Polarization-sensitive interleaved optical coherence tomography Lian Duan, Tahereh Marvdashti, and Audrey K. Ellerbee∗ E. L. Ginzton Laboratory and Department of Electrical Engineering Stanford University, Stanford, CA 94305, USA ∗ [email protected]
Abstract: We introduce a new strategy for single-mode fiber based polarization-sensitive (PS-) optical coherence tomography (OCT) using orthogonally polarized optical frequency combs (OFC) in the sample arm. The two OFCs are tuned to be interleaved in the spectral domain, permitting simultaneous measurement of both polarization states from the same spatial region C close to the location of zero pathlength delay. The two polarization states of the beam in the sample arm are demultiplexed by interpolation after performing wavelength stabilization via a two-mirror calibration method. The system uses Jones matrix methods to measure quantitatively the round-trip phase retardation B-scans in the sample. A glass plate and quarter-wave plate were measured to validate the accuracy of the birefringence measurement. Further, we demonstrated the potential of this system for biomedical applications by measurement of chicken breast muscle. © 2015 Optical Society of America OCIS codes: (170.4500) Optical coherence tomography; (170.3880) Medical and biological imaging; (260.5430) Polarization; (260.1440) Birefringence.
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1.
Introduction
Polarization-sensitive (PS-) optical coherence tomography (OCT) is a technique that adds birefringence contrast to the structural contrast provided by conventional OCT [1–3]. It is well known that biomedical tissues containing aligned collagen fibers exhibit strong birefringence; thus, changes in the birefringent properties of these tissues are often indicative of disease. Consequently, PS-OCT has been used to enhance disease diagnosis in several clinical areas, including ophthalmology [4–7], dermatology [8–11], dentistry [12–14], cardiology [15, 16], and cancer detection [17]. Free-space or polarization-maintaining (PM)-fiber-based PS-OCT systems can directly measure birefringence using a single incident polarization state (SIPS) [1,4,18–23]. However, freespace systems require delicate alignment and have limited utility for applications requiring flexible sample delivery, such as endoscopy, while PM-fiber-based systems can be costly, bulky and suffer from polarization cross-talk and alignment challenges [8,18,20,24,25]. Recently, singlemode (SM) fiber-based SIPS PS-OCT systems have been proposed by several groups [26, 27]. A major challenge of these systems is the need for all segments of the SM fiber need to be well fixed (e.g., using polarization controllers) to prevent polarization scrambling in the fiber; hence, such systems are not suitable for portable, handheld or endoscopic applications. In contrast, SM fiber-based PS-OCT systems that rely on illumination of the sample with multiple input polarization states (MIPS) have been successfully utilized for endoscopy [16, 28, 29]. Unfortunately, designs that exploit time-encoding of the input polarization state suffer from reduced imaging speed due to the need for sequential measurements [16, 30], and strategies invoking multiple light sources or polarization modulators incur significant additional costs [31–33]. Some have used depth-encoding to separate images from different polarization states [34–36], but these methods lead to lower sensitivity for at least one polarization channel, which in turn leads to a lower effective signal-to-noise ratio (SNR) and increased systematic and random errors in the birefringence measurement [37, 38]. The recent emergence of shortcavity swept source lasers with long coherence lengths, low sensitivity roll-off versus depth and passive delay units does mitigate some of the asymmetric sensitivity challenges in depth#238168 - $15.00 USD Received 16 Apr 2015; revised 8 May 2015; accepted 11 May 2015; published 15 May 2015 © 2015 OSA 18 May 2015 | Vol. 23, No. 10 | DOI:10.1364/OE.23.013693 | OPTICS EXPRESS 13695
PC 3
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Fig. 1. Schematic of the PS-iOCT system with orthogonally polarized optical frequency combs. FPE: Fabry-Perot etalon; PC: polarization controller; FPBS, fiber polarizing beam splitter; CIR: circulator; BS: non-polarizing beam splitter; PBS: polarizing beam splitter; G, galvanometers.
encoded systems [35,36], but these systems still suffer from sensitivity decay in the polarization channel far from the zero-delay-line. In this work, we present a new configuration for MIPS, SM-fiber PS-OCT whereby two orthogonally polarized inputs are generated in the form of interleaved optical frequency combs from a single source and interrogate the sample simultaneously (i.e., maintaining imaging speed). We thus call the method PS-interleaved OCT (PS-iOCT). In contrast with aforementioned systems, polarization multiplexing in PS-iOCT is achieved by passive optics (standard etalons) rather than active modulators (e.g., electro-optic modulator or acousto-optic modulator), making the scheme robust, economical, and immune to the sensitivity issues caused by modulated waveforms. Moreover, because of the interleaved strategy, both polarization channels appear at the same depth in the OCT image; hence, the two channels are not separately affected by the inherent sensitivity roll-off associated with depth-encoded schemes. In a proof-of-principle experiment, we demonstrate the ability of PS-iOCT to measure birefringence in calibrated and biological samples. We also demonstrate how this method can be implemented using a novel design for phase calibration using a single digitizer having only two detection channels, a potential cost savings. 2.
System design
Figure 1 shows the schematic of the PS-iOCT system. The light source is a MEMS short-cavity swept source (VCSEL, Thorlabs) with an average power of 27 mW, a center wavelength of 1310 nm, a 200-kHz sweep rate over 100 nm of full-width at half-maximum bandwidth, and > 50mm coherence length. The light is connected to an unbalanced Mach-Zehnder interferometer with a 90/10 coupler, which serves as the main interferometer for the system. The 10% arm is connected to the reference arm, while the 90% light is evenly split by a 50/50 coupler. Each portion of the latter is sent to a Fabry-Perot etalon (FPE) via a circulator to generate an optical frequency comb (OFC). The two combs emerging from this circulator then illuminate both the sample and a calibration mirror, whose use in wavelength stabilization will be described later. Light returning from the sample and reference arms are mixed and detected by a freespace polarization-sensitive detection unit. The illumination power at the sample surface was measured to be 2.4 mW. The two Fabry-Perot etalons in the sample arm both comprise a highly reflective silver mirror and an optical flat. The optical flats were uniformly coated with 35A˚ CR and 45A˚ Au, yielding reflectivities of 31% at the air/Au and 22% at the glass/Cr interfaces, respectively. The thick-
#238168 - $15.00 USD Received 16 Apr 2015; revised 8 May 2015; accepted 11 May 2015; published 15 May 2015 © 2015 OSA 18 May 2015 | Vol. 23, No. 10 | DOI:10.1364/OE.23.013693 | OPTICS EXPRESS 13696
Raw spectrum
Phase stabilization and spectral calibration
Interpolation and dispersion compensation
FFT
Surface Jones matrix extraction
Jones analysis
Retardation
Calibration data
Fig. 2. PS-iOCT data processing pipeline.
nesses of FPE1 and FPE2 were both around 6.6 mm, which resulted in two interleaved OFCs with 1020 peaks each incident on the third port of the circulator leading to the sample arm. The free spectral range and finesse of the two FPEs were around 0.1 nm and 2.6 under this setting, respectively. A clear advantage of the use of air-spaced FPEs is the ability to tune the number of desired peaks by changing the thickness of the etalons. Tuning the peaks leads to adjustment of the imaging depth range and sensitivity, as shown previously [39]. The polarization states of the combs were made orthogonal using separate polarization controllers (PC) in each arm; the two combs were then combined with a (fiber-based) polarizing beam splitter. The output powers of the two combs were maximized and balanced by adjusting PC1 and PC2, and the signal was digitized using a 1.2-GHz digitizer (ATS9360, AlazarTech, Canada). 3.
Data processing
Figure 2 outlines the main steps of the data processing framework, which are discussed in the following subsections. Data processing was slightly complicated compared to traditional OCT and iOCT processing because of the need for high phase (wavelength) stability from the light source to avoid polarization cross-talk. Moreover, to avoid the cost of a second high-speed digitizer, we developed a method to perform wavelength stabilization, spectral calibration and dispersion compensation using the single, two-channel digitizer that we had available. 3.1.
Wavelength stabilization and spectral calibration
In post-processing, each raw PS-iOCT interferogram captured on a single detection channel contains a mix of two interferograms, one each from the two input polarization states. Therefore, each raw interferogram must be demultiplexed to yield two interferograms that separately encode data from orthogonal polarization states. The demultiplexing can be done as described previously for interleaved OCT [40]. To minimize polarization cross-talk, which occurs when the demultiplexing is not done correctly, accurate resampling is needed. Furthermore, we must ensure that interferograms for each A-line are well registered in wavelength space. The use of a swept-wavelength source presents two main challenges for this task: firstly, jitter in the synchronization between the wavelength sweep and digitizer can result in a random spectral shift of the interferogram in each spectrum signal; secondly, the time-to-wavelength calibration function can vary among A-lines because of the limited repeatability of wavelength-swept sources. These two problems can be solved by implementing a technique for wavelength stabilization and spectral calibration, respectively. Others have solved these problems by incorporating an additional Mach-Zehnder interferometer to perform both spectral calibration and wavelength stabilization simultaneously [41]. However, their designs require an extra detection channel, which can significantly increase the cost and complexity of the system, as was not possible for us. Alternatively, calibration mirrors have been adopted to compensate the phase jitter between A-lines, but this method is only suitable in the case of a source with good repeatability of the wavelength sweep [33, 42]. In the absence of a second digitizer and third digitization channel, we developed a novel strategy to perform spectral calibration and wavelength stabilization for each A-scan. The calibration mirror in the sample arm of our system was illuminated with light from the 10% output
#238168 - $15.00 USD Received 16 Apr 2015; revised 8 May 2015; accepted 11 May 2015; published 15 May 2015 © 2015 OSA 18 May 2015 | Vol. 23, No. 10 | DOI:10.1364/OE.23.013693 | OPTICS EXPRESS 13697
Amplitude (A. U.)
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Fig. 3. An example A-line signal showing the two peak signals generated by the calibration mirror. The red and blue dashed lines indicate roughly the filtering windows used to extract the interferograms for wavelength stabilization and spectral calibration.
of the combined OFC. We positioned the calibration mirror so that the peak associated with reflection from the proximal surface of one of the two FPEs appeared at a depth of 50-70 pixels, which was relatively close to the zero-delay line (the whole range of each A-line was 4096 pixels). Meanwhile, the calibration mirror also generated a second signal due to interference with the second surface in the FPE; this signal appeared deeper in the A-line (around pixel 900-1200), as shown in Fig. 3. The spatial difference between these two signals was twice the thickness of the FPE cavity because of round-trip travels in the FPE cavity. Note the significant broadening of the second peak due to dispersion. The interferograms associated with the two calibration peaks were individually recovered by 1) windowing the data in the spatial domain over the areas shown in the dashed rectangles of Fig. 3 and 2) performing an inverse Fourier transform on the windowed signals. Assuming the first peak appears at a depth of zm , the thickness of the FPE is L, and the reflectivity of the calibration mirror is R, the interference signals of these two peaks in the time domain are respectively given by √ St,1 ∝ R cos(2kt (t)zm + ϕt (t)), and √ St,2 ∝ R cos(2kt (t)(zm + L) + ϕt (t)),
(1) (2)
where kt and ϕ t are the spectral calibration and dispersion functions, sampled in time as the source sweeps in wavelength. Because the dispersion function is the same for both peaks, the spectral calibration function can be obtained by taking the difference in phase between the interferograms generated from the two windowed functions [43]. This result, which is then unwrapped, is a function of wavenumber and the thickness of FPE cavity, which are fixed. All A-lines were then zero-padded and interpolated using their corresponding unwrapped phase difference function, which is directly proportional to the function that yields a linear-in-k wavelength mapping. This information also allowed us to assign the same wavelength to the same data points in all A-lines, which achieves the necessary wavelength stabilization. To verify the effectiveness of our wavelength stabilization algorithm, one of the two OFC signals, shown in Fig. 4(a), was observed as a function of time. The signal at pixels 1515-1640, which approximately corresponds to a spectral range from 7000 cm−1 to 7150 cm−1 , is shown prior to stabilization in Fig. 4(b). Figures 4(c) and 4(d) show the OFC signal in the selected range extracted from 128 continuous A-lines without and with wavelength stabilization and spectral calibration. The shift in the peak indices over time (x-axis) without spectral calibration manifests as a jagged signal in Fig. 4(c), while the clear horizontal lines in Fig. 4(d) show that the peak signals are consistently mapped to the same indices by our algorithm. The spectral #238168 - $15.00 USD Received 16 Apr 2015; revised 8 May 2015; accepted 11 May 2015; published 15 May 2015 © 2015 OSA 18 May 2015 | Vol. 23, No. 10 | DOI:10.1364/OE.23.013693 | OPTICS EXPRESS 13698
Amplitude (A. U.)
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Fig. 4. (a) The optical comb signal from a single FPE; (b) zoom image of the segment from the red box in (a); (c) and (d) evolution of the comb in time without and with wavelength alignment and spectral calibration. The bottom right images show the zoom-in of yellow rectangles marked in areas (c) and (d).
signals were well aligned and stable at all wavelengths. We found that the error in alignment over time for sets of data containing 128 OFC peaks as shown in Fig. 4(b) had a maximum error of less than one pixel. This performance ensured the two polarization states could be demultiplexed with the minimum cross-talk in downstream processing. 3.2.
Demultiplexing via interpolation
The two interferograms from orthogonal polarizations were separated by interpolating the peak indices of each FPE signal. The indices for interpolation were determined after the FPEs and calibration mirror were aligned. A calibration spectrum data containing the two mirror signals from the calibration was acquired prior to imaging and the interpolation indices were used for all subsequent measurements. Figure 5(a) shows the effectiveness of the resampling procedure: the red and blue curves show spectral signals obtained by taking the peak and valley indices, respectively, from a single optical comb signal shown as the black curve; hence, the two curves in Fig. 5(b) depict the upper and lower envelopes of the optical comb and demonstrate the efficiency with which we can separate the two polarization states generated by FPEs. The curve shown in blue in Fig. 5(b) represents cross-talk between two incident channels; however, this effect will be cancelled once we implement methods to cancel system birefringence. 3.3.
Dispersion compensation and fourier transformation
The spectrally calibrated and phase-stabilized interferograms were processed using standard Fourier domain OCT routines including zero-padding and Fourier transformation. 3.4.
Jones analysis
Since we detected separately the H and V components of the OCT signal for both orthogonal polarizations, the global Jones matrices from the FPBS to the free-space detection unit could be extracted from each pixel of the sample. After the extraction of surface Jones matrices, standard Jones analysis was implemented to extract the birefringence parameters in the sample [44, 45]. #238168 - $15.00 USD Received 16 Apr 2015; revised 8 May 2015; accepted 11 May 2015; published 15 May 2015 © 2015 OSA 18 May 2015 | Vol. 23, No. 10 | DOI:10.1364/OE.23.013693 | OPTICS EXPRESS 13699
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Fig. 5. (a) Demultiplexing of the two interferograms via resampling. (b) Extracting the peaks from the optical comb signal for a single channel yields the interferogram (red curve) for a single polarization state. Since we expect the optical comb represents a unique polarization state, we regard the red curve as signal and the blue one, obtained from the signal valleys, as cross-talk.
4. 4.1.
Results Measurement of standard samples
To validate the performance of the system using samples with known birefringence, we measured the round-trip birefringence of a quarter-wave plate for different orientations of the optic axis (OA). The OA of the QWP was rotated from 0 to π rad in increments of π /9, and we collected 128 A-scans at each position. The measurement results are shown in the red curve in Fig. 6. The mean across all measurements was 2.98 ± 0.06 rad, which is close to the expected value of 3.14 rad. We observed a minor fluctuation in the measured birefringence as the QWP was rotated, which we believe was due to an effective change in the signal-to-noise ratio during rotation of the sample [37,38]. Separately, we also measured the birefringence of a glass sample 1024 times, averaging in sets of 128 retardation measurements (blue curve). The mean across all measurements was 0.10 ± 0.02 rad, which is close to the expected value of 0 rad. 4.2.
Measurement of biomedical samples
We also performed an ex vivo measurement of chicken breast muscle to demonstrate the potential of the system for imaging biological samples. The lateral scanning range was 6 mm and corresponds to 850 A-lines. The intensity and phase retardation images are shown in Figs. 7(a) and 7(b), respectively. Fixed pattern noise associated with spurious reflections in the system was removed by subtracting the median of the complex values at each depth in each B-scan channel. The intensity image was calculated as the summation of the squared amplitude of the four Jones elements and displayed in dB scale. Based on the intensity information, the sample surface was identified using a dynamic programming segmentation algorithm [6]. The system
#238168 - $15.00 USD Received 16 Apr 2015; revised 8 May 2015; accepted 11 May 2015; published 15 May 2015 © 2015 OSA 18 May 2015 | Vol. 23, No. 10 | DOI:10.1364/OE.23.013693 | OPTICS EXPRESS 13700
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Fig. 6. Round-trip retardation measurement of a glass plate and QWP. The optic axis of the QWP was rotated between measurements, showing the stability of the measurement.
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birefringence was then canceled by using the Jones matrix information calculated for the surface, and the round-trip phase retardation at each pixel was derived from the corrected Jones matrices. The phase retardation B-scan is shown in Fig. 7(b), and clearly exhibits the banding structure that is typically observed in chicken breast by most traditional PS-OCT systems. 5.
Discussion
As demonstrated above, the penetration depth of light in the chicken breast muscle was around 1 - 1.5 mm, which is comparable to some reported PS-OCT systems [21,31] but lower than some others [33]. This suggests that the current system suffers somewhat from reduced sensitivity. We hypothesize several factors likely contribute to this decreased sensitivity, most of which stem from limitations of the digitizer configuration we had available. Firstly, the sensitivity of Fourier domain OCT is well known to be a function of the number of data points in each A-line [46–48]. The number of data points used to generate A-scans in PS-iOCT is determined by the number of peaks generated by a single OFC, which in turn is determined by the thickness of the FPE cavity. Thus, one way to enhance the sensitivity is to increase the thickness of the FPEs. In this case, the number of possible peaks will ultimately be restricted by the sampling rate of the digitizer. The frequency bandwidth of the OFC signal increases linearly with the number of peaks, and too high a peak number leads to insufficient sampling and signal decay. With our digitizer, this occurs when the number of peaks exceeds 1200. To address this issue, a faster digitizer or slower wavelength-swept light source can be
#238168 - $15.00 USD Received 16 Apr 2015; revised 8 May 2015; accepted 11 May 2015; published 15 May 2015 © 2015 OSA 18 May 2015 | Vol. 23, No. 10 | DOI:10.1364/OE.23.013693 | OPTICS EXPRESS 13701
used. Secondly, the interferograms generated by the calibration mirror contribute to a baseline signal that reduces the effective dynamic range available to measure the actual sample signal: that is, the maximum power we can measure from the sample is lowered by the additional signal contribution from the calibration mirror. This issue could be resolved by implementing the system using a digitizer with additional detection channels or by adding a second digitizer to the system. In the current work, we did not do this due to cost constraints. Instead, we developed a novel method to perform the measurements using a single two-channel digitizer. Thirdly, the power illuminating the sample is relatively low, which affects the sensitivity. The loss of power partially associated with the need for additional components to perform the wavelength stabilization and dispersion compensation. This too could be avoided with the availability of another digitizer channel. Moreover, we fabricated our own metal-coated FPEs, but the metal absorbs a significant portion of beam power, lowering its coupling efficiency. In the future, the FPEs could be replaced with commercial etalons or dielectric-coated etalons with better performance. Application of well-designed dielectric coating on the optical flats can offer lower optical power loss in FPEs as well as reduce the amount of cross-talk between the two polarization channels. 6.
Conclusion
In conclusion, we have introduced a novel scheme for SM fiber-based PS-OCT using two optical frequency combs encoding orthogonal polarization states. The proposed SM fiber-based method multiplexes polarization-encoded information by interleaving the spectra of the two polarized channels during a single sweep of a swept laser source. It therefore maintains the advantage of the robustness and flexibility of SMF MIPS PS-OCT systems for endoscopic or intravascular imaging while reducing the imaging time, cost and effects of sensitivity roll-off compared to other SMF MIPS schemes. As the polarization multiplexing is achieved by passive optics, the system is simple. Quantitative birefringence information can be extracted because the detectors allow reconstruction of the full Jones matrix of the sample. The interleaved OFC design ensures that the sensitivities of two polarization channels are equal and are not affected by roll-off, unlike depth-encoded methods. Several parameters of the system (sensitivity, imaging range, and degree of cross-talk) are flexible and may be changed by adjusting the thickness of the etalons. Finally, the system was designed for use with a single digitizer, which can save significantly on cost. One inherent disadvantage of the system, however, is that use of the etalons for polarization encoding can result in reduced optical power at the input of the system. However, modern light sources for OCT usually provide much higher power than is required for biomedical imaging, which mitigates this concern. In our system, we achieved a sensitivity of 90 dB despite having fewer points per A-scan than traditional OCT systems, which is on par with standard sensitivities in PS-OCT schemes. Higher sensitivity could be obtained by using a higher power source or increasing the number of peaks in our A-scan. A second potential disadvantage is that the iOCT strategy trades ranging depth for improved imaging speed through multiplexing. As such, the number of available points for each A-line is reduced, which can result in reduced imaging sensitivity. This issue does not pose a problem for light sources possessing a longer coherence length than the thickness of your sample, such as the VCSEL used in this work, and can be compensated by using a digitizer with a higher sampling rate. In this work, we used a digitizer with a 1.2 GHz sampling rate and managed to still have 1020 points in each A-scan and a usable ranging depth of > 6mm, which is typical of OCT systems. Finally, the current implementation uses a single, two-channel digitizer, which contributes to lower sensitivity. In the future, this can be addressed by using a different digitizer with another channel.
#238168 - $15.00 USD Received 16 Apr 2015; revised 8 May 2015; accepted 11 May 2015; published 15 May 2015 © 2015 OSA 18 May 2015 | Vol. 23, No. 10 | DOI:10.1364/OE.23.013693 | OPTICS EXPRESS 13702
Acknowledgments We thank Thomas Carver and Saara Khan for their assistance on fabrication of optical components, Hee-Yoon Lee for helpful discussion about the system setup, and members of the SBO research group for insightful discussion. Tahereh Marvdashti was funded by a Stanford Graduate Fellowship. Lian Duan was partly funded by Air Force Research Award FA9550-12-1-0269.
#238168 - $15.00 USD Received 16 Apr 2015; revised 8 May 2015; accepted 11 May 2015; published 15 May 2015 © 2015 OSA 18 May 2015 | Vol. 23, No. 10 | DOI:10.1364/OE.23.013693 | OPTICS EXPRESS 13703
Abstract: We introduce a new strategy for single-mode fiber based polarization-sensitive (PS-) optical coherence tomography (OCT) using orthogonally polarized optical frequency combs (OFC) in the sample arm. The two OFCs are tuned to be interleaved in the spectral domain, permitting simultaneous measurement of both polarization states from the same spatial region C close to the location of zero pathlength delay. The two polarization states of the beam in the sample arm are demultiplexed by interpolation after performing wavelength stabilization via a two-mirror calibration method. The system uses Jones matrix methods to measure quantitatively the round-trip phase retardation B-scans in the sample. A glass plate and quarter-wave plate were measured to validate the accuracy of the birefringence measurement. Further, we demonstrated the potential of this system for biomedical applications by measurement of chicken breast muscle. © 2015 Optical Society of America OCIS codes: (170.4500) Optical coherence tomography; (170.3880) Medical and biological imaging; (260.5430) Polarization; (260.1440) Birefringence.
References and links 1. M. R. Hee, D. Huang, E. A. Swanson, and J. G. Fujimoto, “Polarization-sensitive low-coherence reflectometer for birefringence characterization and ranging,” J. Opt. Soc. Am. B 9, 903–908 (1992). 2. J. F. de Boer and T. E. Milner, “Review of polarization sensitive optical coherence tomography and Stokes vector determination,” J. Biomed. Opt. 7, 359–371 (2002). 3. J. de Boer, Polarization Sensitive Optical Coherence Tomography (Springer, 2013). 4. E. G¨otzinger, M. Pircher, M. Sticker, A. F. Fercher, and C. K. Hitzenberger, “Measurement and imaging of birefringent properties of the human cornea with phase-resolved, polarization-sensitive optical coherence tomography,” J. Biomed. Opt. 9, 94–102 (2004). 5. E. G¨otzinger, M. Pircher, and C. K. Hitzenberger, “High speed spectral domain polarization sensitive optical coherence tomography of the human retina,” Opt. Express 13, 10217–10229 (2005). 6. L. Duan, M. Yamanari, and Y. Yasuno, “Automated phase retardation oriented segmentation of chorio-scleral interface by polarization sensitive optical coherence tomography,” Opt. Express 20, 3353–3366 (2012). 7. M. Yamanari, K. Ishii, S. Fukuda, Y. Lim, L. Duan, S. Makita, M. Miura, T. Oshika, and Y. Yasuno, “Optical rheology of porcine sclera by birefringence imaging,” PLoS ONE 7, e44026 (2012). 8. C. E. Saxer, J. F. de Boer, B. H. Park, Y. Zhao, Z. Chen, and J. S. Nelson, “High-speed fiber based polarizationsensitive optical coherence tomography of in vivo human skin,” Opt. Lett. 25, 1355–1357 (2000). 9. B. H. Park, C. Saxer, S. M. Srinivas, J. S. Nelson, and J. F. de Boer, “In vivo burn depth determination by highspeed fiber-based polarization sensitive optical coherence tomography,” J. Biomed. Opt. 6, 474–479 (2001). 10. M. C. Pierce, J. Strasswimmer, B. Hyle Park, B. Cense, and J. F. De Boer, “Birefringence measurements in human skin using polarization-sensitive optical coherence tomography,” J. Biomed. Opt. 9, 287–291 (2004). 11. L. Duan, T. Marvdashti, A. Lee, J. Y. Tang, and A. K. Ellerbee, “Automated identification of basal cell carcinoma by polarization-sensitive optical coherence tomography,” Biomed. Opt. Express 5, 3717–3729 (2014).
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Potsaid, J. Liu, V. Jayaraman, A. Cable, M. F. Kraus, K. Liang, J. Hornegger, and J. G. Fujimoto, “Depth-encoded all-fiber swept source polarization sensitive OCT,” Biomed. Opt. Express 5, 2931–2949 (2014). 30. B. Park, M. C. Pierce, B. Cense, S.-H. Yun, M. Mujat, G. Tearney, B. Bouma, and J. de Boer, “Real-time fiberbased multi-functional spectral-domain optical coherence tomography at 1.3 µm,” Opt. Express 13, 3931–3944 (2005). 31. S. Jiao and L. V. Wang, “Two-dimensional depth-resolved Mueller matrix of biological tissue measured with double-beam polarization-sensitive optical coherence tomography,” Opt. Lett. 27, 101–103 (2002). 32. S. Jiao, M. Todorovic, G. Stoica, and L. V. Wang, “Fiber-based polarization-sensitive Mueller matrix optical coherence tomography with continuous source polarization modulation,” Appl. Opt. 44, 5463–5467 (2005). 33. M. Yamanari, S. Makita, and Y. 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and local birefringence imaging,” Opt. Express 18, 854–876 (2010). 38. L. Duan, S. Makita, M. Yamanari, Y. Lim, and Y. Yasuno, “Monte-Carlo-based phase retardation estimator for polarization sensitive optical coherence tomography,” Opt. Express 19, 16330–16345 (2011). 39. H. Y. Lee, T. Marvdashti, L. Duan, S. A. Khan, and A. K. Ellerbee, “Scalable multiplexing for parallel imaging with interleaved optical coherence tomography,” Biomed. Opt. Express 5, 3192–3203 (2014). 40. H. Y. Lee, H. Sudkamp, T. Marvdashti, and A. K. Ellerbee, “Interleaved optical coherence tomography,” Opt. Express 21, 26542–26556 (2013). 41. R. Huber, M. Wojtkowski, and J. G. Fujimoto, “Fourier domain mode locking ( FDML ): A new laser operating regime and applications for optical coherence tomography,” Opt. Express 14, 1981–1983 (2006). 42. B. Vakoc, S. Yun, J. de Boer, G. Tearney, and B. Bouma, “Phase-resolved optical frequency domain imaging,” Opt. Express 13, 5483–5493 (2005). 43. S. Makita, T. Fabritius, and Y. Yasuno, “Full-range, high-speed, high-resolution 1-μ m spectral-domain optical coherence tomography using BM-scan for volumetric imaging of the human posterior eye,” Opt. Express 16, 8406–8420 (2008). 44. S. Jiao, W. Yu, G. Stoica, and L. Wang, “Optical-fiber-based Mueller optical coherence tomography,” Opt. Lett. 28, 1206–1208 (2003). 45. B. H. Park, M. C. Pierce, B. Cense, and J. F. de Boer, “Jones matrix analysis for a polarization-sensitive optical coherence tomography system using fiber-optic components,” Opt. Lett. 29, 2512–2514 (2004). 46. R. Leitgeb, C. Hitzenberger, and A. Fercher, “Performance of fourier domain vs. time domain optical coherence tomography,” Opt. Express 11, 889–894 (2003). 47. 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,” Opt. Lett. 28, 2067–2069 (2003). 48. M. Choma, M. Sarunic, C. Yang, and J. Izatt, “Sensitivity advantage of swept source and Fourier domain optical coherence tomography,” Opt. Express 11, 2183–2189 (2003).
1.
Introduction
Polarization-sensitive (PS-) optical coherence tomography (OCT) is a technique that adds birefringence contrast to the structural contrast provided by conventional OCT [1–3]. It is well known that biomedical tissues containing aligned collagen fibers exhibit strong birefringence; thus, changes in the birefringent properties of these tissues are often indicative of disease. Consequently, PS-OCT has been used to enhance disease diagnosis in several clinical areas, including ophthalmology [4–7], dermatology [8–11], dentistry [12–14], cardiology [15, 16], and cancer detection [17]. Free-space or polarization-maintaining (PM)-fiber-based PS-OCT systems can directly measure birefringence using a single incident polarization state (SIPS) [1,4,18–23]. However, freespace systems require delicate alignment and have limited utility for applications requiring flexible sample delivery, such as endoscopy, while PM-fiber-based systems can be costly, bulky and suffer from polarization cross-talk and alignment challenges [8,18,20,24,25]. Recently, singlemode (SM) fiber-based SIPS PS-OCT systems have been proposed by several groups [26, 27]. A major challenge of these systems is the need for all segments of the SM fiber need to be well fixed (e.g., using polarization controllers) to prevent polarization scrambling in the fiber; hence, such systems are not suitable for portable, handheld or endoscopic applications. In contrast, SM fiber-based PS-OCT systems that rely on illumination of the sample with multiple input polarization states (MIPS) have been successfully utilized for endoscopy [16, 28, 29]. Unfortunately, designs that exploit time-encoding of the input polarization state suffer from reduced imaging speed due to the need for sequential measurements [16, 30], and strategies invoking multiple light sources or polarization modulators incur significant additional costs [31–33]. Some have used depth-encoding to separate images from different polarization states [34–36], but these methods lead to lower sensitivity for at least one polarization channel, which in turn leads to a lower effective signal-to-noise ratio (SNR) and increased systematic and random errors in the birefringence measurement [37, 38]. The recent emergence of shortcavity swept source lasers with long coherence lengths, low sensitivity roll-off versus depth and passive delay units does mitigate some of the asymmetric sensitivity challenges in depth#238168 - $15.00 USD Received 16 Apr 2015; revised 8 May 2015; accepted 11 May 2015; published 15 May 2015 © 2015 OSA 18 May 2015 | Vol. 23, No. 10 | DOI:10.1364/OE.23.013693 | OPTICS EXPRESS 13695
PC 3
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Fig. 1. Schematic of the PS-iOCT system with orthogonally polarized optical frequency combs. FPE: Fabry-Perot etalon; PC: polarization controller; FPBS, fiber polarizing beam splitter; CIR: circulator; BS: non-polarizing beam splitter; PBS: polarizing beam splitter; G, galvanometers.
encoded systems [35,36], but these systems still suffer from sensitivity decay in the polarization channel far from the zero-delay-line. In this work, we present a new configuration for MIPS, SM-fiber PS-OCT whereby two orthogonally polarized inputs are generated in the form of interleaved optical frequency combs from a single source and interrogate the sample simultaneously (i.e., maintaining imaging speed). We thus call the method PS-interleaved OCT (PS-iOCT). In contrast with aforementioned systems, polarization multiplexing in PS-iOCT is achieved by passive optics (standard etalons) rather than active modulators (e.g., electro-optic modulator or acousto-optic modulator), making the scheme robust, economical, and immune to the sensitivity issues caused by modulated waveforms. Moreover, because of the interleaved strategy, both polarization channels appear at the same depth in the OCT image; hence, the two channels are not separately affected by the inherent sensitivity roll-off associated with depth-encoded schemes. In a proof-of-principle experiment, we demonstrate the ability of PS-iOCT to measure birefringence in calibrated and biological samples. We also demonstrate how this method can be implemented using a novel design for phase calibration using a single digitizer having only two detection channels, a potential cost savings. 2.
System design
Figure 1 shows the schematic of the PS-iOCT system. The light source is a MEMS short-cavity swept source (VCSEL, Thorlabs) with an average power of 27 mW, a center wavelength of 1310 nm, a 200-kHz sweep rate over 100 nm of full-width at half-maximum bandwidth, and > 50mm coherence length. The light is connected to an unbalanced Mach-Zehnder interferometer with a 90/10 coupler, which serves as the main interferometer for the system. The 10% arm is connected to the reference arm, while the 90% light is evenly split by a 50/50 coupler. Each portion of the latter is sent to a Fabry-Perot etalon (FPE) via a circulator to generate an optical frequency comb (OFC). The two combs emerging from this circulator then illuminate both the sample and a calibration mirror, whose use in wavelength stabilization will be described later. Light returning from the sample and reference arms are mixed and detected by a freespace polarization-sensitive detection unit. The illumination power at the sample surface was measured to be 2.4 mW. The two Fabry-Perot etalons in the sample arm both comprise a highly reflective silver mirror and an optical flat. The optical flats were uniformly coated with 35A˚ CR and 45A˚ Au, yielding reflectivities of 31% at the air/Au and 22% at the glass/Cr interfaces, respectively. The thick-
#238168 - $15.00 USD Received 16 Apr 2015; revised 8 May 2015; accepted 11 May 2015; published 15 May 2015 © 2015 OSA 18 May 2015 | Vol. 23, No. 10 | DOI:10.1364/OE.23.013693 | OPTICS EXPRESS 13696
Raw spectrum
Phase stabilization and spectral calibration
Interpolation and dispersion compensation
FFT
Surface Jones matrix extraction
Jones analysis
Retardation
Calibration data
Fig. 2. PS-iOCT data processing pipeline.
nesses of FPE1 and FPE2 were both around 6.6 mm, which resulted in two interleaved OFCs with 1020 peaks each incident on the third port of the circulator leading to the sample arm. The free spectral range and finesse of the two FPEs were around 0.1 nm and 2.6 under this setting, respectively. A clear advantage of the use of air-spaced FPEs is the ability to tune the number of desired peaks by changing the thickness of the etalons. Tuning the peaks leads to adjustment of the imaging depth range and sensitivity, as shown previously [39]. The polarization states of the combs were made orthogonal using separate polarization controllers (PC) in each arm; the two combs were then combined with a (fiber-based) polarizing beam splitter. The output powers of the two combs were maximized and balanced by adjusting PC1 and PC2, and the signal was digitized using a 1.2-GHz digitizer (ATS9360, AlazarTech, Canada). 3.
Data processing
Figure 2 outlines the main steps of the data processing framework, which are discussed in the following subsections. Data processing was slightly complicated compared to traditional OCT and iOCT processing because of the need for high phase (wavelength) stability from the light source to avoid polarization cross-talk. Moreover, to avoid the cost of a second high-speed digitizer, we developed a method to perform wavelength stabilization, spectral calibration and dispersion compensation using the single, two-channel digitizer that we had available. 3.1.
Wavelength stabilization and spectral calibration
In post-processing, each raw PS-iOCT interferogram captured on a single detection channel contains a mix of two interferograms, one each from the two input polarization states. Therefore, each raw interferogram must be demultiplexed to yield two interferograms that separately encode data from orthogonal polarization states. The demultiplexing can be done as described previously for interleaved OCT [40]. To minimize polarization cross-talk, which occurs when the demultiplexing is not done correctly, accurate resampling is needed. Furthermore, we must ensure that interferograms for each A-line are well registered in wavelength space. The use of a swept-wavelength source presents two main challenges for this task: firstly, jitter in the synchronization between the wavelength sweep and digitizer can result in a random spectral shift of the interferogram in each spectrum signal; secondly, the time-to-wavelength calibration function can vary among A-lines because of the limited repeatability of wavelength-swept sources. These two problems can be solved by implementing a technique for wavelength stabilization and spectral calibration, respectively. Others have solved these problems by incorporating an additional Mach-Zehnder interferometer to perform both spectral calibration and wavelength stabilization simultaneously [41]. However, their designs require an extra detection channel, which can significantly increase the cost and complexity of the system, as was not possible for us. Alternatively, calibration mirrors have been adopted to compensate the phase jitter between A-lines, but this method is only suitable in the case of a source with good repeatability of the wavelength sweep [33, 42]. In the absence of a second digitizer and third digitization channel, we developed a novel strategy to perform spectral calibration and wavelength stabilization for each A-scan. The calibration mirror in the sample arm of our system was illuminated with light from the 10% output
#238168 - $15.00 USD Received 16 Apr 2015; revised 8 May 2015; accepted 11 May 2015; published 15 May 2015 © 2015 OSA 18 May 2015 | Vol. 23, No. 10 | DOI:10.1364/OE.23.013693 | OPTICS EXPRESS 13697
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Fig. 3. An example A-line signal showing the two peak signals generated by the calibration mirror. The red and blue dashed lines indicate roughly the filtering windows used to extract the interferograms for wavelength stabilization and spectral calibration.
of the combined OFC. We positioned the calibration mirror so that the peak associated with reflection from the proximal surface of one of the two FPEs appeared at a depth of 50-70 pixels, which was relatively close to the zero-delay line (the whole range of each A-line was 4096 pixels). Meanwhile, the calibration mirror also generated a second signal due to interference with the second surface in the FPE; this signal appeared deeper in the A-line (around pixel 900-1200), as shown in Fig. 3. The spatial difference between these two signals was twice the thickness of the FPE cavity because of round-trip travels in the FPE cavity. Note the significant broadening of the second peak due to dispersion. The interferograms associated with the two calibration peaks were individually recovered by 1) windowing the data in the spatial domain over the areas shown in the dashed rectangles of Fig. 3 and 2) performing an inverse Fourier transform on the windowed signals. Assuming the first peak appears at a depth of zm , the thickness of the FPE is L, and the reflectivity of the calibration mirror is R, the interference signals of these two peaks in the time domain are respectively given by √ St,1 ∝ R cos(2kt (t)zm + ϕt (t)), and √ St,2 ∝ R cos(2kt (t)(zm + L) + ϕt (t)),
(1) (2)
where kt and ϕ t are the spectral calibration and dispersion functions, sampled in time as the source sweeps in wavelength. Because the dispersion function is the same for both peaks, the spectral calibration function can be obtained by taking the difference in phase between the interferograms generated from the two windowed functions [43]. This result, which is then unwrapped, is a function of wavenumber and the thickness of FPE cavity, which are fixed. All A-lines were then zero-padded and interpolated using their corresponding unwrapped phase difference function, which is directly proportional to the function that yields a linear-in-k wavelength mapping. This information also allowed us to assign the same wavelength to the same data points in all A-lines, which achieves the necessary wavelength stabilization. To verify the effectiveness of our wavelength stabilization algorithm, one of the two OFC signals, shown in Fig. 4(a), was observed as a function of time. The signal at pixels 1515-1640, which approximately corresponds to a spectral range from 7000 cm−1 to 7150 cm−1 , is shown prior to stabilization in Fig. 4(b). Figures 4(c) and 4(d) show the OFC signal in the selected range extracted from 128 continuous A-lines without and with wavelength stabilization and spectral calibration. The shift in the peak indices over time (x-axis) without spectral calibration manifests as a jagged signal in Fig. 4(c), while the clear horizontal lines in Fig. 4(d) show that the peak signals are consistently mapped to the same indices by our algorithm. The spectral #238168 - $15.00 USD Received 16 Apr 2015; revised 8 May 2015; accepted 11 May 2015; published 15 May 2015 © 2015 OSA 18 May 2015 | Vol. 23, No. 10 | DOI:10.1364/OE.23.013693 | OPTICS EXPRESS 13698
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0
Fig. 4. (a) The optical comb signal from a single FPE; (b) zoom image of the segment from the red box in (a); (c) and (d) evolution of the comb in time without and with wavelength alignment and spectral calibration. The bottom right images show the zoom-in of yellow rectangles marked in areas (c) and (d).
signals were well aligned and stable at all wavelengths. We found that the error in alignment over time for sets of data containing 128 OFC peaks as shown in Fig. 4(b) had a maximum error of less than one pixel. This performance ensured the two polarization states could be demultiplexed with the minimum cross-talk in downstream processing. 3.2.
Demultiplexing via interpolation
The two interferograms from orthogonal polarizations were separated by interpolating the peak indices of each FPE signal. The indices for interpolation were determined after the FPEs and calibration mirror were aligned. A calibration spectrum data containing the two mirror signals from the calibration was acquired prior to imaging and the interpolation indices were used for all subsequent measurements. Figure 5(a) shows the effectiveness of the resampling procedure: the red and blue curves show spectral signals obtained by taking the peak and valley indices, respectively, from a single optical comb signal shown as the black curve; hence, the two curves in Fig. 5(b) depict the upper and lower envelopes of the optical comb and demonstrate the efficiency with which we can separate the two polarization states generated by FPEs. The curve shown in blue in Fig. 5(b) represents cross-talk between two incident channels; however, this effect will be cancelled once we implement methods to cancel system birefringence. 3.3.
Dispersion compensation and fourier transformation
The spectrally calibrated and phase-stabilized interferograms were processed using standard Fourier domain OCT routines including zero-padding and Fourier transformation. 3.4.
Jones analysis
Since we detected separately the H and V components of the OCT signal for both orthogonal polarizations, the global Jones matrices from the FPBS to the free-space detection unit could be extracted from each pixel of the sample. After the extraction of surface Jones matrices, standard Jones analysis was implemented to extract the birefringence parameters in the sample [44, 45]. #238168 - $15.00 USD Received 16 Apr 2015; revised 8 May 2015; accepted 11 May 2015; published 15 May 2015 © 2015 OSA 18 May 2015 | Vol. 23, No. 10 | DOI:10.1364/OE.23.013693 | OPTICS EXPRESS 13699
Amplitude (A. U.)
1
(a)
0
Amplitude (A. U.)
7.7
1
Wavenumber [x103cm-1]
7.715
Wavenumber [x103cm-1]
7.9
(b)
0
7.4
Fig. 5. (a) Demultiplexing of the two interferograms via resampling. (b) Extracting the peaks from the optical comb signal for a single channel yields the interferogram (red curve) for a single polarization state. Since we expect the optical comb represents a unique polarization state, we regard the red curve as signal and the blue one, obtained from the signal valleys, as cross-talk.
4. 4.1.
Results Measurement of standard samples
To validate the performance of the system using samples with known birefringence, we measured the round-trip birefringence of a quarter-wave plate for different orientations of the optic axis (OA). The OA of the QWP was rotated from 0 to π rad in increments of π /9, and we collected 128 A-scans at each position. The measurement results are shown in the red curve in Fig. 6. The mean across all measurements was 2.98 ± 0.06 rad, which is close to the expected value of 3.14 rad. We observed a minor fluctuation in the measured birefringence as the QWP was rotated, which we believe was due to an effective change in the signal-to-noise ratio during rotation of the sample [37,38]. Separately, we also measured the birefringence of a glass sample 1024 times, averaging in sets of 128 retardation measurements (blue curve). The mean across all measurements was 0.10 ± 0.02 rad, which is close to the expected value of 0 rad. 4.2.
Measurement of biomedical samples
We also performed an ex vivo measurement of chicken breast muscle to demonstrate the potential of the system for imaging biological samples. The lateral scanning range was 6 mm and corresponds to 850 A-lines. The intensity and phase retardation images are shown in Figs. 7(a) and 7(b), respectively. Fixed pattern noise associated with spurious reflections in the system was removed by subtracting the median of the complex values at each depth in each B-scan channel. The intensity image was calculated as the summation of the squared amplitude of the four Jones elements and displayed in dB scale. Based on the intensity information, the sample surface was identified using a dynamic programming segmentation algorithm [6]. The system
#238168 - $15.00 USD Received 16 Apr 2015; revised 8 May 2015; accepted 11 May 2015; published 15 May 2015 © 2015 OSA 18 May 2015 | Vol. 23, No. 10 | DOI:10.1364/OE.23.013693 | OPTICS EXPRESS 13700
QWP orientation [rad]
Retardation [rad]
π
0
π
QWP
Glass 0 0
Glass measurement ID
7
Fig. 6. Round-trip retardation measurement of a glass plate and QWP. The optic axis of the QWP was rotated between measurements, showing the stability of the measurement.
(a)
(b)
0.5 mm
π
0.5 mm 0 Fig. 7. Retardation measurement of glass and QWP.
birefringence was then canceled by using the Jones matrix information calculated for the surface, and the round-trip phase retardation at each pixel was derived from the corrected Jones matrices. The phase retardation B-scan is shown in Fig. 7(b), and clearly exhibits the banding structure that is typically observed in chicken breast by most traditional PS-OCT systems. 5.
Discussion
As demonstrated above, the penetration depth of light in the chicken breast muscle was around 1 - 1.5 mm, which is comparable to some reported PS-OCT systems [21,31] but lower than some others [33]. This suggests that the current system suffers somewhat from reduced sensitivity. We hypothesize several factors likely contribute to this decreased sensitivity, most of which stem from limitations of the digitizer configuration we had available. Firstly, the sensitivity of Fourier domain OCT is well known to be a function of the number of data points in each A-line [46–48]. The number of data points used to generate A-scans in PS-iOCT is determined by the number of peaks generated by a single OFC, which in turn is determined by the thickness of the FPE cavity. Thus, one way to enhance the sensitivity is to increase the thickness of the FPEs. In this case, the number of possible peaks will ultimately be restricted by the sampling rate of the digitizer. The frequency bandwidth of the OFC signal increases linearly with the number of peaks, and too high a peak number leads to insufficient sampling and signal decay. With our digitizer, this occurs when the number of peaks exceeds 1200. To address this issue, a faster digitizer or slower wavelength-swept light source can be
#238168 - $15.00 USD Received 16 Apr 2015; revised 8 May 2015; accepted 11 May 2015; published 15 May 2015 © 2015 OSA 18 May 2015 | Vol. 23, No. 10 | DOI:10.1364/OE.23.013693 | OPTICS EXPRESS 13701
used. Secondly, the interferograms generated by the calibration mirror contribute to a baseline signal that reduces the effective dynamic range available to measure the actual sample signal: that is, the maximum power we can measure from the sample is lowered by the additional signal contribution from the calibration mirror. This issue could be resolved by implementing the system using a digitizer with additional detection channels or by adding a second digitizer to the system. In the current work, we did not do this due to cost constraints. Instead, we developed a novel method to perform the measurements using a single two-channel digitizer. Thirdly, the power illuminating the sample is relatively low, which affects the sensitivity. The loss of power partially associated with the need for additional components to perform the wavelength stabilization and dispersion compensation. This too could be avoided with the availability of another digitizer channel. Moreover, we fabricated our own metal-coated FPEs, but the metal absorbs a significant portion of beam power, lowering its coupling efficiency. In the future, the FPEs could be replaced with commercial etalons or dielectric-coated etalons with better performance. Application of well-designed dielectric coating on the optical flats can offer lower optical power loss in FPEs as well as reduce the amount of cross-talk between the two polarization channels. 6.
Conclusion
In conclusion, we have introduced a novel scheme for SM fiber-based PS-OCT using two optical frequency combs encoding orthogonal polarization states. The proposed SM fiber-based method multiplexes polarization-encoded information by interleaving the spectra of the two polarized channels during a single sweep of a swept laser source. It therefore maintains the advantage of the robustness and flexibility of SMF MIPS PS-OCT systems for endoscopic or intravascular imaging while reducing the imaging time, cost and effects of sensitivity roll-off compared to other SMF MIPS schemes. As the polarization multiplexing is achieved by passive optics, the system is simple. Quantitative birefringence information can be extracted because the detectors allow reconstruction of the full Jones matrix of the sample. The interleaved OFC design ensures that the sensitivities of two polarization channels are equal and are not affected by roll-off, unlike depth-encoded methods. Several parameters of the system (sensitivity, imaging range, and degree of cross-talk) are flexible and may be changed by adjusting the thickness of the etalons. Finally, the system was designed for use with a single digitizer, which can save significantly on cost. One inherent disadvantage of the system, however, is that use of the etalons for polarization encoding can result in reduced optical power at the input of the system. However, modern light sources for OCT usually provide much higher power than is required for biomedical imaging, which mitigates this concern. In our system, we achieved a sensitivity of 90 dB despite having fewer points per A-scan than traditional OCT systems, which is on par with standard sensitivities in PS-OCT schemes. Higher sensitivity could be obtained by using a higher power source or increasing the number of peaks in our A-scan. A second potential disadvantage is that the iOCT strategy trades ranging depth for improved imaging speed through multiplexing. As such, the number of available points for each A-line is reduced, which can result in reduced imaging sensitivity. This issue does not pose a problem for light sources possessing a longer coherence length than the thickness of your sample, such as the VCSEL used in this work, and can be compensated by using a digitizer with a higher sampling rate. In this work, we used a digitizer with a 1.2 GHz sampling rate and managed to still have 1020 points in each A-scan and a usable ranging depth of > 6mm, which is typical of OCT systems. Finally, the current implementation uses a single, two-channel digitizer, which contributes to lower sensitivity. In the future, this can be addressed by using a different digitizer with another channel.
#238168 - $15.00 USD Received 16 Apr 2015; revised 8 May 2015; accepted 11 May 2015; published 15 May 2015 © 2015 OSA 18 May 2015 | Vol. 23, No. 10 | DOI:10.1364/OE.23.013693 | OPTICS EXPRESS 13702
Acknowledgments We thank Thomas Carver and Saara Khan for their assistance on fabrication of optical components, Hee-Yoon Lee for helpful discussion about the system setup, and members of the SBO research group for insightful discussion. Tahereh Marvdashti was funded by a Stanford Graduate Fellowship. Lian Duan was partly funded by Air Force Research Award FA9550-12-1-0269.
#238168 - $15.00 USD Received 16 Apr 2015; revised 8 May 2015; accepted 11 May 2015; published 15 May 2015 © 2015 OSA 18 May 2015 | Vol. 23, No. 10 | DOI:10.1364/OE.23.013693 | OPTICS EXPRESS 13703
