Simultaneous dual-band optical coherence tomography in the spectral domain for high resolution in...

Simultaneous dual-band optical coherence tomography in the spectral domain for high

resolution in vivo imaging

Peter Cimalla,* Julia Walther, Mirko Mehner, Maximiliano Cuevas, and Edmund Koch

Department of Clinical Sensoring and Monitoring, Medical Faculty Carl Gustav Carus, University of Technology Dresden, Fetscherstrasse 74, 01307 Dresden, Germany

*[email protected]

Abstract: Optical coherence tomography (OCT) in the spectral domain is demonstrated simultaneously at two wavelength bands centered at 800 nm and 1250 nm. A novel commercial supercontinuum laser is applied as a single low coherence broadband light source. The emission spectrum of the source is shaped by optical and spatial filtering in order to achieve an adequate double peak spectrum containing the wavelength bands 700 - 900 nm and 1100 - 1400 nm for dual-band OCT imaging and thus reducing the radiation exposure of the sample. Each wavelength band is analyzed with an individual spectrometer at an A-scan rate of about 12 kHz which enables real-time imaging for the examination of moving samples. A common path optical setup optimized for both spectral regions with a separate single fiber-based scanning unit was realized which facilitates flexible handling and easy access to the measurement area. The free-space axial resolutions were measured to be less than 4.5 µm and 7 µm at 800 nm and 1250 nm, respectively. Three-dimensional imaging ten times faster than previously reported with a signal-to-noise-ratio of above 90 dB is achieved simultaneously in both wavelength bands. Spectral domain dual-band OCT combines real-time imaging with high resolution at 800 nm and enhanced penetration depth at 1250 nm and therefore provides a well suited tool for in vivo vasodynamic measurements. Further, spatially resolved spectral features of the sample are obtained by means of comparing the backscattering properties at two different wavelength bands. The ability of dual-band OCT to enhance tissue contrast and the sensitivity of this imaging modality to wavelength-dependent sample birefringence is demonstrated.

©2009 Optical Society of America

OCIS codes: (110.4500) Optical coherence tomography; (170.3880) Medical and biological imaging; (170.6510) Spectroscopy, tissue diagnostics; (170.4580) Optical diagnostics for medicine.

References and links

1. J. G. Fujimoto, “Optical coherence tomography for ultrahigh resolution in vivo imaging,” Nat. Biotechnol. 21(11), 1361–1367 (2003).

2. G. Maguluri, M. Mujat, B. H. Park, K. H. Kim, W. Sun, N. V. Iftimia, R. D. Ferguson, D. X. Hammer, T. C. Chen, and J. F. de Boer, “Three dimensional tracking for volumetric spectral-domain optical coherence tomography,” Opt. Express 15(25), 16808–16817 (2007).

3. M. Yamanari, M. Miura, S. Makita, T. Yatagai, and Y. Yasuno, “Phase retardation measurement of retinal nerve fiber layer by polarization-sensitive spectral-domain optical coherence tomography and scanning laser polarimetry,” J. Biomed. Opt. 13(1), 014013 (2008).

4. T. Xie, S. Guo, Z. Chen, D. Mukai, and M. Brenner, “GRIN lens rod based probe for endoscopic spectral domain optical coherence tomography with fast dynamic focus tracking,” Opt. Express 14(8), 3238–3246 (2006).

5. R. K. Wang, and S. Hurst, “Mapping of cerebro-vascular blood perfusion in mice with skin and skull intact by Optical Micro-AngioGraphy at 1.3 mum wavelength,” Opt. Express 15(18), 11402–11412 (2007).

#114443 - $15.00 USD Received 17 Jul 2009; revised 16 Sep 2009; accepted 8 Oct 2009; published 13 Oct 2009

(C) 2009 OSA 26 October 2009 / Vol. 17, No. 22 / OPTICS EXPRESS 19486

6. T. H. Ko, D. C. Adler, J. G. Fujimoto, D. Mamedov, V. Prokhorov, V. Shidlovski, and S. Yakubovich, “Ultrahigh resolution optical coherence tomography imaging with a broadband superluminescent diode light source,” Opt. Express 12(10), 2112–2119 (2004).

7. B. E. Bouma, G. J. Tearney, I. P. Bilinsky, B. Golubovic, and J. G. Fujimoto, “Self-phase-modulated Kerr-lens mode-locked Cr:forsterite laser source for optical coherence tomography,” Opt. Lett. 21(22), 1839–1841 (1996).

8. W. Drexler, U. Morgner, F. X. Kärtner, C. Pitris, S. A. Boppart, X. D. Li, E. P. Ippen, and J. G. Fujimoto, “In vivo ultrahigh-resolution optical coherence tomography,” Opt. Lett. 24(17), 1221–1223 (1999).

9. R. Leitgeb, W. Drexler, A. Unterhuber, B. Hermann, T. Bajraszewski, T. Le, A. Stingl, and A. Fercher, “Ultrahigh resolution Fourier domain optical coherence tomography,” Opt. Express 12(10), 2156–2165 (2004).

10. B. Povazay, K. Bizheva, A. Unterhuber, B. Hermann, H. Sattmann, A. F. Fercher, W. Drexler, A. Apolonski, W. J. Wadsworth, J. C. Knight, P. St. J. Russell, M. Vetterlein, and E. Scherzer, “Submicrometer axial resolution optical coherence tomography,” Opt. Lett. 27(20), 1800–1802 (2002).

11. A. Aguirre, N. Nishizawa, J. Fujimoto, W. Seitz, M. Lederer, and D. Kopf, “Continuum generation in a novel photonic crystal fiber for ultrahigh resolution optical coherence tomography at 800 nm and 1300 nm,” Opt. Express 14(3), 1145–1160 (2006).

12. H. Wang, and A. M. Rollins, “Optimization of dual-band continuum light source for ultrahigh-resolution optical coherence tomography,” Appl. Opt. 46(10), 1787–1794 (2007).

13. J. M. Schmitt, A. Knüttel, M. Yadlowsky, and M. A. Eckhaus, “Optical-coherence tomography of a dense tissue: statistics of attenuation and backscattering,” Phys. Med. Biol. 39(10), 1705–1720 (1994).

14. Y. Pan, and D. L. Farkas, “Noninvasive imaging of living human skin with dual-wavelength optical coherence tomography in two and three dimensions,” J. Biomed. Opt. 3(4), 446–455 (1998).

15. F. Feldchtein, V. Gelikonov, R. Iksanov, G. Gelikonov, R. Kuranov, A. Sergeev, N. Gladkova, M. Ourutina, D. Reitze, and J. Warren, “In vivo OCT imaging of hard and soft tissue of the oral cavity,” Opt. Express 3(6), 239–250 (1998).

16. M. Pircher, E. Götzinger, R. Leitgeb, A. F. Fercher, and C. K. Hitzenberger, “Speckle reduction in optical coherence tomography by frequency compounding,” J. Biomed. Opt. 8(3), 565–569 (2003).

17. J. M. Schmitt, S. H. Xiang, and K. M. Yung, “Differential absorption imaging with optical coherence tomography,” J. Opt. Soc. Am. 15(9), 2288–2296 (1998).

18. M. Pircher, E. Götzinger, R. Leitgeb, A. Fercher, and C. Hitzenberger, “Measurement and imaging of water concentration in human cornea with differential absorption optical coherence tomography,” Opt. Express 11(18), 2190–2197 (2003).

19. F. Spöler, S. Kray, P. Grychtol, B. Hermes, J. Bornemann, M. Först, and H. Kurz, “Simultaneous dual-band ultra-high resolution optical coherence tomography,” Opt. Express 15(17), 10832–10841 (2007).

20. D. Sacchet, J. Moreau, P. Georges, and A. Dubois, “Simultaneous dual-band ultra-high resolution full-field optical coherence tomography,” Opt. Express 16(24), 19434–19446 (2008).

21. S. Kray, F. Spöler, M. Först, and H. Kurz, “High-resolution simultaneous dual-band spectral domain optical coherence tomography,” Opt. Lett. 34(13), 1970–1972 (2009).

22. R. Leitgeb, C. Hitzenberger, and A. Fercher, “Performance of fourier domain vs. time domain optical coherence tomography,” Opt. Express 11(8), 889–894 (2003).

23. R. Cernat, G. M. Dobre, I. Trifanov, L. Neagu, A. Bradu, M. Hughes, and A. Gh, “Podoleanu, “Investigations of OCT imaging performance using a unique source providing several spectral wavebands,” Proc. SPIE 6847, 68470U (2008).

24. S. Yun, G. Tearney, B. Bouma, B. Park, and J. de Boer, “High-speed spectral-domain optical coherence tomography at 1.3 mum wavelength,” Opt. Express 11(26), 3598–3604 (2003).

25. N. Nassif, B. Cense, B. Park, M. Pierce, S. Yun, B. Bouma, G. Tearney, T. Chen, and J. de Boer, “In vivo high-resolution video-rate spectral-domain optical coherence tomography of the human retina and optic nerve,” Opt. Express 12(3), 367–376 (2004).

26. T. Bajraszewski, M. Wojtkowski, M. Szkulmowski, A. Szkulmowska, R. Huber, and A. Kowalczyk, “Improved spectral optical coherence tomography using optical frequency comb,” Opt. Express 16(6), 4163–4176 (2008).

27. I. K. Jang, B. E. Bouma, D. H. Kang, S. J. Park, S. W. Park, K. B. Seung, K. B. Choi, M. Shishkov, K. Schlendorf, E. Pomerantsev, S. L. Houser, H. T. Aretz, and G. J. Tearney, “Visualization of coronary atherosclerotic plaques in patients using optical coherence tomography: comparison with intravascular ultrasound,” J. Am. Coll. Cardiol. 39(4), 604–609 (2002).

28. M. Mogensen, J. B. Thomsen, L. T. Skovgaard, and G. B. E. Jemec, “Nail thickness measurements using optical coherence tomography and 20-MHz ultrasonography,” Br. J. Dermatol. 157(5), 894–900 (2007).

29. B. Park, M. Pierce, B. Cense, and J. de Boer, “Real-time multi-functional optical coherence tomography,” Opt. Express 11(7), 782–793 (2003).

30. S. Meissner, G. Müller, J. Walther, H. Morawietz, and E. Koch, “In-vivo Fourier domain optical coherence tomography as a new tool for investigation of vasodynamics in the mouse model,” J. Biomed. Opt. 14(3), 034027 (2009).

31. J. Walther, G. Mueller, H. Morawietz, and E. Koch, “Analysis of in vitro and in vivo bidirectional flow velocities by phase-resolved Doppler Fourier-domain OCT,” Sens. Actuators A Phys. in press.

32. U. Morgner, W. Drexler, F. X. Kärtner, X. D. Li, C. Pitris, E. P. Ippen, and J. G. Fujimoto, “Spectroscopic optical coherence tomography,” Opt. Lett. 25(2), 111–113 (2000).

33. D. Adler, T. Ko, P. Herz, and J. Fujimoto, “Optical coherence tomography contrast enhancement using spectroscopic analysis with spectral autocorrelation,” Opt. Express 12(22), 5487–5501 (2004).



34. A. Dubois, J. Moreau, and C. Boccara, “Spectroscopic ultrahigh-resolution full-field optical coherence microscopy,” Opt. Express 16(21), 17082–17091 (2008).

1. Introduction

In optical coherence tomography (OCT), important imaging parameters like axial resolution and penetration depth into tissue depend on the spectral properties of the applied light source [1]. The axial resolution is given by the coherence length of the light, which is inverse proportional to the spectral width of the source. The penetration depth is related to the spectral absorption and scattering properties of biological tissue and thus depends on the wavelength range of the applied light. Due to the fact that the near-infrared region shows the highest penetration depth into tissue, conventional OCT systems are commonly equipped with compact superluminescent diodes (SLDs) emitting in the spectral regions centered around 800 nm or 1300 nm. In general, imaging at shorter wavelengths around 800 nm delivers higher axial resolutions, while wavelengths around 1300 nm show an enhanced penetration depth due to less tissue scattering and absorption of endogenous chromophores like hemoglobin and melanin. Recently reported OCT systems, illuminated by a SLD, show a full-width at half maximum spectral bandwidth (FWHM) of about 50 - 80 nm which enable axial resolutions of 7 - 8 µm in the range of 800 nm [2,3] and 10 - 15 µm in the range of 1300 nm [4,5]. In order to increase axial resolution, spectrally broader low coherence light sources have to be applied. Combining SLDs of different central wavelengths to one broadband source is one suitable way for resolution improvement [6]. Alternatively, mode-locked solid state lasers like Cr:forsterite and Ti:sapphire lasers allow high resolutions of 6 µm at 1300 nm [7] down to 1 µm at 800 nm [8,9]. Supercontinuum (SC) lasers which are based on a mode-locked pump source coupled into a highly nonlinear photonic crystal fiber (PCF) emit from the visible to the infrared. Due to their spectral properties, they are suitable for OCT imaging with sub-micrometer axial resolution [10]. SC light sources generating two separate spectral bands were also introduced into OCT allowing separate scanning at 800 nm and 1300 nm with axial resolutions of 3 µm and 5 µm, respectively [11,12]. OCT imaging using two wavelength bands in combination (dual-band OCT) was demonstrated first with two sequentially working SLDs at 800 nm and 1300 nm sharing one optical setup [13]. Simultaneous dual-band OCT imaging was performed later by means of two SLDs and two interferometers used in parallel [14] as well as two SLDs with a single interferometer setup [15]. This technique allows the combination of enhanced resolution at 800 nm, increased imaging depth at 1300 nm, speckle reduction by means of frequency compounding [16] and the extraction of spatially resolved spectral sample features by means of comparing the backscattering properties of the sample at two wavelength bands. Dual-band imaging was also applied for water concentration measurements by differential absorption OCT [17,18].

High resolution simultaneous dual-band OCT was demonstrated using a commercial SC laser [19] and a halogen lamp [20], both in combination with a free-space time domain (TD) setup. Recently, high resolution OCT in the spectral domain (SD) for simultaneous imaging at 740 nm and 1300 nm was also described which employs a fiber-based setup consisting of two PCFs to connect a separate scanning unit to the SC laser source and the spectrometers [21]. The benefits of SD OCT are an enhanced sensitivity compared to TD OCT [22] and the achievable high scan speed which enables real-time examination of moving samples. In the present work, a single fiber-based SD OCT system is introduced using a novel commercial SC laser source for simultaneous three-dimensional imaging at 800 nm and 1250 nm at an A-scan rate ten times faster than previously reported. Due to the application of a high power light source, a combination of filtering techniques is used to reduce the radiation exposure of the sample to harmless levels. Furthermore, the optical setup based on a single fiber, facilitates easy and flexible access to the measurement area.



2. Experimental setup

2.1 System specification

The applied light source is a SuperK Versa Super Continuum Source (Koheras A/S, Denmark) which is based on a Nd:YAG laser pump source and provides an output spectrum from 460 nm to 2400 nm with a total optical power of 1.5 W.

SC sources are employed in the field of OCT imaging under laboratory conditions for several years now [10–12]. Commercially SC sources are available from Fianium and Koheras. The suitability of the Fianium source for time domain and spectral domain OCT imaging has already been demonstrated [19,21]. The authors, to their best knowledge, describe for the first time the use of the Koheras SC source.

From the SC source output spectrum, two separate wavelength bands are extracted with center wavelengths of 800 nm and 1250 nm and a full spectral width of 200 nm and 300 nm, resulting in a theoretical free-space axial resolution of 3.2 µm and 5.2 µm, respectively. Each band is analyzed with an individual spectrometer which contains a linear image sensor with 1024 active pixels. As a consequence, the maximum scanning depth amounts to 0.82 mm at 800 nm and 1.33 mm at 1250 nm. These system parameters were chosen in order to find a convenient trade-off between axial resolution and maximum scanning depth which is inherent in spectrometer-based OCT. In order to avoid unnecessary thermal stress to the sample as a result of infrared radiation exposure, the unused light outside of the two mentioned wavelength bands is blocked, especially the powerful pump source peak at 1064 nm is eliminated.

2.2 Optical setup

Adequate spectral shaping of the SC source’s output can be achieved by means of optical filters [23], or a combination of optical and spatial aperture filters [19]. In this work, the second method is preferred since it allows handling the entire required spectral range (700 - 1400 nm) in a single beam setup and thus facilitates subsequent fiber-coupling. Fig. 1(a) shows the applied optical setup. The SC source emits a collimated beam with a diameter of 0.8 mm. In a first step, this beam is directed to a reflective notch filter which blocks the pump source peak at 1064 nm. The reflected light is neutralized by means of a beam dump. After passing the notch filter, the beam is spectrally split up by a dispersion prism and focused by a cylindrical lens. In the focal plane of the lens the upper and lower part of the emission spectrum is blocked by an adjustable mechanical slit. The unblocked light is retro-reflected by a right angle prism behind the slit and the reflected beam travels the optical way of the incident beam backwards (Fig. 1(a), side view). Thus, the light is spectrally recombined after traversing the cylindrical lens and the dispersion prism again and the initial beam shape is reconstructed. Since the incident and the returning beam are spatially separated by means of the right angle prism, no light is directly sent back into the SC laser source. At a power level of 70% the total power output of the SC source was measured to be 845 mW. An amount of 177 mW is rejected by the notch filter and 663 mW are rejected by the notch filter in combination with the prism-lens-setup. Thus, 22% of the initial optical power is available at the exit of the prism-lens-sequence.

After passing through the prism-lens-sequence, the spectrally shaped beam is coupled into a single mode (SM) fiber via a beam splitter shown by Fig. 1(b). To further decrease the optical power that is guided on the sample, an 80:20 beam splitter is employed which reduces the power to 31 mW. Another benefit of this beam splitter is that 80% of the light returning from the scanning unit is guided to the spectrometers. The fiber with a length of 5 m leads to the scanning unit.



Fig. 1. (a) Prism-lens-sequence for spectral shaping shown in the side view and the top view: collimator of the SC source (C1), notch filter (NF) to block the pump peak at 1064 nm which is then reflected to the beam dump (BD), dispersion prism (DP), cylindrical lens (CL), aperture slit (AS) to block the upper and lower part of the spectrum and retro-reflector prism (RP). (b) Scanner head and spectrometers: beam splitter (BS1) to separate the incident and the returning beam from scanner head, collimator (C2) for fiber-coupling, scanning unit with internal collimator (C3), beam splitter (BS2), reference mirror (RM) and galvanometer scanner (GS). The returning light is sent to the spectrometers via beam expanding optics with a pinhole (PH) and a dichroic mirror (DM) to separate the wavelength bands. The bands are spectrally split up by reflective gratings (G1, G2) and focused on the corresponding Si (D1) and InGaAs (D2) line scan sensor.

A Michelson interferometer inside the scanning unit splits the incoming light into a reference beam and a sample beam and generates the interference light which contains the depth information of the sample. Lateral displacement of the sample beam by means of galvanometer mounted mirrors and a telecentric scan lens setup enables two- and three-dimensional imaging. The light returning from the scanning unit is separated from the incident light by the beam splitter and directed to a dichroic mirror which isolates the two wavelength bands from each other. Each spectral band is sent to its corresponding grating-based spectrometer. To ensure sufficient spectral resolutions, a certain amount of grating grooves has to be illuminated. Therefore, the initial beam diameter of 0.8 mm is enlarged by means of a Kepler beam expander lens setup. A pinhole located in the focal plane of the beam expander additionally blocks undesired oblique stray light from previous optical surfaces. Inside the spectrometers, the 800 nm-band is focused on a silicon (Si) linear image sensor and the 1250 nm-band is focused on an indium gallium arsenide (InGaAs) linear image sensor which achieve line scan rates of up to 12 kHz and 47 kHz, respectively. Fig. 2(a) shows the emission



spectrum of the SC source with and without spectral shaping by means of the prism-lens-sequence after being coupled into the SM fiber. The use of a 780 nm SM fiber ensures the operation in single mode over the entire required spectral range (700 - 1400 nm). The spectra were measured with a fiber-coupled prism spectrometer designed for a wavelength range of 550 - 1800 nm. Without spectral shaping (black line), the peak of the Nd:YAG pump source at 1064 nm can be clearly identified. Wavelengths above 1450 nm are attenuated either due to enhanced fiber damping or as a result of insufficient fiber-coupling in this spectral range. With spectral shaping (gray line), a double peak spectrum is created. The optical power in the obtained spectral bands amounts to 2.1 mW in the range 700 - 1040 nm and 9.4 mW in the range 1090 - 1400 nm at a SC source power level of 70% which is suitable for OCT imaging. In Fig. 2(b), the double peak spectrum is plotted in a linear scale (gray solid line: 700 - 1050 nm with fivefold magnification, gray dashed line: 1050 - 1400 nm) in combination with the measured transmission curve of the applied dichroic mirror at an angle of incidence of 45 degree (black line). It can be seen that the spectral range 1100 - 1400 nm (1250 nm-band) is transmitted while the range 700 - 950 nm is reflected. Due to the fact that only the wavelength range 700 - 900 nm (800 nm-band) is analyzed by the corresponding spectrometer (section 2.1), the spectral range 900 - 1040 nm is currently not used for OCT imaging. The optical power of this unused region amounts to 0.9 mW which corresponds to 8% of the total optical power on the sample.

2.3 System timing and synchronization

For simultaneous dual-band OCT imaging, the InGaAs line scan sensor and the galvanometer scanners are synchronized to the Si line scan sensor with an A-scan rate of 11.88 kHz reaching a B-scan rate of 23.2 Hz with 512 A-scans per B-scan. Due to the integrate-while-readout layout of the line scan sensor electronics, the exposure time can be almost as long as the A-scan period. The maximum exposure time of the Si and InGaAs line scan sensor at the current A-scan rate amounts to 84 µs and 80.2 µs, respectively which corresponds to a duty cycle of 100% and 95%. The acquisition and processing of the line scan sensor data as well as the control of the galvanometer scanners is done by means of a personal computer and custom software developed with LabVIEW (National Instruments, USA).

Fig. 2. (a) Emission spectrum of the SC laser measured after coupling into a 780 nm single mode fiber with and without spectral shaping by means of the prism-lens-sequence (the spectra were measured with a prism spectrometer at different camera settings in order to use the entire dynamic range of the image sensor). (b) Double peak spectrum after spectral shaping in a linear scale plotted in combination with the measured transmission curve of the dichroic mirror (angle of incidence: 45 degree) which separates the OCT bands 700 - 900 nm and 1100 - 1400 nm. The spectral range 900 - 1040 nm is currently not used for OCT imaging.

Both spectrometers can be operated independently as stand-alone OCT systems as well. At this point, a fivefold higher optical power in the 1250 nm-band compared to the 800 nm-band



facilitates shorter exposure times of the InGaAs sensor and enables enhanced high speed imaging with an A-scan rate of up to 47 kHz.

Fig. 3. A-scans obtained from a −14 dB reflector at different axial positions at 800 nm (a) and 1250 nm (b). For deeper positions, partial aliasing as a result of fringe frequency chirping can be clearly identified. (c) Depth dependent sensitivity loss due to finite spectrometer resolution.

The data is in a first step approximated by a parabolic function. The signals decay −13.3 dB at

800 nm and −11.7 dB at 1250 nm. (d) Measured free-space axial resolution (FWHM of A-scan

peak) at 800 nm (≤ 4.5 µm) and 1250 nm (≤ 7 µm).

3. System performance

In order to characterize the developed OCT system, several measurements concerning axial resolution, signal-to-noise-ratio and spectrometer-induced depth dependent sensitivity loss

were performed. Figure 3 shows the A-scans obtained from a glass substrate surface (−14 dB reflector) at various depths. At both wavelength bands significant partial aliasing occurs at the upper third of the scanning range as a result of fringe frequency chirping. The spectrum of the interference light is focused on the line scan sensor in the wavelength space (λ-space) and has to be transferred to the wave number space (k-space) first by means of signal processing for OCT imaging in the spectral domain. The relationship between λ-space and k-space is nonlinear. Due to this nonlinearity and further influence of the applied optics, the fringe frequency of the interference light spectrum is changing along the line scan sensor (chirping). Consequently, for increasing scanning depths and corresponding higher fringe frequencies, the Nyquist frequency of the line scan sensor will be exceeded earlier on one side of the sensor than on the other side leading to partial aliasing effects. As another result of the finite spectrometer resolution, a depth dependent sensitivity loss is observed. Between zero and



maximum scanning depth, the signals at 800 nm and 1250 nm decay −13.3 dB and −11.7 dB, respectively which is comparable to other spectrometer-based systems [24–26].

The measured free-space axial resolutions as the full-width at half maximum (FWHM) of the A-scan peak are in the range of 3.8 - 4.5 µm at 800 nm and 5.7 - 7 µm at 1250 nm which is up to 40% higher than the theoretical values. A reason for this could be an insufficient compensation of the fringe frequency chirping by means of the applied internal signal processing. The signal-to-noise-ratio (SNR) is determined between the peak of the A-scan of a mirror surface and the noise baseline. To simulate the attenuation of the optical power caused by biological tissue, a combination of neutral density (ND) filters was inserted into the sample

arm of the interferometer. The damping of the ND filters amounts to −31 dB and −29 dB in the 800 nm- and 1250 nm-band, respectively. These values have to be added to the A-scan peak-to-baseline-difference to obtain the SNR of the system. The SNR of a SD OCT system depends on various parameters like the exposure time, the optical power in the reference arm and sample arm as well as the noise characteristics of the light source and the image sensor [22]. Figure 4 shows the measured SNR at an exposure time of 84 µs at 800 nm and 80.2 µs at 1250 nm with a variable optical power in the reference arm and a fixed optical power of 0.5 µW at 800 nm and 2 µW at 1250 nm returning from the sample arm to the corresponding spectrometer. The power in the reference arm was varied by means of an adjustable mechanical aperture. For low optical powers in the reference arm, the SNR enhances with increasing reference arm power because the elevated signal level moves away from the fixed noise level of the image sensor. For higher optical powers, the SNR stagnates and even drops off when the excess noise of the light source becomes the dominant noise process. At optimum power settings in the reference arm, a SNR of 94 dB at 800 nm and 93 dB at 1250 nm at 10% of the respective maximum scanning depth was achieved.

Fig. 4. (a) A-scans at 800 nm (left scale, exposure time 84 µs) and 1250 nm (right scale, exposure time 80.2 µs) obtained from a mirror surface with attenuated sample arm power by

means of neutral density filters. The damping of the filters amounts to −31 dB and −29 dB, respectively. The SNR of the system is calculated by adding the filter’s damping to the determined difference of A-scan peak to noise baseline. (b) Measured SNR for different reference arm power (PR) settings with a fixed sample arm power (PS) of 0.5 µW at 800 nm and 2 µW at 1250 nm returning to the spectrometers. At optimum power settings a SNR of 94 dB at 800 nm and 93 dB at 1250 nm is achieved.

4. Simultaneous dual-band OCT imaging

Simultaneous dual-band imaging in an in vivo mouse model is demonstrated in Fig. 5. A B-scan stack was recorded at 800 nm and 1250 nm with a temporal resolution of 23.2 frames per second. The images show the murine saphenous artery and vein surrounded by perivascular fat tissue during the diastole, obtained by averaging the corresponding B-scans of seven consecutive heart cycles. The B-scans that represent the diastole were selected by the lowest



flow velocities found with phase-resolved Doppler analysis [31]. The pixel intensity represents the logarithmic reflectance of the corresponding sample compared to an ideal mirror.

Fig. 5. Simultaneous in vivo scans of murine saphenous artery (A), vein (V) and perivascular fat tissue (FT) at 800 nm (a) and 1250 nm (b) during the diastole, obtained by averaging the corresponding B-scans of 7 consecutive heart cycles. The scan at 1250 nm shows an increased imaging depth especially in the blood vessels while the scan at 800 nm shows a higher resolution in the superficial regions of the cross-section. (c) Frequency compounded image of (a) and (b). (d) Color-encoded differential image of (a) and (b): blue and orange represent enhanced scattering at 800 nm and 1250 nm, respectively.

As a result of the different scanning depth ranges that are imaged on the corresponding line scan sensors, the two OCT images have to be scaled in depth in order to be congruent to each other. The corresponding factor that scales the physical depth information of 800 nm to that at 1250 nm was measured to be 0.589 in air which coincides with the theoretical value of 0.614 calculated from the given system parameters (section 2.1). This scaling factor may differ among probes with different refractive index. Its value has been empirically determined for the tissue shown in Fig. 5 by evaluating the congruency of the structural information in the in vivo images at 800 nm and 1250 nm, and resulted to be 0.582. A smaller scaling factor in tissue compared to air can be explained by considering the wavelength-dependency of the refractive index. In spectral domain OCT, the axial resolution and the scanning depth are inverse proportional to the refractive index (n) of the sample which results in a reduced imaged depth range in tissue (n = 1.33) compared to air (n = 1). In tissue – like in water – the refractive index gets smaller for increasing wavelengths. As a consequence, the depth range reduction at 800 nm is higher compared to 1250 nm, resulting in a smaller scaling factor in tissue.

Another issue that has to be considered for objective image comparison, is the depth-dependent signal decay due to the finite spectrometer resolution (spectrometer roll-off) which is different in the two wavelength bands. In order to correct the signal intensities in depth, a depth-resolved gain curve was calculated using the parabolic functions which were fitted to the measured signal decay data shown in Fig. 3(c). As a side effect, the noise especially in the profound regions of the cross-section is also increased.

Comparing the images of the two wavelength bands to each other, the scan at 1250 nm shows an enhanced imaging depth especially in and underneath the blood vessels. This is consistent with the reduced tissue scattering and absorption of hemoglobin in that spectral range. On the other hand, the scan at 800 nm shows a higher resolution caused by the higher spectral range in k-space. Especially, the signal-poor smooth muscle layer of the tunica media (dark ring structure in the arterial wall) [27] can be better distinguished from the surrounding tissue.

In order to reduce the speckle noise, the frequency compounded image shown in Fig. 5(c) was calculated by pixelwise averaging of the intensity values recorded at 800 nm and 1250 nm. In this situation, scattering structures e.g. the skin covering the blood vessels, appear with



enhanced homogeneity which allows a better discrimination from the underlying fat tissue. Furthermore, this technique also combines the higher image resolution obtained at 800 nm with the enhanced penetration depth into tissue achieved at 1250 nm.

To visualize the backscattering in both spectral regions, a color-encoded differential image is calculated as shown in Fig. 5(d). The color information is obtained from the intensity difference of the scan at 1250 nm and 800 nm and the color map is adjusted in a way that orange represent enhanced scattering at 1250 nm, gray represent equal scattering in both wavelength bands and blue represent enhanced scattering at 800 nm. In order to emphasize the general structural information as well, each RGB channel of the color-encoded image is weighted with the average intensity of 800 nm and 1250 nm (frequency compounded image). Due to the fact that the spectral contrast is very sensitive to speckle noise, seven B-scans of the recorded time-resolved image stack were averaged in order to reduce the speckle noise in the blood vessels.

The high temporal resolution and the three-dimensional imaging capability of the developed OCT system is demonstrated in Fig. 6. In this figure M-scans for the 800 nm-band, the 1250 nm-band, the compounded image and the color-encoded differential image obtained from the time-resolved B-scan stack imaged in Fig. 5 are plotted in combination with a transverse and a longitudinal cross-section obtained from a three-dimensional scan of the murine saphenous artery in vivo. In all M-scans the breathing motion of the mouse can clearly be identified. At 800 nm also the heart beat is identifiable which is represented by periodical fringe washout due to high flow velocities during the systole. At 1250 nm no fringe washout is observable due to the longer wavelength and a sixfold shorter exposure time of 13.9 µs compared to 84 µs in the 800 nm-band at this measurement. With the known time step per pixel of 43.1 ms the breathing rate and the heart rate of the mouse were determined to be 73 breaths per minute and 464 beats per minute, respectively. With a three-dimensional scan the shape and the orientation of the vessel can be investigated. In this example the longitudinal cross-section of the artery can be used to measure the Doppler angle for flow velocity determination.

Fig. 6. Time- and spatial resolved simultaneous in vivo scans of the murine saphenous artery. At the red line in the frequency compounded image (a) the recorded B-scan stack is resliced and the obtained M-scans for the 800 nm-band (b), 1250 nm-band (c), compounded image (d) and color-encoded differential image (e) are plotted with threefold magnification. The signal-poor smooth muscle layer (SM) of the tunica media and the arterial lumen (L) can clearly be identified. The green arrows show the breathing motion and the red arrows indicate the heart beat which is represented by periodical fringe washout due to high flow velocities during the systole. The breathing rate amounts to 73 breaths per minute and the heart rate is 464 beats per minute. The frequency compounded images (f) and (g) show a transverse and a longitudinal cross-section of the artery obtained from a 3D scan. The longitudinal scan is useful to measure the Doppler angle for flow velocity determination.

In a second example, simultaneous dual-band imaging is demonstrated at human skin in

vivo. Figure 7 shows the superficial epidermal layers stratum corneum with the helix-shaped sweat gland ducts and the underlying stratum granulosum imaged at the tip of the small



finger. Again, the scan at 1250 nm shows an enhanced penetration depth with a better visualization of the profound stratum granulosum while the scan at 800 nm is characterized by enhanced scattering in the superficial region of the cross-section and a higher resolution expressed by finer sweat gland duct structures. The frequency compounded image shows an improved homogeneity which allows a better identification of the border between the stratum corneum and stratum granulosum. An additional contrast between these two epidermal layers is achieved due to the superposition of the intensity difference at 1250 nm and 800 nm in the color-encoded differential image. The evolvement of the spectral contrast in the third dimension can be seen in Fig. 7(e) (Media 1). Figure 8 (Media 2) shows a three-dimensional frequency compounded scan from 800 nm and 1250 nm of the human finger tip in vivo. Due to the system’s high resolution and the speckle noise reduction by means of frequency compounding, the helix-like shape of the sweat gland ducts is very well observable.

Fig. 7. Simultaneous in vivo scans of human skin at 800 nm (a) and 1250 nm (b) showing the stratum corneum (C), sweat gland ducts (D) and the stratum granulosum (G) located at the tip of the small finger. The images were obtained by averaging 5 adjacent B-scans of a three-dimensional stack resulting in a displayed cross-section thickness of 20 µm. (c) Frequency compounded image of (a) and (b). (d) Color-encoded differential image of (a) and (b): blue and orange represent enhanced scattering at 800 nm and 1250 nm, respectively. (e) Fly-through of color-encoded differential image stack (Media 1).



http://www.opticsexpress.org/viewmedia.cfm?URI=oe-17-22-19486-1



Fig. 8. Three-dimensional frequency compounded scan from 800 nm and 1250 nm of the human finger tip in vivo (Media 2). The helix-like shape of the sweat gland ducts is clearly visible.

In Fig. 9, simultaneous dual-band imaging is performed in vivo at the human fingernail. The images show the nail plate which appears as a layered structure containing horizontal homogeneous bands of varying intensity [28]. The number and thickness of these bands is different at 800 nm and 1250 nm leading to a strong contrast of the layered structure in color-encoded differential image which was also observed by others [21]. The authors think that birefringence and optical activity are responsible for this effect, in contrast to other assumptions which involve spectroscopic sample features i.e. wavelength-dependent absorption and backscattering. It was shown that the human nail plate also appears as layered structure in polarization-sensitive (PS) OCT [29]. Although the light source is not polarized, the system presented in this work is sensitive to polarization. If the polarization state of the light in the sample arm in comparison to the reference arm is changed due to sample

Fig. 9. Simultaneous in vivo scans of a human fingernail at 800 nm (a) and 1250 nm (b) showing the nail plate (N), the nail bed (B) and the cuticle (C). The images were obtained by averaging 10 adjacent B-scans of a recorded three-dimensional stack resulting in a displayed cross-section thickness of 40 µm. (c) Frequency compounded image of (a) and (b). (d) Color-encoded differential image of (a) and (b).




birefringence, the interference contrast is reduced. In a birefringent medium the phase shift ∆φ between light traveling along the fast axis and the slow axis is given by Eq. (1):

2n d

ϕ πλ

∆ ⋅∆ = (1)

where ∆n is the difference of the refractive index along the fast and the slow axis, d is the thickness of the medium and λ is the wavelength of the light. As the polarization state of the light backscattered from the sample changes periodically as a function of depth, there are discrete depth levels where the congruity of the polarization states in the reference arm and the sample arm shows a minimum. As a consequence, these depth levels appear as signal-poor layers in the nail plate. Since the polarization state returning from each dark layer is identical, the round-trip phase difference between two subsequent layers equals 2π. With the corresponding single-sided phase difference π concerning only the optical way of the light into tissue and Eq. (1), the distance d0 between the dark layers can be calculated by Eq. (2):

0

.2

dn

λ=

⋅∆ (2)

Consequently, the distance between the signal-poor layers should be proportional to the wavelength of the applied light if the imaging is mostly influenced by sample birefringence. To proof this hypothesis, d0 was measured at six different wavelength bands centered at 750 nm, 800 nm, 850 nm, 1175 nm, 1250 nm and 1325 nm by analyzing only certain pixel segments of the line scan sensors (Fig. 10). The refractive index of the nail plate was assumed

Fig. 10. Wavelength-dependent appearance of the layered structure of the human nail plate. The figure shows the OCT scan at 800 nm (a) identical to Fig. 9(a) with a selected region (red rectangle) that was imaged using 6 different wavelength bands: (b) 700 - 800 nm, (c) 750 - 850 nm, (d) 800 - 900 nm, (e) 1100 - 1250 nm, (f) 1175 - 1325 nm, (g) 1250 - 1400 nm. The white bars represent the distance between the signal-poor layers inside the nail plate.

Fig. 11. Distance between dark layer structures in the human nail plate measured at six different spectral regions centered at 750 nm, 800 nm, 850 nm, 1175 nm, 1250 nm and 1325 nm (Fig. 10) using a fixed refractive index of n = 1.47. The error bars of the data points indicate the measurement quantization error of ± 1 pixel which represents ± 4 µm. By fitting a model represented by Eq. (2) to the data points a refractive index difference of ∆n = 0.0057 is obtained for the nail plate.



to be approximately constant in the entire observed spectral range with a value of 1.47 [28]. It can be observed that with increasing wavelength d0 is enlarged (Fig. 11). Assuming that also the birefringence is approximately constant in the spectral range, the model represented by Eq. (2) with the free parameter ∆n was fitted to the data points, resulting in a refractive index difference of 0.0057. The error bars of the data points indicate the measurement quantization error of ± 1 pixel which represents ± 4 µm. The indication of a higher slope of the data points compared to the fit is in agreement with an expected reduction of the birefringence i.e. a smaller ∆n towards higher wavelengths. Optical activity caused by chiral molecules also turns the axes of polarization. However, the effect is not visible even in PS OCT because it is neutralized due to the fact that the light travels the same way in the sample forward and backward. Nevertheless, optical activity may influence the OCT image of the sample since the polarization state of the light is changing along the optical path leading to a depth-dependent impact of the sample’s birefringence.

5. Discussion

The obtained results confirm that simultaneous dual-band OCT in the spectral domain presented in this work enables real-time imaging under in vivo conditions combining the high resolution at 800 nm and enhanced penetration depth at 1250 nm. The major achievements of the presented system to previously reported work are the higher imaging speed compared to [19–21] and the feature to measure in three dimensions with high resolution compared to [14]. Thus, the system is well suited to carry out non-invasive vasodynamic measurements in the in

vivo mouse model [30]. The images of entire saphenous vessels in Fig. 5 obtained with the 1250 nm system are encouraging this wavelength band for flow measurements using phase-resolved Doppler OCT [31]. On the other hand, the higher resolution at 800 nm allows a more precise observation of the thickness of the smooth muscle layer in the arterial wall which is of interest for the investigation of the effects of hypertension. Furthermore, speckle reduction and image quality improvement by means of frequency compounding, implicit in dual-band OCT, facilitates the discrimination of different tissue types. At the same time, frequency compounding combines high image resolution with the enhanced penetration depth into tissue although the higher resolution at 800 nm is slightly reduced due to the image averaging. As a third point, dual-band OCT delivers spectral information of the sample.

The extraction of spatially resolved spectroscopic sample features by means of time-frequency analysis of the interference pattern in a single wavelength band is referred to as spectroscopic OCT (SOCT). It was shown that this technique provides contrast enhancement in comparison with conventional, solely intensity-based OCT [32–34]. Comparing the backscattering properties of a sample at two different wavelength bands is also suitable as a spectroscopic measure in OCT and has been applied to determine the water concentration in the human cornea [18] and for contrast enhancement of soft tissue performed in the oral cavity [15], the human nail fold, the rabbit eye [19,21] and the rabbit trachea [20]. The combination of this technique with the demonstrated high resolution and high speed imaging at two wavelength bands is subject of future research.

The color-encoded differential image reveals the additional spectral sample features that are obtained by dual-band OCT. In Fig. 5(d), the profound regions in and underneath the blood vessels are represented mostly by orange color tones due to the lack of intensity at 800 nm. The superficial regions of the cross-section, like the skin covering the blood vessels, are represented mainly in gray and blue color tones, which is probably a result of the enhanced tissue scattering at shorter wavelengths. A sharp distinction between different types of tissue by means of coincident structural and color information cannot be observed. The color shift shows a continuous progress from blue to gray to orange as the light propagates into tissue and the blood while the structural images show an intensity step from one type of tissue to another. This leads to the conclusion that the spectral properties of tissue change with the wavelength but the impact of this change in the presented samples is to low to generate a sharp spectroscopic contrast that is observed for example at melanocytes by others [32,34]. A similar situation is present in Fig. 7(d). Although there is an additional contrast between the



stratum corneum and the stratum granulosum due to the color representation of the differential image, this is again a result of the enhanced penetration depth of the 1250 nm wavelength band and not a proper spectroscopic feature of the sample.

The layered structure of the human nail plate shown in Fig. 9 and Fig. 10 can be better explained with sample birefringence than with wavelength-specific absorption and scattering i.e. spectroscopic sample features. This is confirmed by the good match between the measured data and the found model. Although the applied model is simple since no dispersion in the sample is considered, it is suitable to illustrate the fundamental relation between the imaging wavelength and the distance between the signal-poor layers inside the human nail plate in OCT which is somehow sensitive to polarization.

6. Conclusion

The use of a novel commercial SC laser source for dual-band OCT in the spectral domain was demonstrated. The required double peak spectrum was obtained from the source by means of spectral shaping using spatial and optical filtering. High axial resolutions better than 4.5 µm at 800 nm and 7 µm at 1250 nm could be achieved simultaneously. At an A-scan rate of about 12 kHz and an exposure time of approximately 80 µs the signal-to-noise-ratio is above 90 dB in both wavelength bands. The system allows in vivo scans with real-time resolution for imaging of fast physiological processes like ventilation and heart beat in mice. The single fiber-based setup facilitates the use of different scanner heads. It was shown that simultaneous dual-band imaging combines the benefits of high resolution at 800 nm and enhanced penetration depth into tissue at 1250 nm. Dual-band OCT has the ability to improve image quality using frequency compounding and to enhance tissue contrast by means of color-encoded representation of the differential image data. Careful interpretation concerning the spectroscopic content of this data is necessary especially in birefringent samples. Therefore, further investigations have to be carried out to evaluate the additional spectral sample features which are obtained by this technique.

Acknowledgements

The authors would like to thank Dr. Gregor Müller from the Vascular Endothelium and Microcirculation group of Prof. Morawietz (Faculty of Medicine Carl Gustav Carus, Dresden University of Technology, Germany) and Björn Fischer from the Fraunhofer Institut Zerstörungsfreie Prüfverfahren (Institutsteil Dresden, Germany) for their support and the providing of the biological samples. This project was supported by SAB (Sächsische Aufbaubank project 11261/1759) and BMBF (Bundesministerium für Bildung und Forschung, NBL 3).



Date post:	07-Oct-2016
Category:	Documents
Upload:	edmund
View:	212 times
Download:	0 times

Simultaneous dual-band optical coherence tomography in the spectral domain for high resolution in...

Documents