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Optics Express

Optics Express

  • Editor: Andrew M. Weiner
  • Vol. 21, Iss. 14 — Jul. 15, 2013
  • pp: 16648–16656
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Miniature spectrometer and beam splitter for an optical coherence tomography on a silicon chip

B. I. Akca, B. Považay, A. Alex, K. Wörhoff, R. M. de Ridder, W. Drexler, and M. Pollnau  »View Author Affiliations


Optics Express, Vol. 21, Issue 14, pp. 16648-16656 (2013)
http://dx.doi.org/10.1364/OE.21.016648


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Abstract

Optical coherence tomography (OCT) has enabled clinical applications that revolutionized in vivo medical diagnostics. Nevertheless, its current limitations owing to cost, size, complexity, and the need for accurate alignment must be overcome by radically novel approaches. Exploiting integrated optics, we assemble the central components of a spectral-domain OCT system on a silicon chip. The spectrometer comprises an arrayed-waveguide grating with 136-nm free spectral range and 0.21-nm wavelength resolution. The beam splitter is realized by a non-uniform adiabatic coupler with its 3-dB splitting ratio being nearly constant over 150 nm. With this device whose overall volume is 0.36 cm3 we demonstrate high-quality in vivo imaging in human skin with 1.4-mm penetration depth, 7.5-µm axial resolution, and a signal-to-noise ratio of 74 dB. Considering the reasonable performance of this early OCT on-a-chip system and the anticipated improvements in this technology, a completely different range of devices and new fields of applications may become feasible.

© 2013 OSA

1. Introduction

Optical coherence tomography (OCT) is a well-established optical technique in the medical sciences for acquiring micrometer-scale-resolution cross-sectional images of specimen in a non-invasive way [1

1. D. Huang, E. A. Swanson, C. P. Lin, J. S. Schuman, W. G. Stinson, W. Chang, M. R. Hee, T. Flotte, K. Gregory, C. A. Puliafito, and J. G. Fujimoto, “Optical coherence tomography,” Science 254(5035), 1178–1181 (1991). [CrossRef] [PubMed]

]. State-of-the-art OCT systems operate in the frequency domain, either with a broad-band light source and a spectrometer, called “spectral-domain OCT” (SD-OCT), or with a rapidly wavelength-tuned laser, called “swept-source OCT” (SS-OCT) [2

2. A. F. Fercher, “Optical coherence tomography - development, principles, applications,” Z. Med. Phys. 20(4), 251–276 (2010). [CrossRef] [PubMed]

]. Both systems typically contain a combination of fiber and free-space components which add to the instrument size and cost, affect its mechanical stability, and therefore require active alignment. Integrating several complex optical devices as miniturized components on a single microchip improves mechanical stability for maintenance-free operation and accesses lithographic high-volume fabrication for dramatic cost reduction and improved repeatability. Integrated systems hold promise for wider deployment of OCT technology and open new fields of applications. Despite these opportunities, OCT based upon integrated optical components has as yet yielded only preliminary results. For time-domain OCT, a microchip integrating eight parallel Michelson interferometers was demonstrated [3

3. D. Culemann, A. Knuettel, and E. Voges, “Integrated optical sensor in glass for optical coherence tomography,” IEEE J. Sel. Top. Quantum Electron. 6(5), 730–734 (2000). [CrossRef]

]; also a rapid scanning delay line with a 10-kHz line rate and 1-mm scan range was realized [4

4. E. Margallo-Balbás, M. Geljon, G. Pandraud, and P. J. French, “Miniature 10 kHz thermo-optic delay line in silicon,” Opt. Lett. 35(23), 4027–4029 (2010). [CrossRef] [PubMed]

]. A Michelson interferometer for an SS-OCT system was demonstrated with an insufficient 40-μm axial resolution and 25-dB sensitivity [5

5. G. Yurtsever, P. Dumon, W. Bogaerts, and R. Baets, “Integrated photonic circuit in silicon on insulator for Fourier domain optical coherence tomography,” Proc. SPIE 7554, 75541B (2010). [CrossRef]

]. SD-OCT imaging was achieved with an arrayed-waveguide-grating (AWG) spectrometer [6

6. M. K. Smit, “New focusing and dispersive planar component based on an optical phased array,” Electron. Lett. 24(7), 385–386 (1988). [CrossRef]

]; however, sensitivity had to be boosted by semiconductor optical amplifiers [7

7. D. Choi, H. Hiro-Oka, H. Furukawa, R. Yoshimura, M. Nakanishi, K. Shimizu, and K. Ohbayashi, “Fourier domain optical coherence tomography using optical demultiplexers imaging at 60,000,000 lines/s,” Opt. Lett. 33(12), 1318–1320 (2008). [CrossRef] [PubMed]

,8

8. D. H. Choi, H. Hiro-Oka, K. Shimizu, and K. Ohbayashi, “Spectral domain optical coherence tomography of multi-MHz A-scan rates at 1310 nm range and real-time 4D-display up to 41 volumes/second,” Biomed. Opt. Express 3(12), 3067–3086 (2012). [CrossRef] [PubMed]

]. InAs/InP quantum-dot-based waveguide photodetectors and a tunable laser source for SS-OCT systems operating around 1.7 µm were presented [9

9. Y. Jiao, B. W. Tilma, J. Kotani, R. Nötzel, M. K. Smit, S. He, and E. A. Bente, “InAs/InP(100) quantum dot waveguide photodetectors for swept-source optical coherence tomography around 1.7 µm,” Opt. Express 20(4), 3675–3692 (2012). [CrossRef] [PubMed]

,10

10. B. W. Tilma, Y. Jiao, J. Kotani, E. Smalbrugge, H. P. M. M. Ambrosius, P. J. Thijs, X. J. M. Leijtens, R. Nötzel, M. K. Smit, and E. A. J. M. Bente, “Integrated tunable quantum-dot laser for optical coherence tomography in the 1.7 µm wavelength region,” IEEE J. Quantum Electron. 48(2), 87–98 (2012). [CrossRef]

], but OCT imaging was not reported. We demonstrated SD-OCT cross-sectional imaging of a multilayered phantom with an AWG [11

11. B. I. Akca, V. D. Nguyen, J. Kalkman, N. Ismail, G. Sengo, F. Sun, T. G. van Leeuwen, A. Driessen, M. Pollnau, K. Wörhoff, and R. M. de Ridder, “Toward spectral-domain optical coherence tomography on a chip,” IEEE J. Sel. Top. Quantum Electron. 18(3), 1223–1233 (2012). [CrossRef]

,12

12. V. D. Nguyen, B. I. Akca, K. Wörhoff, R. M. de Ridder, M. Pollnau, T. G. van Leeuwen, and J. Kalkman, “Spectral domain optical coherence tomography imaging with an integrated optics spectrometer,” Opt. Lett. 36(7), 1293–1295 (2011). [CrossRef] [PubMed]

] and improved its in-depth detection range toward the level of existing bulk optical systems [13

13. B. I. Akca, L. Chang, G. Sengo, K. Wörhoff, R. M. de Ridder, and M. Pollnau, “Polarization-independent enhanced-resolution arrayed-waveguide grating used in spectral-domain optical low-coherence reflectometry,” IEEE Photon. Technol. Lett. 24, 848–850 (2012).

]. Recently, a combination of Michelson interferometer, reference arm, and directional coupler for balanced detection demonstrated SS-OCT of a phantom with an 80-dB sensitivity and 5-mm depth range in air [14

14. V. D. Nguyen, N. Weiss, W. Beeker, M. Hoekman, A. Leinse, R. G. Heideman, T. G. van Leeuwen, and J. Kalkman, “Integrated-optics-based swept-source optical coherence tomography,” Opt. Lett. 37(23), 4820–4822 (2012). [CrossRef] [PubMed]

].

Here we present an important step toward a cheap, compact, and maintenance-free SD-OCT system by integrating its central components, the beam splitter and spectrometer, on a silicon chip [Fig. 1(a)
Fig. 1 Schematic of the partially integrated SD-OCT system. (a) The complete SD-OCT set-up comprising a broadband light source, the microchip with its optical circuitry consisting of a broadband beam splitter and the spectrometer (purple plate, magnified for viewing purposes), line-scan camera, and reference and sample arms, the latter including a scanner unit. (b) Details of the integrated broadband beam splitter, a 3-dB non-uniform adiabatic coupler with specific waveguide widths of w1 = 2 µm, w2 = 1.8 µm, w3 = 1.6 µm, a waveguide separation of d = 0.8 µm, and a taper length of Ltaper = 3.5 mm. (c) A conventional AWG which, in contrast to our device, includes output channels. (d) Scanning electron microscope image of the arrayed-waveguide section of the AWG before top-cladding deposition. (e) Optical microscope image of linear tapers at the waveguide/FPR interface of the fabricated AWG spectrometer.
]. With this device whose overall volume is 0.36 cm3 which is significantly (~two orders of magnitude) smaller compared to bulky counterparts (spectrometer and splitter) we demonstrate in vivo imaging in human skin, delivering averaged cross-sections of sufficient quality for medical diagnosis. Integrated optics holds a great promise for mass-produced OCT systems with significantly reduced costs and smaller footprints which can make them affordable and accessible for wider groups of researchers.

2. Design, fabrication, and characterization

2.1 Microchip fabrication

The integrated components, coupler and AWG, were realized in silicon oxynitride (SiON) [15

15. K. Wörhoff, E. J. Klein, M. G. Hussein, and A. Driessen, “Silicon oxynitride based photonics,” in Proceedings of IEEE International Conference on Transparent Optical Networks (IEEE, 2008), pp. 266–269.

]. An 8-µm-thick oxide layer was thermally grown onto a 100-mm-diameter, <100> oriented silicon substrate. A SiON layer was deposited in an Oxford Plasmalab System 133 plasma-enhanced chemical vapor deposition (PECVD) reactor at 300°C substrate temperature, 60 W power (187.5 kHz), and 500 mTorr chamber pressure. Silane (2% SiH4 diluted in N2) and N2O served as precursors, with a relative gas-flow ratio N2O/SiH4 of 0.58, followed by thermal annealing at 1150°C for 3 h. A 500-nm-thick photoresist layer was spin-coated and patterned by standard lithography and a development step. Single-mode channel waveguides with a 1.8-µm width, 1-µm height, 1.54 core refractive index, and 1.4485 cladding refractive index, enabling bending radii down to 0.5 mm, were etched in a Plasma Therm 790 reactive-ion etching reactor applying a CHF3/O2 gas mixture (100 sccm / 2 sccm) at 28 mTorr pressure, 350 W plasma power, and 20°C substrate temperature. After removing the photoresist, growth of a 1-µm-thick oxide layer by low-pressure chemical vapor deposition using tetraethyl orthosilicate as precursor and a 3-µm-thick PECVD oxide layer were each followed by thermal annealing at 1150°C for 3 h. The device size is 3.0 × 1.2 × 0.1 cm3.

2.2 Broadband non-uniform adiabatic coupler

A directional coupler, consisting of two coupled parallel uniform waveguides, exhibits a wavelength-dependent coupling ratio [Fig. 2(a)
Fig. 2 Wavelength dependence of the integrated beam splitter. (a) Simulated transmission of a uniform, i.e., intrinsically highly wavelength-dependent coupler (dashed lines) and the non-uniform adiabatic coupler in the ideal case, i.e., without fabrication imperfections, and including incomplete etching of the SiON layer in the gap region. (b) Measured transmission of the fabricated non-uniform adiabatic coupler. An oscillatory behavior can be observed, with an excess loss of 2.5 dB.
, “uniform”], making it unsuitable for use as an integrated beam splitter in broad-band OCT applications. This wavelength dependence can be significantly reduced by use of a non-uniform adiabatic coupler with linearly tapered waveguides [16

16. W. H. Louisell, “Analysis of the single tapered mode coupler,” Bell Syst. Tech. J. 33, 853–870 (1955).

], see Fig. 1(b). At their input, both the fundamental (even) and first-order (odd) mode of the composite structure are excited by light from either of the two isolated waveguides. If the propagation constants of the isolated waveguides are sufficiently different, the power in each of them predominantly couples to only one system mode.

We designed a linearly tapered directional coupler [Fig. 1(b)]. The coupler was simulated as a two-dimensional five-layer slab using the beam propagation method (Phoenix, OptoDesigner), showing a weak wavelength dependence over a range larger than 150 nm [Fig. 2(a), “ideal”]. In the straight coupler section, the waveguides were tapered adiabatically, resulting in negligible power conversion between the system modes, from w1 = 2 µm down to w3 = 1.8 µm and from w2 = 1.6 µm up to w3 = 1.8 µm, respectively, over a length of Ltaper = 3.5 mm. The gap between waveguides was d = 0.8 µm.

TE-polarized light from a supercontinuum source (Fianium SC450) was coupled to the input waveguide of the fabricated coupler by a polarization-maintaining single-mode fiber. The output signal was fiber-coupled to an optical spectrum analyzer (iHR 550, Horiba Jobin Yvon) and analyzed. To eliminate fiber-to-chip coupling losses, the transmission response was normalized to a straight waveguide. The fabricated coupler showed a coupling ratio exhibiting an oscillatory wavelength dependence with a splitting ratio of ± 1.76 dB and 2.5 dB excess loss [Fig. 2(b)], evoked by incomplete etching of the gap between the waveguides. Simulation of this fabrication error qualitatively reproduced this behavior [Fig. 2(a), “imperfect”]. Since the light fractions traveling via the reference and sample arm both pass twice through the coupler, once in “cross”- and once in “bar”-direction, the ripple in the splitting ratio results in a variation of the signal strength over the full bandwidth of < 0.2 dB at the detector array.

2.3 Arrayed-waveguide grating for enhanced depth range

In an AWG spectrometer light from an input waveguide horizontally diverges in a planar-waveguide free-propagation region (FPR) and illuminates the input facets of a number M of arrayed waveguides [Fig. 1(c)]. A constant path-length increment ΔL between adjacent arrayed waveguides causes a phase gradient across their outputs, depending linearly on frequency. These outputs, arranged on a circle, excite approximately cylindrical wavefronts in the second FPR, focusing different wavelengths onto different output waveguides. As the integrated spectrometer of the miniaturized SD-OCT system, an AWG operating at a center wavelength of λcAWG = 1250 nm, with a large FSR of 136 nm and a high overall wavelength resolution of δλ = 0.21 nm was designed by choosing [17

17. M. K. Smit and C. van Dam, “PHASAR-based WDM-devices: Principles, design and applications,” IEEE J. Sel. Top. Quantum Electron. 2(2), 236–250 (1996). [CrossRef]

] a grating order m = 9 and ΔL = 7.6 µm. Linear tapers at the waveguide/FPR interfaces [Fig. 1(e)] diminish the excess loss to 1.5 dB. A scanning electron microscope image of the arrayed waveguides before the top oxide deposition is shown in Fig. 1(d).

The depth detection range in SD-OCT is limited by the wavelength resolution of its spectrometer [2

2. A. F. Fercher, “Optical coherence tomography - development, principles, applications,” Z. Med. Phys. 20(4), 251–276 (2010). [CrossRef] [PubMed]

]. In conventional AWGs [Fig. 1(c)] this resolution is determined by the number of output channels, i.e. sampling points, which is typically limited to a few hundreds [6

6. M. K. Smit, “New focusing and dispersive planar component based on an optical phased array,” Electron. Lett. 24(7), 385–386 (1988). [CrossRef]

]. By imaging the AWG output focal plane directly onto a camera [Fig. 1(a)], the number of sampling points increases to the number of camera pixels that are addressed. In this way the additional signal roll-off due to discretely located output channels of the AWG spectrometer can be eliminated [11

11. B. I. Akca, V. D. Nguyen, J. Kalkman, N. Ismail, G. Sengo, F. Sun, T. G. van Leeuwen, A. Driessen, M. Pollnau, K. Wörhoff, and R. M. de Ridder, “Toward spectral-domain optical coherence tomography on a chip,” IEEE J. Sel. Top. Quantum Electron. 18(3), 1223–1233 (2012). [CrossRef]

]. The overall spectral resolution of the imaging system is given as the combination of the detector-limited and the arrayed-waveguide-limited resolution [13

13. B. I. Akca, L. Chang, G. Sengo, K. Wörhoff, R. M. de Ridder, and M. Pollnau, “Polarization-independent enhanced-resolution arrayed-waveguide grating used in spectral-domain optical low-coherence reflectometry,” IEEE Photon. Technol. Lett. 24, 848–850 (2012).

]. With a × 20 objective lens the usable light-source bandwidth of 75 nm, centered at λcSource ~1320 nm, was imaged onto the entire 1024 detector pixels, resulting in a wavelength spacing of 0.073 nm/pixel. Combined with the Sparrow-criterion-based arrayed-waveguide-limited resolution [18

18. C. M. Sparrow, “On spectroscopic resolving power,” Astrophys. J. 44, 76–86 (1916). [CrossRef]

,19

19. H. Takahashi, S. Suzuki, K. Kato, and I. Nishi, “Arrayed-waveguide grating for wavelength division multi/demultiplexer with nanometer resolution,” Electron. Lett. 26(2), 87–88 (1990). [CrossRef]

] of Δλ = λcAWG/1.26Mm = 0.22 nm, where M = 510, the overall spectral resolution [20

20. Z. Hu, Y. Pan, and A. M. Rollins, “Analytical model of spectrometer-based two-beam spectral interferometry,” Appl. Opt. 46(35), 8499–8505 (2007). [CrossRef] [PubMed]

] was calculated as δλ = 0.225 nm, theoretically resulting in a depth range of 1.9 mm in air, as compared to 0.5 mm with output channels. The free spectral range (FSR) can be extended and simultaneously δλ enhanced by cascading several AWGs [21

21. K. Takada, H. Yamada, and K. Okamoto, “320-channel multiplexer consisting of a 100 GHz-spaced parent AWG and 10 GHz-spaced subsidiary AWGs,” Electron. Lett. 35(10), 824–826 (1999). [CrossRef]

,22

22. B. I. Akca, C. R. Doerr, G. Sengo, K. Wörhoff, M. Pollnau, and R. M. de Ridder, “Broad-spectral-range synchronized flat-top arrayed-waveguide grating applied in a 225-channel cascaded spectrometer,” Opt. Express 20(16), 18313–18318 (2012). [CrossRef] [PubMed]

], thereby significantly improving the axial resolution and depth range in OCT.

3. OCT Performance

3.1 Partially integrated SD-OCT system

The integrated device was interfaced to the light source, external reference arm, two-dimensional scanning system, and linescan camera [Fig. 1(a)] and the system performance was analyzed. The SNR, axial resolution, and imaging range were measured by placing a mirror in the sample arm, which was moved for depth-ranging measurements. An in-tissue depth range of 1.4 mm and axial resolution of 7.5 µm was achieved. Close to the zero delay where both interferometer arms are equal in length, a SNR of 74 dB was measured for 0.5 mW of optical power on the sample.

3.2 In vivo imaging with a partially integrated OCT system

The feasibility of on-chip OCT was demonstrated by in vivo imaging of human skin above the proximal interphalangeal joint of the middle finger by applying contact gel to the imaging sites as an index-matching medium to decrease the surface reflectivity. The averaged en-face and corresponding cross-sectional tomograms are shown in Fig. 3
Fig. 3 Images of glabrous skin at interdigital joint taken using the partially integrated SD-OCT system. En face section at (a) the deeper epidermal layers featuring the living epidermis on top of the dermal papillae (yellow), (b) rete subpapillare where fibrous components dominate the basis of the dermal papillae (orange) and (c) the deeper dermis with vessels (violet). (d) Cross-section as indicated by the dotted white line in the en face sections. Colored indicators depict the location of the en face views.
, with corresponding regions indicated by colored marks or dashed lines, respectively. The morphology of several layers including the stratum corneum and uncornified layers of the epidermis [Fig. 3(a)], stratum papillare and reticulare of the dermis [Fig. 3(b)], and deeper dermis featuring vessels [Fig. 3(c)] has been visualized within a 1.4-mm depth range, sufficient for performing diagnostic procedures due to the fact that most skin malignancies are of epidermal origin [24

24. E. Fuchs and S. Raghavan, “Getting under the skin of epidermal morphogenesis,” Nat. Rev. Genet. 3(3), 199–209 (2002). [CrossRef] [PubMed]

]. In pigmented thin skin (Fig. 4
Fig. 4 Images of pigmented thin skin taken using the partially integrated SD-OCT system. (a-c) Cross-sectional views of a three-dimensional volume obtained at a location with increased melanin concentration. (f-h) En face views at different depths. (e) Three-dimensional volume-rendered representation of OCT image data. The yellow markers delineate the corresponding positions of the en face views.
), additionally to their bright appearance in OCT above the skin, hairs are identified as dark, almost vertical channels. The layered structure of the pigmented thin skin is well visualized throughout the cross-sectional series and a three-dimensional volume-rendered representation is depicted in Fig. 4(e).

3.3 Performance comparison of on-chip with bulk system

Performance of the partially integrated system was compared to a 1320-nm fiber-based, custom-designed SD-OCT system [25

25. A. Alex, B. Povazay, B. Hofer, S. Popov, C. Glittenberg, S. Binder, and W. Drexler, “Multispectral in vivo three-dimensional optical coherence tomography of human skin,” J. Biomed. Opt. 15(2), 026025 (2010). [CrossRef] [PubMed]

] with a conventional spectrometer by acquiring OCT images of scar tissue utilizing the same light source, detection system, patient interface, sample, and processing method (Fig. 5
Fig. 5 Cross-sectional tomograms of the scar tissue at the index finger. Images taken with 32 × average using (a) the 1300-nm custom-designed SD-OCT system and (b) the partially integrated SD-OCT system. For the latter image, the zero delay is offset (ZD-offset) by ~800 μm toward the region of interest to compensate for ~10 dB loss caused by signal roll-off. (c) The signal roll-off curves of the partially integrated OCT system using an AWG spectrometer (blue dashed line) and the custom-designed OCT system using a conventional spectrometer (black solid line). The resulting signal roll-off of the partially integrated OCT system after compensation is given by the blue solid line.
). The axial resolution of the fiber-based system was ~5.7 µm in tissue, slightly better than the 7.5-μm resolution of the on-chip system. However, the 74-dB SNR of the on-chip system lags behind the 94-dB value of the fiber-based system. The theoretical shot-noise-limited SNR [26

26. J. A. Izatt and M. A. Choma, “Theory of optical coherence tomography,” in Optical Coherence Tomography: Technology and Applications, W. Drexler and J. G. Fujimoto, eds. (Springer, Berlin, New York, 2008), pp. 47–72.

] of both devices is 107 dB due to the shared camera with >70% quantum efficiency and 13.9 µs exposure time at 0.5 mW optical power incident on the sample. Figure 5(a) shows the OCT image taken with 32 × averages using the fiber-based system. The zero delay of the on-chip system was offset by ~800 μm toward the region of interest to compensate for ~10 dB loss caused by signal roll-off. The resulting OCT image, also taken with 32 × averages, is shown in Fig. 5(b). A signal roll-off of 16 dB was measured at 1.4 mm depth range. The signal roll-off curves obtained by measuring the signal strength of a known reflector at different locations of the on-chip (blue dashed line) and fiber-based (black solid line) system are compared in Fig. 5(c). The resulting effective signal roll-off of the on-chip system after compensation, i.e., shifting its zero delay by ~800 μm and averaging of 32 frames (blue solid line), comes close to the signal roll-off of the fiber-based system.

3.4 Current imperfections and possible improvements

The obtained volumetric images of human skin in multiple locations and subjects demonstrate that our partially integrated SD-OCT system with a 7.5-µm axial resolution, 1.4-mm depth range, and 74-dB SNR is capable of acquiring high-quality in vivo OCT images of human skin. Despite the fact that this is the first demonstration of in vivo imaging with an on-chip OCT system, the performance parameters come surprisingly close to a state-of-the-art bulk OCT system; nevertheless, there is significant room for improvement.

Though 74-dB SNR is sufficient for biomedical imaging, in principle it can be improved to 94 dB by diminishing the non-adiabatic-coupler loss from 2.5 dB to its theoretical limit of 0 dB with an optimized lithography and etching procedure, fiber-to-chip coupling losses from 3 dB to < 0.5 dB by spot-size converters [27

27. M. M. Spühler, B. J. Offrein, G. Bona, R. Germann, I. Massarek, and D. Erni, “A very short planar silica spot-size converter using a nonperiodic segmented waveguide,” J. Lightwave Technol. 16(9), 1680–1685 (1998). [CrossRef]

], e.g. tapered input waveguides [28

28. O. Mitomi, K. Kasaya, and H. Miyazawa, “Design of a single-mode tapered waveguide for low-loss chip-to-fiber coupling,” IEEE J. Quantum Electron. 30(8), 1787–1793 (1994). [CrossRef]

], objective-lens transmission loss from 1.5 dB to < 1 dB by use of near-infrared objective lenses, and AWG excess loss from 1.5 dB to < 0.5 dB by applying vertical tapers at the arrayed-waveguide/FPR interfaces [29

29. A. Sugita, A. Kaneko, K. Okamoto, M. Itoh, A. Himeno, and Y. Ohmori, “Very low insertion loss arrayed-waveguide grating with vertically tapered waveguides,” IEEE Photon. Technol. Lett. 12(9), 1180–1182 (2000). [CrossRef]

].

The AWG and coupler were designed to operate at a 1250-nm center wavelength, whereas later a 1320-nm light source was chosen for the experiments, thereby degrading the coupler splitting ratio by 5%, while slightly improving the depth range by ~50 µm due to the longer operating wavelength, despite the reduced AWG-limited resolution of 0.29 nm.

At SiON/air interfaces, waveguides were tapered from 1.8 µm to 3 µm to efficiently couple light to the chip. The specific taper design caused a wavelength-dependent destructive interference, thereby reducing the 100-nm bandwidth of the light source to a usable bandwidth of 75 nm (with a Gaussian-like spectrum), consequently increasing the axial resolution in tissue from 5 µm to 7.5 µm. An improved taper design can avoid this effect.

The internal scattering caused by waveguide surface roughness and back-reflections at SiON/air interfaces resulted in enlarged background in the measurements, which limits the usable dynamic range of the detector. Polishing the waveguide facets at an 8° angle reduced, albeit could not eliminate these back-reflections. Scattering can be diminished with a higher-quality e-beam mask and optimized etching procedure.

4. Conclusions

In vivo imaging with a partially integrated SD-OCT system has been demonstrated by acquiring volumetric images of human skin in multiple locations and subjects. In-tissue axial resolution of 7.5 µm and in-tissue depth range of 1.4 mm was achieved. Current performance of the partially integrated OCT system can be further improved to the level of bulky commercial OCT systems with an optimized design and high-quality fabrication facilities. Considering the ability of lithography to mass-produce optimized optical systems, full on-chip integration of a complete OCT instrument will ultimately allow for significantly lower fabrication costs and pave the road toward a much wider distribution of OCT systems and new fields of applications.

Acknowledgments

This work was financially supported by the Smart Mix Program of the Netherlands Ministry of Economic Affairs and the Netherlands Ministry of Education, Culture and Science, Medical University Vienna, and the European Union projects FUN OCT (FP7 HEALTH, contract no. 201880) and FAMOS (FP7 ICT, contract no. 317744). The authors thank Anton Hollink, Henk van Wolferen, Milorad Jevremovic, Bernhard Rosenauer, and Angelika Unterhuber for technical support, Gabriel Sengo for device fabrication, Dimitri Geskus and XiOPhotonics for angle-polishing the device, Gunay Yurtsever and Alfred Driessen for fruitful discussions, and the Vereniging voor Biofysica en Biomedische Technologie for providing a travel grant.

References and links

1.

D. Huang, E. A. Swanson, C. P. Lin, J. S. Schuman, W. G. Stinson, W. Chang, M. R. Hee, T. Flotte, K. Gregory, C. A. Puliafito, and J. G. Fujimoto, “Optical coherence tomography,” Science 254(5035), 1178–1181 (1991). [CrossRef] [PubMed]

2.

A. F. Fercher, “Optical coherence tomography - development, principles, applications,” Z. Med. Phys. 20(4), 251–276 (2010). [CrossRef] [PubMed]

3.

D. Culemann, A. Knuettel, and E. Voges, “Integrated optical sensor in glass for optical coherence tomography,” IEEE J. Sel. Top. Quantum Electron. 6(5), 730–734 (2000). [CrossRef]

4.

E. Margallo-Balbás, M. Geljon, G. Pandraud, and P. J. French, “Miniature 10 kHz thermo-optic delay line in silicon,” Opt. Lett. 35(23), 4027–4029 (2010). [CrossRef] [PubMed]

5.

G. Yurtsever, P. Dumon, W. Bogaerts, and R. Baets, “Integrated photonic circuit in silicon on insulator for Fourier domain optical coherence tomography,” Proc. SPIE 7554, 75541B (2010). [CrossRef]

6.

M. K. Smit, “New focusing and dispersive planar component based on an optical phased array,” Electron. Lett. 24(7), 385–386 (1988). [CrossRef]

7.

D. Choi, H. Hiro-Oka, H. Furukawa, R. Yoshimura, M. Nakanishi, K. Shimizu, and K. Ohbayashi, “Fourier domain optical coherence tomography using optical demultiplexers imaging at 60,000,000 lines/s,” Opt. Lett. 33(12), 1318–1320 (2008). [CrossRef] [PubMed]

8.

D. H. Choi, H. Hiro-Oka, K. Shimizu, and K. Ohbayashi, “Spectral domain optical coherence tomography of multi-MHz A-scan rates at 1310 nm range and real-time 4D-display up to 41 volumes/second,” Biomed. Opt. Express 3(12), 3067–3086 (2012). [CrossRef] [PubMed]

9.

Y. Jiao, B. W. Tilma, J. Kotani, R. Nötzel, M. K. Smit, S. He, and E. A. Bente, “InAs/InP(100) quantum dot waveguide photodetectors for swept-source optical coherence tomography around 1.7 µm,” Opt. Express 20(4), 3675–3692 (2012). [CrossRef] [PubMed]

10.

B. W. Tilma, Y. Jiao, J. Kotani, E. Smalbrugge, H. P. M. M. Ambrosius, P. J. Thijs, X. J. M. Leijtens, R. Nötzel, M. K. Smit, and E. A. J. M. Bente, “Integrated tunable quantum-dot laser for optical coherence tomography in the 1.7 µm wavelength region,” IEEE J. Quantum Electron. 48(2), 87–98 (2012). [CrossRef]

11.

B. I. Akca, V. D. Nguyen, J. Kalkman, N. Ismail, G. Sengo, F. Sun, T. G. van Leeuwen, A. Driessen, M. Pollnau, K. Wörhoff, and R. M. de Ridder, “Toward spectral-domain optical coherence tomography on a chip,” IEEE J. Sel. Top. Quantum Electron. 18(3), 1223–1233 (2012). [CrossRef]

12.

V. D. Nguyen, B. I. Akca, K. Wörhoff, R. M. de Ridder, M. Pollnau, T. G. van Leeuwen, and J. Kalkman, “Spectral domain optical coherence tomography imaging with an integrated optics spectrometer,” Opt. Lett. 36(7), 1293–1295 (2011). [CrossRef] [PubMed]

13.

B. I. Akca, L. Chang, G. Sengo, K. Wörhoff, R. M. de Ridder, and M. Pollnau, “Polarization-independent enhanced-resolution arrayed-waveguide grating used in spectral-domain optical low-coherence reflectometry,” IEEE Photon. Technol. Lett. 24, 848–850 (2012).

14.

V. D. Nguyen, N. Weiss, W. Beeker, M. Hoekman, A. Leinse, R. G. Heideman, T. G. van Leeuwen, and J. Kalkman, “Integrated-optics-based swept-source optical coherence tomography,” Opt. Lett. 37(23), 4820–4822 (2012). [CrossRef] [PubMed]

15.

K. Wörhoff, E. J. Klein, M. G. Hussein, and A. Driessen, “Silicon oxynitride based photonics,” in Proceedings of IEEE International Conference on Transparent Optical Networks (IEEE, 2008), pp. 266–269.

16.

W. H. Louisell, “Analysis of the single tapered mode coupler,” Bell Syst. Tech. J. 33, 853–870 (1955).

17.

M. K. Smit and C. van Dam, “PHASAR-based WDM-devices: Principles, design and applications,” IEEE J. Sel. Top. Quantum Electron. 2(2), 236–250 (1996). [CrossRef]

18.

C. M. Sparrow, “On spectroscopic resolving power,” Astrophys. J. 44, 76–86 (1916). [CrossRef]

19.

H. Takahashi, S. Suzuki, K. Kato, and I. Nishi, “Arrayed-waveguide grating for wavelength division multi/demultiplexer with nanometer resolution,” Electron. Lett. 26(2), 87–88 (1990). [CrossRef]

20.

Z. Hu, Y. Pan, and A. M. Rollins, “Analytical model of spectrometer-based two-beam spectral interferometry,” Appl. Opt. 46(35), 8499–8505 (2007). [CrossRef] [PubMed]

21.

K. Takada, H. Yamada, and K. Okamoto, “320-channel multiplexer consisting of a 100 GHz-spaced parent AWG and 10 GHz-spaced subsidiary AWGs,” Electron. Lett. 35(10), 824–826 (1999). [CrossRef]

22.

B. I. Akca, C. R. Doerr, G. Sengo, K. Wörhoff, M. Pollnau, and R. M. de Ridder, “Broad-spectral-range synchronized flat-top arrayed-waveguide grating applied in a 225-channel cascaded spectrometer,” Opt. Express 20(16), 18313–18318 (2012). [CrossRef] [PubMed]

23.

B. Hofer, B. Povazay, B. Hermann, A. Unterhuber, G. Matz, and W. Drexler, “Dispersion encoded full range frequency domain optical coherence tomography,” Opt. Express 17(1), 7–24 (2009). [CrossRef] [PubMed]

24.

E. Fuchs and S. Raghavan, “Getting under the skin of epidermal morphogenesis,” Nat. Rev. Genet. 3(3), 199–209 (2002). [CrossRef] [PubMed]

25.

A. Alex, B. Povazay, B. Hofer, S. Popov, C. Glittenberg, S. Binder, and W. Drexler, “Multispectral in vivo three-dimensional optical coherence tomography of human skin,” J. Biomed. Opt. 15(2), 026025 (2010). [CrossRef] [PubMed]

26.

J. A. Izatt and M. A. Choma, “Theory of optical coherence tomography,” in Optical Coherence Tomography: Technology and Applications, W. Drexler and J. G. Fujimoto, eds. (Springer, Berlin, New York, 2008), pp. 47–72.

27.

M. M. Spühler, B. J. Offrein, G. Bona, R. Germann, I. Massarek, and D. Erni, “A very short planar silica spot-size converter using a nonperiodic segmented waveguide,” J. Lightwave Technol. 16(9), 1680–1685 (1998). [CrossRef]

28.

O. Mitomi, K. Kasaya, and H. Miyazawa, “Design of a single-mode tapered waveguide for low-loss chip-to-fiber coupling,” IEEE J. Quantum Electron. 30(8), 1787–1793 (1994). [CrossRef]

29.

A. Sugita, A. Kaneko, K. Okamoto, M. Itoh, A. Himeno, and Y. Ohmori, “Very low insertion loss arrayed-waveguide grating with vertically tapered waveguides,” IEEE Photon. Technol. Lett. 12(9), 1180–1182 (2000). [CrossRef]

OCIS Codes
(170.4500) Medical optics and biotechnology : Optical coherence tomography
(230.3120) Optical devices : Integrated optics devices

ToC Category:
Medical Optics and Biotechnology

History
Original Manuscript: May 3, 2013
Revised Manuscript: June 28, 2013
Manuscript Accepted: June 29, 2013
Published: July 3, 2013

Virtual Issues
Vol. 8, Iss. 8 Virtual Journal for Biomedical Optics

Citation
B. I. Akca, B. Považay, A. Alex, K. Wörhoff, R. M. de Ridder, W. Drexler, and M. Pollnau, "Miniature spectrometer and beam splitter for an optical coherence tomography on a silicon chip," Opt. Express 21, 16648-16656 (2013)
http://www.opticsinfobase.org/oe/abstract.cfm?URI=oe-21-14-16648


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References

  1. D.  Huang, E. A.  Swanson, C. P.  Lin, J. S.  Schuman, W. G.  Stinson, W.  Chang, M. R.  Hee, T.  Flotte, K.  Gregory, C. A.  Puliafito, J. G.  Fujimoto, “Optical coherence tomography,” Science 254(5035), 1178–1181 (1991). [CrossRef] [PubMed]
  2. A. F.  Fercher, “Optical coherence tomography - development, principles, applications,” Z. Med. Phys. 20(4), 251–276 (2010). [CrossRef] [PubMed]
  3. D.  Culemann, A.  Knuettel, E.  Voges, “Integrated optical sensor in glass for optical coherence tomography,” IEEE J. Sel. Top. Quantum Electron. 6(5), 730–734 (2000). [CrossRef]
  4. E.  Margallo-Balbás, M.  Geljon, G.  Pandraud, P. J.  French, “Miniature 10 kHz thermo-optic delay line in silicon,” Opt. Lett. 35(23), 4027–4029 (2010). [CrossRef] [PubMed]
  5. G.  Yurtsever, P.  Dumon, W.  Bogaerts, R.  Baets, “Integrated photonic circuit in silicon on insulator for Fourier domain optical coherence tomography,” Proc. SPIE 7554, 75541B (2010). [CrossRef]
  6. M. K.  Smit, “New focusing and dispersive planar component based on an optical phased array,” Electron. Lett. 24(7), 385–386 (1988). [CrossRef]
  7. D.  Choi, H.  Hiro-Oka, H.  Furukawa, R.  Yoshimura, M.  Nakanishi, K.  Shimizu, K.  Ohbayashi, “Fourier domain optical coherence tomography using optical demultiplexers imaging at 60,000,000 lines/s,” Opt. Lett. 33(12), 1318–1320 (2008). [CrossRef] [PubMed]
  8. D. H.  Choi, H.  Hiro-Oka, K.  Shimizu, K.  Ohbayashi, “Spectral domain optical coherence tomography of multi-MHz A-scan rates at 1310 nm range and real-time 4D-display up to 41 volumes/second,” Biomed. Opt. Express 3(12), 3067–3086 (2012). [CrossRef] [PubMed]
  9. Y.  Jiao, B. W.  Tilma, J.  Kotani, R.  Nötzel, M. K.  Smit, S.  He, E. A.  Bente, “InAs/InP(100) quantum dot waveguide photodetectors for swept-source optical coherence tomography around 1.7 µm,” Opt. Express 20(4), 3675–3692 (2012). [CrossRef] [PubMed]
  10. B. W.  Tilma, Y.  Jiao, J.  Kotani, E.  Smalbrugge, H. P. M. M.  Ambrosius, P. J.  Thijs, X. J. M.  Leijtens, R.  Nötzel, M. K.  Smit, E. A. J. M.  Bente, “Integrated tunable quantum-dot laser for optical coherence tomography in the 1.7 µm wavelength region,” IEEE J. Quantum Electron. 48(2), 87–98 (2012). [CrossRef]
  11. B. I.  Akca, V. D.  Nguyen, J.  Kalkman, N.  Ismail, G.  Sengo, F.  Sun, T. G.  van Leeuwen, A.  Driessen, M.  Pollnau, K.  Wörhoff, R. M.  de Ridder, “Toward spectral-domain optical coherence tomography on a chip,” IEEE J. Sel. Top. Quantum Electron. 18(3), 1223–1233 (2012). [CrossRef]
  12. V. D.  Nguyen, B. I.  Akca, K.  Wörhoff, R. M.  de Ridder, M.  Pollnau, T. G.  van Leeuwen, J.  Kalkman, “Spectral domain optical coherence tomography imaging with an integrated optics spectrometer,” Opt. Lett. 36(7), 1293–1295 (2011). [CrossRef] [PubMed]
  13. B. I.  Akca, L.  Chang, G.  Sengo, K.  Wörhoff, R. M.  de Ridder, M.  Pollnau, “Polarization-independent enhanced-resolution arrayed-waveguide grating used in spectral-domain optical low-coherence reflectometry,” IEEE Photon. Technol. Lett. 24, 848–850 (2012).
  14. V. D.  Nguyen, N.  Weiss, W.  Beeker, M.  Hoekman, A.  Leinse, R. G.  Heideman, T. G.  van Leeuwen, J.  Kalkman, “Integrated-optics-based swept-source optical coherence tomography,” Opt. Lett. 37(23), 4820–4822 (2012). [CrossRef] [PubMed]
  15. K.  Wörhoff, E. J.  Klein, M. G.  Hussein, A.  Driessen, “Silicon oxynitride based photonics,” in Proceedings of IEEE International Conference on Transparent Optical Networks (IEEE, 2008), pp. 266–269.
  16. W. H.  Louisell, “Analysis of the single tapered mode coupler,” Bell Syst. Tech. J. 33, 853–870 (1955).
  17. M. K.  Smit, C.  van Dam, “PHASAR-based WDM-devices: Principles, design and applications,” IEEE J. Sel. Top. Quantum Electron. 2(2), 236–250 (1996). [CrossRef]
  18. C. M.  Sparrow, “On spectroscopic resolving power,” Astrophys. J. 44, 76–86 (1916). [CrossRef]
  19. H.  Takahashi, S.  Suzuki, K.  Kato, I.  Nishi, “Arrayed-waveguide grating for wavelength division multi/demultiplexer with nanometer resolution,” Electron. Lett. 26(2), 87–88 (1990). [CrossRef]
  20. Z.  Hu, Y.  Pan, A. M.  Rollins, “Analytical model of spectrometer-based two-beam spectral interferometry,” Appl. Opt. 46(35), 8499–8505 (2007). [CrossRef] [PubMed]
  21. K.  Takada, H.  Yamada, K.  Okamoto, “320-channel multiplexer consisting of a 100 GHz-spaced parent AWG and 10 GHz-spaced subsidiary AWGs,” Electron. Lett. 35(10), 824–826 (1999). [CrossRef]
  22. B. I.  Akca, C. R.  Doerr, G.  Sengo, K.  Wörhoff, M.  Pollnau, R. M.  de Ridder, “Broad-spectral-range synchronized flat-top arrayed-waveguide grating applied in a 225-channel cascaded spectrometer,” Opt. Express 20(16), 18313–18318 (2012). [CrossRef] [PubMed]
  23. B.  Hofer, B.  Povazay, B.  Hermann, A.  Unterhuber, G.  Matz, W.  Drexler, “Dispersion encoded full range frequency domain optical coherence tomography,” Opt. Express 17(1), 7–24 (2009). [CrossRef] [PubMed]
  24. E.  Fuchs, S.  Raghavan, “Getting under the skin of epidermal morphogenesis,” Nat. Rev. Genet. 3(3), 199–209 (2002). [CrossRef] [PubMed]
  25. A.  Alex, B.  Povazay, B.  Hofer, S.  Popov, C.  Glittenberg, S.  Binder, W.  Drexler, “Multispectral in vivo three-dimensional optical coherence tomography of human skin,” J. Biomed. Opt. 15(2), 026025 (2010). [CrossRef] [PubMed]
  26. J. A. Izatt and M. A. Choma, “Theory of optical coherence tomography,” in Optical Coherence Tomography: Technology and Applications, W. Drexler and J. G. Fujimoto, eds. (Springer, Berlin, New York, 2008), pp. 47–72.
  27. M. M.  Spühler, B. J.  Offrein, G.  Bona, R.  Germann, I.  Massarek, D.  Erni, “A very short planar silica spot-size converter using a nonperiodic segmented waveguide,” J. Lightwave Technol. 16(9), 1680–1685 (1998). [CrossRef]
  28. O.  Mitomi, K.  Kasaya, H.  Miyazawa, “Design of a single-mode tapered waveguide for low-loss chip-to-fiber coupling,” IEEE J. Quantum Electron. 30(8), 1787–1793 (1994). [CrossRef]
  29. A.  Sugita, A.  Kaneko, K.  Okamoto, M.  Itoh, A.  Himeno, Y.  Ohmori, “Very low insertion loss arrayed-waveguide grating with vertically tapered waveguides,” IEEE Photon. Technol. Lett. 12(9), 1180–1182 (2000). [CrossRef]

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