Polarization-sensitive optical frequency domain imaging based on unpolarized light

Polarization-sensitive optical coherence tomography (PS-OCT) is an augmented form of OCT, providing 3D images of both tissue structure and polarization properties. We developed a new method of polarizationsensitive optical frequency domain imaging (PS-OFDI), which is based on a wavelength-swept source. In this method the sample was illuminated with unpolarized light, which was composed of two orthogonal polarization states (i.e., separated by 180° in the Poincaré sphere) that are uncorrelated to each other. Reflection of these polarization states from within the sample was detected simultaneously and independently using a frequency multiplexing scheme. This simultaneous sample probing with two polarization states enabled determination of the depth-resolved Jones matrices of the sample. Polarization properties of the sample were obtained by analyzing the sample Jones matrices through eigenvector decomposition. The new PS-OFDI system ran at 31K wavelength-scans/s with 3072 pixels per wavelength-scan, and was tested by imaging a polarizer and several birefringent tissues such as chicken muscle and human skin. Lastly the new PS-OFDI was applied to imaging two cancer animal models: a mouse model by injecting cancer cells and a hamster cheek pouch model. These animal model studies demonstrated the significant differences in tissue polarization properties between cancer and normal tissues in vivo. ©2011 Optical Society of America OCIS codes: (170.4500) optical coherence tomography; (170.3880) medical and biological imaging; (260.5430) polarization; (260.1440) birefringence; (110.4500) optical coherence tomography. References and links 1. J. F. de Boer, T. E. Milner, M. J. C. van Gemert, and J. S. Nelson, “Two-dimensional birefringence imaging in biological tissue by polarization-sensitive optical coherence tomography,” Opt. 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Introduction
Polarization-sensitive optical coherence tomography (PS-OCT) is an augmented form of OCT [1,2]. OCT is a 3D imaging technique based on light back-reflected from within tissues [3]. While OCT provides structural information of tissues, PS-OCT provides additional information of tissue polarization properties simultaneously with structure by detecting depth resolved polarization state changes of reflected light. Biological tissues have various polarization properties such as birefringence, diattenuation, and depolarization. Tissue birefringence is originated from either organized microstructures or collagen composition [2]. There are various birefringent tissues, e.g. muscle, tendon, cartilage, dermis of skin, coronary artery, anterior eye segment, retinal nerve fiber layer, and vocal fold. The potential of PS-OCT has been demonstrated in pre-clinical and clinical studies of various organs such as the eye [4][5][6][7][8], skin [9][10][11][12], coronary artery [13], and vocal fold [14,15].
There have been various PS-OCT methods developed [1,[16][17][18][19][20][21][22][23][24][25][26][27][28][29][30][31]. PS-OCT measures sample polarization properties by detecting depth-resolved polarization state changes of reflected light with respect to the incident polarization. Although the amount of polarization change varies depending on the incident polarization state and sample orientation (optic axis), several PS-OCT methods have been developed to reliably determine sample polarization properties. In a bulk-optic setup, circular polarization was used as the incident polarization state to detect linear birefringence of the sample [16,17]. Various optical fiber based methods have been developed for flexible assessment of internal and external organs [18][19][20][22][23][24]26,27]. Unlike bulk-optic based methods, fiber based methods do not have control over the incident polarization state(s) onto the sample, due to the birefringence of optical fibers in the sample arm of an interferometer setup. Therefore, these methods used at least two different incident polarization states for sample probing. One method alternated source polarization states in every depth-scan by using a polarization modulator, such that adjacent depth-scans were acquired with different incident polarization states [19]. Sample polarization properties were calculated by combining information of two adjacent depth-scans. This method had a restriction in lateral scanning speed: since two depth-scans with different incident polarization states should be in principle acquired at (nearly) the same position, the speed of lateral scanning was limited to ensure enough overlap between adjacent wavelength-scans [26]. Recently several PS-OCT methods capable of measuring polarization properties within a single wavelength-scan were developed to overcome this limitation [27][28][29]. These methods used two incident polarization states in single wavelength-scans. One method combined two polarization states into one by encoding them at different carrier frequencies, and sample reflections were separated by frequency demodulation, and is referred as frequency multiplexing [27]. Other methods were based on modulation of the source polarization state at speeds much higher than that of the wavelength-scan. One method modulated the source polarization state at one quarter of the data acquisition frequency such that four polarization states alternated to measure the Stokes vector components within single wavelength-scans [29]. The other method modulated the source polarization state at two thirds of the data acquisition frequency, generating two different polarization states and frequency shifting within single wavelength-scans [28].
We developed a new PS-OCT method by extending the first single wavelength-scan method based on frequency multiplexing. This previous method used two different polarization states separated by 90° in the Poincaré sphere for sample probing. This polarization state combination ensures the reliable measurement of sample's polarization properties: even if one of two incident polarization states is aligned with the optic axis of a birefringent tissue and thus its reflected polarization does not change, the other incident polarization state will incur a maximal change of the polarization state in its reflection. This method was originally developed for the PS-OCT method, which measured two adjacent depth-scans sequentially with alternating incident polarization states [19]. Since this single wavelength-scan method probes with two polarization states simultaneously, a phase relationship between two reflected polarization states is available. This phase relationship can be used to calculate sample's polarization properties, and the previous restriction on the choice of polarization combination, which was 90° separation in the Poincaré sphere, is no longer required. Therefore, this new method utilized the phase relationship of two reflected polarization states. This method probed a sample with unpolarized light, which was a combination of two uncorrelated polarization states separated by 180° on the Poincaré sphere. Reflected polarization states from the sample were measured by using a frequency multiplexing scheme. In this manuscript, we will describe the system design and implementation, polarization analysis method, and various imaging results including both exvivo and in-vivo birefringent tissues, and in-vivo animal cancer models.

System configuration
The schematic of the new system is shown in Fig. 1. Its light source was a wavelength-swept source based on a polygon-scanner filter. The repetition rate of the source was 31K wavelength-scans/s with an output of 40 mW, a bandwidth of 120 nm centered at 1295 nm, and a spectral line width of 0.23 nm corresponding to a one sided depth range of 1.6 mm in air. A small portion (10%) of light from the source was tapped to the combination of a circulator (AC photonics, CA) and a fiber Bragg grating (FBG, OE-land, Canada) to generate a trigger signal for each sweep of the source, and the rest (90%) was used for imaging. Light from the source first passed through an unpolarizer. Input light of the unpolarizer was linearly polarized at 45° with the horizontal axis by using a linear polarizer after a polarization controller, and then was split into horizontal and vertical polarization states in equal amounts by a polarizing beam splitter (PBS). These two polarization states were decorrelated with respect to each other by introducing a relative delay of 9 cm, and were recombined by using another PBS. Then, output light was completely unpolarized.
After the unpolarizer, 90% of the power was directed to the sample arm of an interferometer setup for sample probing, and the remaining 10% was directed to the transmission reference arm. In the reference arm, unpolarized light was split into two orthogonal polarization states by using a PBS. A polarization controller between the unpolarizer and reference arm was adjusted such that the same two linear polarization states, which were generated in the unpolarizer, were generated again in the reference arm. This polarization control was not necessary for the principle of the functionality, but was used in order to minimize a high Relative Intensity Noise (RIN) which will be discussed later. These two polarization states were frequency shifted at 20 MHz and 40 MHz respectively by a pair of acousto-optic modulators (Brimrose Inc, NC) for the frequency multiplexing scheme. These AOMs were driven by a single reference clock for synchronization. This frequency shifting was also advantageous in doubling the imaging depth range to 3.2 mm in air [32]. After the AOM's, the two orthogonal polarization states were recombined by a PBS and coupled into a single-mode fiber. Light from the reference transmission arm was combined with reflected light from the sample for interference in a detection arm, and the interference signal was detected by a polarization-diverse balanced detection setup. In the detection arm, light from the reference arm first passed through a polarization controller, and then a polarizer oriented at 45° with respect to the horizontal axis. This combination of the polarization controller and polarizer ensured that each polarization state, which was frequency shifted in the reference arm, was detected by each channel of the balanced receivers with equal intensity. Signals from the two balanced receivers were acquired simultaneously by an ADC board (PDA14, Signatec Inc, CA) running at 100 MHz sampling frequency. Out of the available signal bandwidth of 50 MHz, signals of individual polarization states occupied separate detection bands: one from 10MHz to 30MHz, and the other from 30MHz to 50MHz. Synchronization between actuators (galvo's for scanning the sample arm beam) and ADC board was configured by using a field programmable logic array (FPGA) demo board (Spartan3, Digilent). The FPGA board used the 100MHz sampling clock from the ADC board as a reference clock, and generated triggering and driving signals based on the reference clock.
The acquired spectra were Fourier transformed into the frequency domain, and then individual frequency bands were demodulated. The demodulated spectra were inverse Fourier transformed to the time domain. Then, the time to k-space mapping was applied by using precalibrated wavelength data and interpolation. Dispersion compensation was applied to the spectra by using a pre-calibration of the dispersion mismatch between reference and sample arms. Finally the spectra in equal k-space were Fourier transformed into reflectivity profiles in depth (z) space. The measured sensitivity was 98 dB, about 17 dB lower compared with the theoretical sensitivity of 115dB, and 7-8 dB lower compared to other published OCT systems. We observed that the RIN of the source increased significantly after the unpolarizer, when a linear polarizer was placed at an orientation of 45° with the horizontal axis. This indicated that despite a 9 cm delay, there still was coherence between the orthogonal polarization states. We attributed this coherence to the narrow line width associated with the cavity mode spacing of the laser (100 MHz). The polarizer or the PBS's in the detection unit acted as the polarizer in this measurement setup. The additional 7-8 dB of noise was due to incomplete suppression by the balanced detection of this RIN, and light intensity fluctuations generated by the frequency shifters. Intensity images were obtained by adding up intensities of both channels and bands, and polarization-sensitive (PS) images were generated by analyzing the obtained complex depth profiles in the procedure described in the next paragraph.

PS analysis method
The PS analysis method for this new PS-OFDI was based on a previous Jones matrix analysis method [33]. This analysis method was originally developed for the PS-OCT which measured two adjacent wavelength-scans with alternating source polarization states. The nonunpolarizing polarization properties of an optical system can be completely described by its complex Jones matrix J, which transforms an incident polarization, described by a complex electric field vector  1 and p 2 are attenuation coefficients along the optic axes and η is the amount of phase retardation. PS images were plotted as accumulative phase retardation with respect to the surface states, and displayed in a gray scale, as black for 0°, and white for 180° phase retardations, and then wrapped back to black for 360°. Diattenuation was quantified by the ratio The new PS-OFDI system and analysis method were tested by imaging various specimens. First, a thin polarizer was used as a pure diattenuator to verify the PS analysis method. The optic axis and transmission coefficients p 1 and p 2 of the polarizer were measured by imaging its top and bottom surfaces. This measurement was repeated with a stepwise rotation of the polarizer by 10° for every measurement until 180° rotation was reached. The measured optic axis was plotted with respect to the set optic axis (Fig. 2a). Ratio of p 1 to p 2 was typically 1/1000, and the correlation between the set and measured optic axis was 0.994 and the standard deviation was 0.02 degrees. Second, chicken muscle and human skin were imaged as representative birefringent tissues. These samples were imaged by using a custom hand-held probe, which had an optical window slanted by 10° with respect to the propagation direction of the sample beam in order to avoid back reflection from the window surface. During imaging, the optical window pressed the tissue samples gently, which flattened tissue surfaces.

System verification
Chicken muscle was imaged ex-vivo and a human hand was imaged in vivo. Representative images are shown in Fig. 2(b-g). Chicken muscle images show relatively amorphous structures in the intensity image (2b), rapid change of accumulated phase retardation throughout the tissue in the PS image (2c), a relatively uniform diattenuation (2d) and the optic axis orientation (2e). The optic axis image shows a uniform orientation of the optic axis in the top layer, over the first black-white-black phase retardation band in Fig. 2c. Then the optic axis appears to change (transition to white), which repeatedly appears with the depth. This apparent change in the optic axis is attributed to the instability of the optic axis orientation algorithm: when the polarization state reflected at a particular depth is nearly equal to the polarization state reflected at the surface, the optic axis needs to be calculated from the difference of nearly identical states. This occurs when the phase retardation reaches 360°, which is at the first, second third, etc black band in Fig. 2e. Human hand images show the epidermis and dermis of skin in the intensity image (2f) and PS image (2g). The change of accumulated phase retardation occurs in the dermis, which is associated with collagen composition. After the verification steps with the test specimens, the new PS-OFDI was applied to imaging two animal cancer models shown in Fig. 3. These were a mouse cancer model by injecting cancer cells subcutaneously, and a hamster cheek pouch model which is a well known oral cancer model [35]. In preparation of the mouse cancer model, a human pancreatic cancer cell line (AsPC-1) was obtained from the American Type Culture Collection (ATCC), and was maintained in a medium recommended by the ATCC at 37°C in a humidified atmosphere of 5% CO2. 6 to 8-week-old male nu+/nu + mice (quantity: 4) were purchased from Cox-7 Animal Facility (Boston, MA). Animal use was approved by the Massachusetts General Hospital IACUC. Mice were subcutaneously injected with AsPC-1 cells (50 x 10 6 /ml) in 100uL Phosphate Buffered Saline solution (VWR) containing 50% matrigel (BD Biosciences) in both legs. Imaging was conducted longitudinally from day 1 until day 10 after injection by using the custom hand-held probe. In preparation of the hamster cheek pouch model, golden Syrian hamsters were purchased from Charles River Laboratories, Wilmington, MA. Premalignant and invasive mucosal pouch lesions were induced by application of 0.2ml of 0.5% 9, 10-dimethylbenzanthracene (DMBA) in mineral oil applied 3 x weekly for 10-15 weeks. During DMBA applications the animals were anesthetized with isoflurane by inhalation. PS-OCT imaging was conducted in vivo under anesthesia, by pulling out the lesion in the cheek pouch from the mouth and by using the custom hand-held probe. Out of longitudinal images of mouse cancer models one representative 3D image, which was taken on day 5, is shown in Fig. 4. Intensity image shows the skin on the surface and the muscle/cancer layer below. The muscle layer can be shown in this image, because the mouse skin layer is thin. The skin is highly scattering with layered structures, and the muscle/cancer layer is less scattering with homogeneous structures. The cancer section appears more scattering than the normal muscle section. These two sections are very well distinguishable in the PS image due to the difference in tissue birefringence. The cancer section appears to be separated from the surrounding normal muscle section in both intensity and PS images. This separation was confirmed by histology.

In vivo animal model imaging
A 3D reconstructed image of the hamster cheek pouch model is shown in Fig. 5. In the intensity image, the cancer and normal cheek pouch tissue sections are not distinguishable, because the tissue surface was deformed by the optical window of the hand-held probe. However, the PS image shows significant contrast between them based on the banding pattern, because the normal cheek pouch tissue is birefringent and the cancer tissue is not. Some banding pattern appears at the bottom of the cancer tissue, and this may indicate the presence of normal tissues below the cancer.  Fig. 4.

Discussion
A new PS-OFDI system based on unpolarized light was implemented successfully. Unpolarized light, composed of two uncorrelated polarization states separated by 180° on the Poincaré sphere, probed a sample, and the depth resolved Jones matrix of the sample was constructed by combining the two measurements with different incident polarization states. Because reflection of the two incident polarization states was measured simultaneously, the phase relationship between reflected polarization states was available and the full sample Jones matrix could be measured and analyzed by eigenvector decomposition. This new method was verified by imaging various control specimens such as a polarizer, chicken muscle, and skin.
Since these Jones matrices contain all the information of polarization properties of the sample, individual parameters can be extracted by an analytical decomposition method [36]. However, a simple eigenvector decomposition method provided better immunity to noise and was used in our PS analysis. A more robust analytical analysis method with better noise immunity may need to be developed to take advantage of the full sample Jones matrices to e.g. determine the optic axis of diattenuation and birefringence independently.