Swept source optical coherence microscopy using a Fourier domain mode-locked laser

Download Swept source optical coherence microscopy using a Fourier domain mode-locked laser

Post on 30-Sep-2016

215 views

Category:

Documents

2 download

Embed Size (px)

TRANSCRIPT

  • Swept source optical coherence microscopy using a Fourier domain mode-locked laser

    Shu-Wei Huang1, Aaron D. Aguirre1, Robert A. Huber1,2, Desmond C. Adler1, and James G. Fujimoto1

    1Department of Electrical Engineering and Computer Science and Research Laboratory of Electronics, Massachusetts Institute of Technology, Cambridge, Massachusetts 02139

    2Lehrstuhl fr BioMolekulare Optik, Fakultt fr Physik, Ludwig-Maximilians-Universitt Mnchen, 80538 Mnchen, Germany

    klkla@mit.edu

    Abstract: Swept source optical coherence microscopy (OCM) enables cellular resolution en face imaging as well as integration with optical coherence tomography (OCT) cross sectional imaging. A buffered Fourier domain mode-locked (FDML) laser light source provides high speed, three dimensional imaging. Image resolutions of 1.6 m 8 m (transverse axial) with a 220 m 220 m field of view and sensitivity higher than 98 dB are achieved. Three dimensional cellular imaging is demonstrated in vivo in the Xenopus laevis tadpole and ex vivo in the rat kidney and human colon. 2007 Optical Society of America OCIS codes: (110.4500) Optical coherence tomography; (140.3600) Lasers, tunable; (170.3880) Medical and biological imaging; (180.1790) Confocal microscopy; (180.6900) Three-dimensional microscopy

    References and links 1. T. Wilson, Confocal Microscopy (Academic, London, 1990) 2. J. A. Izatt, M. R. Hee, G. M. Owen, E. A. Swanson, and J. G. Fujimoto, "Optical Coherence Microscopy in

    Scattering Media," Opt. Lett. 19, 590 (1994) 3. J. M. Schmitt, M. J. Yadlowsky, and R. F. Bonner, "Subsurface imaging of living skin with Optical

    Coherence Microscopy," Dermatology 191, 93 (1995) 4. A. D. Aguirre, P. Hsiung, T. H. Ko, I. Hartl, and J. G. Fujimoto, "High-resolution optical coherence

    microscopy for high-speed, in vivo cellular imaging," Opt. Lett. 28, 2064 (2003) 5. A. L. Clark, A. Gillenwater, R. Alizadeh-Naderi, A. K. El-Naggar, and R. Richards-Kortum, "Detection

    and diagnosis of oral neoplasia with an optical coherence microscope," J. Biomed. Opt. 9, 1271 (2004) 6. Y. H. Zhao, Z. P. Chen, C. Saxer, S. H. Xiang, J. F. de Boer, and J. S. Nelson, "Phase-resolved optical

    coherence tomography and optical Doppler tomography for imaging blood flow in human skin with fast scanning speed and high velocity sensitivity," Opt. Lett. 25, 114 (2000)

    7. V. Westphal, S. Yazdanfar, A. M. Rollins, and J. A. Izatt, "Real-time, high velocity-resolution color Doppler optical coherence tomography," Opt. Lett. 27, 34 (2002)

    8. T. Q. Xie, Z. G. Wang, and Y. T. Pan, "High-speed optical coherence tomography using fiberoptic acousto-optic phase modulation," Opt. Express 11, 3210 (2003).

    9. M. Pircher, E. Goetzinger, R. Leitgeb, and C. K. Hitzenberger, "Transversal phase resolved polarization sensitive optical coherence tomography," Phys. Med. Biol. 49, 1257 (2004)

    10. 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 (2003).

    11. M. A. Choma, M. V. Sarunic, C. H. Yang, and J. A. Izatt, "Sensitivity advantage of swept source and Fourier domain optical coherence tomography," Opt. Express 11, 2183 (2003).

    12. R. Leitgeb, C. K. Hitzenberger, and A. F. Fercher, "Performance of Fourier domain vs. time domain optical coherence tomography," Opt. Express 11, 889 (2003).

    13. M. A. Choma, A. K. Ellerbee, C. H. Yang, T. L. Creazzo, and J. A. Izatt, "Spectral-domain phase microscopy," Opt. Lett. 30, 1162 (2005).

    14. C. Joo, T. Akkin, B. Cense, B. H. Park, and J. E. de Boer, "Spectral-domain optical coherence phase microscopy for quantitative phase-contrast imaging," Opt. Lett. 30, 2131 (2005).

    15. C. Y. Xu, C. Vinegoni, T. S. Ralston, W. Luo, W. Tan, and S. A. Boppart, "Spectroscopic spectral-domain optical coherence microscopy," Opt. Lett. 31, 1079 (2006).

    #80984 - $15.00 USD Received 12 Mar 2007; revised 30 Apr 2007; accepted 2 May 2007; published 4 May 2007(C) 2007 OSA 14 May 2007 / Vol. 15, No. 10 / OPTICS EXPRESS 6210

  • 16. S. H. Yun, C. Boudoux, G. J. Tearney, and B. E. Bouma, "High-speed wavelength-swept semiconductor laser with a polygon-scanner-based wavelength filter," Opt. Lett. 28, 1981 (2003).

    17. S. H. Yun, G. J. Tearney, J. F. de Boer, N. Iftimia, and B. E. Bouma, "High-speed optical frequency-domain imaging," Opt. Express 11, 2953 (2003).

    18. R. Huber, M. Wojtkowski, J. G. Fujimoto, J. Y. Jiang, and A. E. Cable, "Three-dimensional and C-mode OCT imaging with a compact, frequency swept laser source at 1300 nm," Opt. Express 13, 10523 (2005).

    19. 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, 3225 (2006).

    20. R. Huber, D. C. Adler, and J. G. Fujimoto, "Buffered Fourier domain mode locking: unidirectional swept laser sources for optical coherence tomography imaging at 370,000 lines/s," Opt. Lett. 31, 2975 (2006).

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

    22. M. Gu, C. J. R. Sheppard, and X. Gan, "Image-Formation in a Fiberoptic Confocal Scanning Microscope," J. Opt. Soc. Am. A 8, 1755 (1991).

    23. M. Rajadhyaksha, R. R. Anderson, and R. H. Webb, "Video-rate confocal scanning laser microscope for imaging human tissues in vivo," Appl. Opt. 38, 2105 (1999).

    24. D. Adler, R. Huber, and J. G. Fujimoto, "Phase-sensitive optical coherence tomography at up to 370,000 lines per second using buffered Fourier domain mode-locked lasers," Opt. Lett. 32, 626-628 (2007).

    1. Introduction Confocal microscopy provides an optical sectioning ability in thick samples and is a powerful technique for biological and biomedical studies [1]. The sectioning ability in confocal microscopy depends heavily on tight focusing with high NA objectives and rejection of out-of-focus light with a pinhole in the conjugate focal plane. Therefore, it is vulnerable to aberration and multiple scattering. Optical coherence microscopy (OCM) improves the resistance to aberration and multiple scattering by combining high-sensitivity, coherence-gated detection with standard confocal microscopy, leading to greater image contrast and imaging depth in highly scattered tissues [2-5]. Moreover, the sectioning ability using coherence gating mainly depends on the spectrum of the light source, reducing the requirement for high NA optics and therefore leading to easier miniaturization for endoscopic studies.

    Traditionally, OCM is implemented with time domain detection [2-5], which allows high speed en face imaging in order to eliminate motion artifacts. However, time domain systems have a number of features that make the optical system undesirably complex. First, a rapid phase modulation scheme in the reference arm is required for time domain OCM systems. Several methods have been proposed, including rapid scanning optical delay lines [4], electro-optic modulators [6, 7], and acousto-optic modulators [8, 9]. All of these methods require sophisticated optical designs and specific issues such as scanner synchronization and dispersion compensation need to be carefully addressed. Second, spatial overlap between the coherence gate and confocal gate is critical to ensure optimal image quality in highly scattered tissues. Due to the inhomogeneous nature of tissues, the matching between the two gates generally can not be set a priori without real time measurement. In addition, when a fiber based endoscope system is used, any variation of stress in the fiber can introduce gate mismatch and degrade the image quality. Therefore, a feedback loop and a fast coordination algorithm are required to maintain optimal image quality during the imaging period. Finally, due to the limited field of view, usually a few hundred m, OCM itself can suffer from sampling error in clinical studies. One solution to this limitation is to use optical coherence tomography (OCT) for large scale survey and to conduct OCM only in the regions where abnormalities are detected using OCT. An imaging modality which incorporates high speed OCT and OCM is thus desirable. However, due to differences between time domain OCT and OCM system designs, it is challenging to achieve both high speed OCT and OCM in one system.

    Recent work has shown that spectral / Fourier domain detection enables OCT imaging with dramatically improved speed and sensitivity over conventional time domain detection [10-12]. Spectral / Fourier domain OCM has also been investigated [13-15]. Different from

    #80984 - $15.00 USD Received 12 Mar 2007; revised 30 Apr 2007; accepted 2 May 2007; published 4 May 2007(C) 2007 OSA 14 May 2007 / Vol. 15, No. 10 / OPTICS EXPRESS 6211

  • time domain OCM, which acquires only a single en face image, spectral / Fourier domain OCM generates an image by acquiring an entire 3D volume and rendering the en face plane. Importantly, and in contrast to OCT, there is no inherent sensitivity advantage of spectral / Fourier domain OCM compared to time domain OCM. The extremely high data rates necessary to acquire an entire volume with spectral / Fourier domain OCM result in reduction of pixel exposure / dwell time for any given pixel in the en face image, which effectively negates the sensitivity advantage. However, OCM and OCT using spectral / Fourier domain detection has the advantage of sharing the same optics in the reference arm, which makes integration of the two techniques relatively straightforward. In addition, spectral / Fourier domain detection measures optical signals from all depths simultaneously and therefore the need for coherence and confocal gate coordination is significantly relaxed. An en face OCM image at a depth matched to the confocal gate can be digitally extracted from the entire 3D dataset. Similarly, a confocal image with reduced speckle can be generated by summing the dataset in the axial direction. Disadvantages of spectral / Fourier domain detection include the need for a high-speed camera and spectrometer, with typical camera readout rates and spectrometer losses limiting the system speed to ~29,000 axial scan per second. This corresponds to >2 seconds acquisition time to generate en face images with 256256 pixels. Thus the image quality will suffer from motion artifacts when in vivo studies are performed.

    Swept source / Fourier domain detection has similar advantages to spectral / Fourier domain detection, but does not require a spectrometer and line scan camera and therefore has higher detection efficiency and higher speed. Different swept source designs have been proposed and implemented in OCT [16-20]. The Fourier domain mode-locked (FDML) laser is a new type of frequency swept laser that is especially promising for high speed OCT imaging [19, 20]. By using a cavity with a long optical delay line and scanning a high finesse, band-pass filter synchronously with the cavity round-trip time, the need to build up lasing from spontaneous emission is circumvented and very high sweep rates can be achieved. Axial scan rates up to 370 kHz have been demonstrated recently with a buffered configuration [20] and could provide OCM frame rates of ~6 Hz with 256 x 256 pixels, which is comparable to high-speed time domain OCM and is sufficient to eliminate motion artifacts for in vivo imaging. In this paper, we demonstrate swept source / Fourier domain OCM imaging with a buffered FDML laser. We demonstrate initial imaging results at ~42 kHz axial scan rates, corresponding to image acquisition times of ~1.5 seconds. Image resolutions of 1.6 m 8 m (transverse axial) are achieved and cellular level resolution imaging is demonstrated in Xenopus laevis tadpoles, rat kidney, and human colon. 2. Experimental setup Figure 1 is a schematic diagram of the swept source OCM system. The buffered FDML laser has a full tuning range of ~110 nm centered at 1290 nm and an output power of ~12 mW. 3% of the laser power is coupled to an optical spectrum analyzer (OSA) for monitoring the spectrum and a Mach-Zehnder interferometer for recalibration of time to optical frequency. The other 97% of the laser power is then split equally and delivered to a reference arm and a sample arm. A polarization controller and a neutral density filter wheel are used to set the reference arm power to obtain the optimum sensitivity performance. Although the types of glasses in the objective lens are unknown, the dispersion in the sample arm is partially balanced in the reference arm by use of SFL6 and can be compensated numerically as described in reference [21]. The sample arm has a scanning confocal microscope with a pair of closely spaced galvanometer-driven mirrors, a pair of relay lenses, and a 60 water-immersion objective. Due to high loss in the optics at this wavelength, the incident power on the sample is reduced to ~1.5 mW. It should be noted that the loss in the optics does not only reduce the illumination power but also the back-coupling efficiency. The back aperture of the objective is intentionally underfilled, such that the effective NA is ~0.35, corresponding to a theoretical confocal gate of ~16 m [22]. This provides a good compromise between transverse resolution and depth of field, such that a series of high quality, coherence gated en

    #80984 - $15.00 USD Received 12 Mar 2007; revised 30 Apr 2007; accepted 2 May 2007; published 4 May 2007(C) 2007 OSA 14 May 2007 / Vol. 15, No. 10 / OPTICS EXPRESS 6212

  • face images can be extracted from the 3D dataset. The excess laser noise is cancelled by dual-balanced detection.

    The confocal gate is measured to be ~20 m [solid line in Fig. 2(A)] by blocking the reference arm and recording the DC photodetector signal while translating a mirror in the sample arm around the focus. The slight difference from the theoretical value is a consequence of residual spherical and chromatic aberrations. The coherence gate is measured to be ~8 m [dashed line in Fig. 2(A)], which is slightly larger than the theoretical value and is the result of the non-Gaussian spectral shape, nonlinear sweep speed, and the frequency calibration errors as described in reference [19]. Figure 2(B) shows the transverse resolution of the microscope system. The smallest elements measuring 2.2 m in width on the USAF 1951 resolution target can be clearly visualized. The e-2 radius was estimated to be ~1.6 m using an edge-scan measurement method [23]. Figure 2(C) shows the measured axial point spread function on a log scale at the depth of ~450 m, which is typically the deepest penetration depth of the OCM system while preserving reasonable image quality in highly scattered tissues. This demonstrates that the sensitivi...

Recommended

View more >