a 1.5-mm-diameter hole in a 8.5-mm-wide and 5.5-mm-thick bone inside a 20-mm-thick chicken tissue for the gate positions at 0, 75, and 225 ps are displayed in Fig. 4(a)–(c), respectively. The bright spot at the center of the first image [Fig. 4(a)] gives the location of the hole. When the time gate is located at 0 ps the detector picks up a large number of snake photons traversing through the hole resulting in the formation of the central bright disk. The shadow of the bone can be seen horizontally across and in the

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(c) Fig. 4. 2-D time-sliced images of a 1.5-mm hole drilled in a 5.5-mm-wide bone embedded in a 20-mm-thick chicken tissue. Two spatial intensity profiles were obtained for each image: one in the region with the hole and the other away from it as displayed in white dashed and solid boxes, respectively, in the images.

middle of all three images. When we go to gate positions at 75 and 225 ps the hole disappears leaving a more uniform shadow in the middle and high intensity distributions from diffuse light on the two sides. In order to find out how the contrast of the image of the hole changed with the gate position we plotted two spatial profiles for each image. The profiles were obtained by integrating the intensity over two different vertical regions (20-pixel-wide each) of the image—one that included the hole (dashed line), and the other was about 20 pixel away from it (solid line). The contrast was calculated by comparing the two spatial intensity profiles for each image: by taking the and intensity through the hole from the first profile as the intensity away from the hole from the second profile as [see the profiles in Fig. 4(a)]. With a strong intensity peak through the hole, the profiles from the first image yield a contrast of 0.4. The profiles from the second image taken at 75-ps gate position show two broad peaks from the diffuse light traversing through tissue and a central valley caused by the highly scattering bone. The intensity peak at the hole is reduced almost completely (contrast of 0.08) as most of the photons traversing through the hole had reached the detector earlier. As we move our gate position further to 225 ps the central valley caused by the bone is reduced as more diffuse photons reach the detector washing out the shadow.

(b) Fig. 5. (a) 2-D images of the metacarpal bones of a human palm. The darker part long the vertical direction and to the left of the bright central region is due to the vein going across the gap between the middle and the ring fingers. (b) Spatial intensity profile integrated over the region enclosed by the dashed box in (a).

Transillumination measurements were performed in vivo on a human palm in order to image the metacarpal bones. Fig. 5(a) displays a 2-D image of the palm obtained using the time-sliced detection. The 3-cm-diameter beam was centered on the middle finger as shown in the figure and the timegate was set at 0 ps to collect the image bearing photons. The images of the three metacarpal bones (ring, middle and index) are clearly resolved at this time-window. The darker parts show the location of the bones and the lighter parts show the gaps between them. A spatial profile, shown in Fig. 5(b), was obtained by integrating the intensity over a 25-pixel-wide vertical region that gave a contrast of about 0.66 for the middle finger. The vein going across the region between the middle and the ring fingers appears as the vertical darker part next to the bright central region. Some of the early arriving photons passing through this region are absorbed by the vein resulting in the dark line. B. Fluorescence Imaging Fluorescence images of the interior of the bone obtained in the transillumination geometry for two different positions of the time-gate are shown in Fig. 6(a) and (b), when the

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Fig. 7. Double exponential fitting of the temporal profile of fluorescence intensity from Coumarin 440 dye solution injected into bone marrow. The dotted line is the theoretical fit to the experimental curve (solid line). The fast decay time constant f was found to be 568 and 730 ps for the dye solution in quartz cell and dye + bone system, respectively. The slow decay constant s was 2577 and 3286 ps for the respective cases. The ratio of amplitudes af =as were 1.33 for the dye in quartz cell and 0.67 for the other case.

(d)

2

Fig. 6. (a), (b) Two magnified fluorescence images of a 2 mm 2 mm region of the bone marrow taken at 750-ps and 10-ns gate positions, respectively. (c) Photograph of the 5-mm-thick bone used in the fluorescence imaging experiment. The black mark in the lower middle region of the bone shows the point of incidence of SH excitation beam. (d) Temporal fluorescence intensity profiles measured from Coumarin 440 dye solution in quartz cell, and dye bone system. The temporal profiles were constructed from a series of such images at gate positions between 0 and 12 ns.

+

laser beam was incident on the bone shown in Fig. 6(c). The fluorescence was collected from the opposite side of the bone. The images recorded at gate positions of 750 ps and 10 ns show a 2 mm 2 mm area of the bone marrow impregnated with the dye. The overall spatial intensity distribution is much higher in 750-ps image than the 10-ns image, a consequence of the higher fluorescence light available at earlier times. Dye fluorescence decays in time and, consequently, less light is available at the later times. To examine the fluorescence decay characteristics and its effect on fluorescence based imaging, we studied the time evolution of Coumarin 440 fluorescence within the bone marrow and compared it with the fluorescence decay of the dye solution in methanol contained in a quartz cell. The two normalized temporal fluorescence intensity profiles are shown in Fig. 6(d). The heavy curve is from the dye solution in methanol and the light curve is from the dye in the bone marrow. These profiles were constructed from the sequence of images that were recorded as the gate position was moved in 25-ps steps between 0–12 ns. The integrated intensity of certain area of each image was plotted as a function of the gate position for which the image was recorded to construct the temporal profiles. Since the signal from the dye solution in

quartz cell was much stronger than that from the dye solution in bone marrow the two profiles were normalized to see the differences between them. As is evident from Fig. 6(d), the two profiles follow almost the same decay curve up to 1 ns and then deviate. The profile bone) extends much of the dye inside bone marrow (dye farther in time than that of the dye alone. This is expected from the nature of photon propagation in turbid media. The fluorescence from dye bone system will have ballistic, snake and diffuse components [2] as the fluorescence making its way out of the bone will undergo scattering. The ballistic component will follow the decay profile of the dye in quartz bone fluorescence will cell more closely. The tail of dye be lengthened because of multiple scattering. The chemical environment presented by bone marrow is also expected to have an effect on the fluorescence of the dye bone system. However, that effect is expected to be weaker than the effect of scattering by bone marrow and bone wall. For a quantitative understanding we obtained fits of the two decay profiles, which on close examination revealed double exponential decays. The fits of the two profiles to a double exponential of the form (2) where the subscripts and stand for the slow and the fast components, respectively, were obtained. Fig. 7 displays the dye system fit of the fluorescence decay from the bone where the dotted line represents the theoretical fit. The fast was found to be 568 and 730 ps for decay time constant bone system, respectively. The the dye solution and dye slow decay constant was found to be 2577 and 3286 ps for was 1.33 the respective cases. The ratio of amplitudes

ZEVALLOS et al.: PICOSECOND ELECTRONIC TIME-GATED IMAGING OF BONES IN TISSUES

for the dye solution in quartz cell and 0.67 for the other case, a very significant difference. It is evident from the profiles and values of fit parameters that the 750-ps image was recorded with the fast component of the fluorescence, while the 10-ns image was obtained with the slow component. This forms the basis for fluorescence lifetime imaging where maps of two or more fluorophores emitting in the same spectral range but with different decay times could be separately imaged. IV. DISCUSSION The results obtained from these time-resolved transillumination and fluorescence imaging measurements underline the potential that time-sliced optical imaging holds for application in biomedical imaging. We have shown the feasibility of time-sliced imaging of bones in thick tissues and a small hole in a bone embedded in tissue. In vivo measurement of metacarpal bones yielding high contrast 2-D images is highly encouraging to pursue this avenue in development of viable alternatives for monitoring bone tumor, bone marrow malignancy, osteoporosis, arthritis, rheumatic disorders and other bone injuries. Our ability to obtain 2-D images directly in a few seconds of data accumulation is significant from patient’s point of view in real life applications. The fluorescence imaging measurements of early and late emissions using Coumarin 440 dye show the feasibility of seeing the interior of a bone. Our ability to obtain 2-D images of a small region of bone marrow and detect significant differences in the decay kinetics is highly promising in developing an optical modality for monitoring small regions of abnormalities in bone and bone marrow. Coumarin 440 dye was used only as a test case. External fluorophores that bind to diseased cells should be investigated in the visible and NIR region. It has been shown previously that breast malignancy [29] and atherosclerotic plaque [30] can be detected by measuring fluorescence decay kinetics. By finding suitable fluorescence markers, native or external, one can similarly study bone marrow. Fluorescence kinetic measurements on diseased and normal marrow may show signature differences in their decay patterns which can be used for fluorescence mapping. By using a broad beam one can obtain 2-D fluorescence image of a large region of bone marrow. Temporal intensity profiles can be constructed for various smaller regions from this image as we have shown above. By measuring the fluorescence decay parameters from each profile various maps of decay amplitudes and/or time constants can be constructed giving information about various abnormalities. It should be noted that the ability to obtain images using the fast and slow components forms the basis for imaging based on fluorescence lifetime differences as has been discussed in the earlier section. Even though this line of research is at an early stage requiring further investigation our preliminary results show a lot of promise for future applications. More work is needed to find suitable external fluorescence markers that bind to various diseased cells of bone marrow and to investigate different excitation wavelengths in visible and NIR region. Further improvements in the electronic time-gated imaging system

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would also enhance the precision in time-slicing, which would, consequently, improve the transillumination image quality. REFERENCES [1] V. V. Tuchin, Selected Papers on Tissue Optics: Applications in Medical Diagnostics and Therapy, SPIE Milestone Series, vol. MS102, 1994. [2] B. Chance, “Optical method,” Annu. Rev. Biophys. Biophys. Chem., vol. 20, pp. 1–28, 1991. [3] S. K. Gayen and R. R. Alfano, “Emerging optical biomedical imaging techniques,” Opt. Photon. News, vol. 7, no. 3, pp. 17–22, 1996. [4] B. B. Das, F. Liu, and R. R. Alfano, “Time-resolved fluorescence and photon migration studies in biomedical and model random media,” Rep. Prog. Phys., vol. 60, pp. 227–292, 1997. [5] J. C. Hebden, S. R. Arridge, and D. T. Delpy, “Optical imaging in medicine—I: Experimental techniques,” Phys. Med. Biol., vol. 42, pp. 825–840, 1997. [6] S. R. Arridge and J. C. Hebden, “Optical imaging in medicine—II: Modeling and reconstruction,” Phys. Med. Biol., vol. 42, pp. 841–853, 1997. [7] S. Anderson-Engels, R. Berg, S. Svanberg, and O. Jarlman, “Timeresolved transillumination for medical diagnostics,” Opt. Lett., vol. 15, pp. 1179–1181, 1990. [8] J. R. Lakowicz and K. Berndt, “Frequency domain mesurements of photon migration in tissues,” Chem. Phys. Lett., vol. 166, no. 3, pp. 246–252, 1990. [9] L. O. Svaasand, B. J. Tromberg, R. C. Haskell, T.-T. Tsay, and M. W. Burns, “Tissue characterization and imaging using photon density waves,” Opt. Engg., vol. 32, no. 2, pp. 258–266, 1993. [10] T. French, E. Gratton, and J. Maier, “Fequency domain imaging of thick tissues using a CCD,” in Time-Resolved Laser Spectroscopy III, Proc. SPIE, 1992, vol. 1640, pp. 254–261. [11] M. Ferrari, R. A. De Blasi, F. Safoue, Q. Wei, and G. Zaccanti, “Toward human brain near infrared imaging: Time resolved and unresolved spectroscopy during hypoxic hypoxia,” in Optical Function of Brain Function and Metabolism. New York: Plenum, 1993, pp. 21–31. [12] B. B. Das, K. M. Yoo, and R. R. Alfano, “Ultrafast time-gated imaging in thick tissues: A step toward optical mammography,” Opt. Lett., vol. 18, pp. 1092–1094, 1993. [13] K. G. Spears, J. Serafin, N. H. Abramson, X. Zhu, and H. Bjelkhagen, “Chronocoherent imaging for medicine,” IEEE Trans. Biomed. Eng., vol. 36, pp. 1210–1221, 1989. [14] E. Leith, E. Arons, H. Chen, Y. Chen, D. Dilworth, J. Lopez, M. Shih, P. C. Sun, and G. Vossler, “Electronic holography for imaging through tissue,” Opt. Photon. News, vol. 4, no. 10, pp. 19–23, 1993. [15] V. V. Tuchin, “Laser light scattering in biomedical diagnostics and therapy,” J. Laser Applicat., vol. 5, no. 2/3, pp. 43–60, 1993. [16] M. E. Brezinski, G. J. Tearny, S. A. Boppart, B. Bouma, and J. G. Fujimoto, “High resolution intra-arterial imaging with optical coherence tomography,” in Tech. Dig. Advances in Optical Imaging and Photon Migration. Washington, DC: Opt. Soc. Amer., 1996, pp. 21–23. [17] S. G. Demos and R. R. Alfano, “Optical polarization imaging,” Appl. Opt., vol. 36, pp. 150–155, 1997. [18] J. F. de Boer, T. E. Milner, M. J. C. van Gemert, and J. S. Nelson, “Two-dimensional birefrigence imaging in biological tissue by polarization-sensitive optical coherence tomography,” Opt. Lett., vol. 22, pp. 934–936, 1997. [19] Y. Guo, P. P. Ho, H. Savage, D. Harris, P. Sacks, S. Schantz, F. Liu, N. Zhadin, and R. R. Alfano, “Second-harmonic tomography of tissues,” Opt. Lett., vol. 22, pp. 1323–1325, 1997. [20] W. Denk, “Two-photon excitation in functional biological imaging,” J. Biomed. Opt., vol. 1, pp. 296–304, 1996. [21] Y. Guo, Q. Z. Wang, N. Zhadin, F. Liu, S. G. Demos, D. Calistru, A. Turksliunas, A. Katz, Y. Budansky, P. P. Ho, and R. R. Alfano, “Twophoton excitation of fluorescence from chicken tissue,” Appl. Opt., vol. 36, pp. 968–970, 1997. [22] S. K. Gayen, M. E. Zevallos, M. Alrubaiee, J. M. Evans, and R. R. Alfano, “Two-dimensional near-infrared transillumination imaging of biomedical media with a chromium-doped forsterite laser,” Appl. Opt., vol. 37, pp. 5327–5336, 1998. [23] B. Devaraj, M. Takeda, M. Kobayashi, M. Usha, K. P. Chan, Y. Watanabe, T. Yuasa, T. Akatsuka, M. Yamada, and H. Inaba, “In vivo laser computed tomographic imaging of human fingers by coherent detection imaging method using different wavelengths in near infrared region,” Appl. Phys. Lett., vol. 69, pp. 3671–3673, 1996. [24] A. W. Wist, P. Moon, S. Meiksin, S. L. Herr, and P. P. Fatouros, “High resolution light imaging system for teeth and tissues,” J. Clinical Laser

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Manuel E. Zevallos was born in Guayaquil, Ecuador, on December 25, 1965. He became a U.S. citizen in 1991. He received the B.E. and M.E. degrees in electrical engineering from the City College of the City University of New York, New York, NY, in 1992 and 1996, respectively, and the Ph.D. degree in electrical engineering from the Graduate School of the City University of New York while working at the Institute for Ultrafast Spectroscopy and Lasers. His dissertation research involved near-infrared optical imaging and light propagation in highly scattering random media. He joined IBM in February 1999 where he is currently working in the Microelectronics Division.

S. K. Gayen received the B.Sc. (honors) and M.Sc. degrees in physics from the University of Dacca, Bangladesh, and the Ph.D. degree in physics from the University of Connecticut, Storrs. He is currently a Senior Member of the Research Staff at the New York State Center for Advanced Technology for Ultrafast Photonic Materials and Applications at the City University of New York. His research interests are in optical biomedical imaging, tunable solid-state lasers, spectroscopy of impurity ions in solids, nonlinear optics, and ultrafast laser spectroscopy. Dr. Gayen is a member of the American Physical Society.

Bidyut Baran Das received the B.Sc. (Honors) degree in physics from Ravenshaw College of Utkal University, Orissa, India, the M.Sc. degree in physics from the Indian Institute of Technology, Kanpur, and the Ph.D. degree in physics from the City University of New York, New York, NY. Currently, he is a Research Scientist at the Institute for Ultrafast Spectroscopy and Lasers, Physics Department, the City College of New York. His research has encompassed studies of timeresolved light scattering in random media, steady-state and time-resolved fluorescence from human tissues for developing optical diagnostic methods, optical biomedical imaging, and inverse image reconstruction.

Mohammad Alrubaiee received the B.S. degree in electric engineering and M.Sc. degree in physics from the City College of the City University of New York. He a Research Assistant at the Institute for Ultrafast Spectroscopy and Lasers, Physics Department, the City College of New York. He is pursuing his graduate research in optical biomedical imaging and laser spectroscopy.

R. R. Alfano (M’87–SM’89) is a Distinguished Professor of Science and Engineering at the City College of the City University of New York, and Director of the Institute for Ultrafast Spectroscopy and Lasers, the New York State Center for Advanced Technology for Ultrafast Photonic Materials and Applications, and the Center for Laser Imaging and Cancer Diagnostics, a Department of Energy Center of Excellence at the City College of New York. His research has encompassed a wide variety of areas including optical biomedical imaging, photon propagation through turbid media, ultrafast laser science and technology, ultrafast supercontinuum generation, tunable solid-state lasers, nonlinear optics, as well as dynamical processes in semiconductors, dielectric crystals, molecular systems, polymers, and biological systems. He has published over 500 papers, edited four books on ultrafast laser science and applications, and holds 66 patents. Prof. Alfano is a Fellow of the American Physical Society and the Optical Society of America.