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High performance, short coherence length light sources with broad bandwidths and high output powers are critical for high-speed, ultrahigh resolution OCT imaging. We demonstrate a new, high performance light source for ultrahigh resolution OCT. Bandwidths of 140 nm at 1300 nm center wavelength with high output powers of 330 mW are generated by an all-fiber Raman light source based on a continuous-wave Yb-fiber laser-pumped microstructure fiber. The light source is compact, robust, turnkey and requires no optical alignment. In vivo, ultrahigh resolution, high-speed, time domain OCT imaging with <5 µm axial resolution is demonstrated.
©2004 Optical Society of America
1. Introduction
Optical coherence tomography (OCT) is an emerging biomedical imaging technology that can perform high-speed, micron scale, noninvasive imaging of tissue morphology in vivo [1, 2]. Since OCT is based on low-coherence interferometry, the axial image resolution, Δz, is determined by the bandwidth Δλ and the center wavelength λ 0 of the light source: Δz=[2ln(2)/π](λ02/Δλ). Standard OCT systems use superluminescent diode light sources, which achieve an axial resolution of 10–15 µm. Compact and portable high performance broadband light sources with sufficient power and bandwidth are important to achieve ultrahigh resolution, high-speed OCT imaging outside the laboratory setting.
Femtosecond solid-state lasers have been demonstrated to directly generate broad bandwidths which can be used for ultrahigh resolution OCT imaging, but are difficult to operate outside the laboratory [3, 4]. Nonlinear and microstructure fibers pumped by femtosecond bulk and fiber systems have also enabled imaging with unprecedented resolutions, but require the use of femtosecond lasers [5–7]. Reductions in cost can be achieved using low-threshold femtosecond lasers which use inexpensive low-power pump lasers, but these systems are still relatively complex [8, 9]. Femtosecond fiber laser-based sources promise to be compact and robust [10]. The 1300 nm wavelength region is of particular interest for biomedical applications because it permits improved imaging depth when compared with shorter wavelengths due to reduced scattering in biological tissue [11]. The development of broadband light sources for OCT imaging in scattering tissue has therefore focused on this wavelength range [5, 7, 10, 12]. Recent work has also investigated light source development and imaging in the 1000–1100 nm wavelength range, which provides a compromise of higher resolution for a given bandwidth at the expense of reduced image penetration [7, 13]. A portable source suitable for in vivo clinical applications has been demonstrated in this wavelength range using a femtosecond diode-pumped Nd:Glass laser with a highly nonlinear fiber [13].
We report a novel approach for broadband continuum generation in microstructure fibers using a high-power, continuous-wave, all-fiber pump light source. This new light source promises to enable ultrahigh resolution, high-speed OCT imaging with lower cost and complexity than with femtosecond laser-based light sources. Microstructure fibers typically have been pumped with femtosecond lasers to provide the peak powers necessary to initiate nonlinear effects for continuum generation under conditions of anomalous or near-zero dispersion. However, the use of high intensity femtosecond pulses for broadband continuum generation under these conditions leads to severe spectral modulation of the continuum in the vicinity of the pump wavelength. This spectral modulation produces sidelobes and reduced contrast in the interferometric axial point spread function. Femtosecond pumping of microstructure fibers may also result in excessive temporal instability of the continuum and nonlinearly amplified quantum noise, which can lead to excess intensity noise [14].
An alternative to using high peak powers is to increase the effective nonlinear interaction length of the Raman interaction, which is governed by optical losses in the fiber and dispersive walk-off between the pump and continuum pulses. The use of longer pump pulses reduces this dispersive walk-off effect. Stimulated Raman scattering has been shown to be the principle nonlinearity for continuum generation by using nanosecond-scale pump pulses [15, 16]. Recently, the possibility of low peak power and even continuous-wave, multiwatt Raman continuum generation in highly nonlinear fibers has been demonstrated [17]. Continuous-wave pumping of nonlinear fibers can enable the development of robust and turnkey continuum light sources which require no optical alignment, enabling high-speed, ultrahigh resolution OCT imaging in a wide range of applications outside the laboratory.
In this paper we demonstrate a new light source for ultrahigh resolution, high-speed OCT imaging using Raman continuum generation from a continuous-wave pumped microstructure fiber. This source achieves bandwidths of ~140 nm in the 1300 nm wavelength range with output powers of 330 mW, higher than achieved using any other technique. The light source is compact (25×25×20 cm), robust, completely turnkey, and requires no optical alignment. Ultrahigh resolution, high-speed OCT imaging is demonstrated with <5 µm axial resolutions. High-speed ultrahigh resolution imaging is demonstrated in vivo in the hamster cheek pouch and in human skin.
2. Methods
The broadband all-fiber Raman continuum light source was based on a 10 W continuous-wave, non-polarized, multimode diode-pumped, single-mode Yb-fiber laser (IPG Photonics) directly spliced to an anomalously dispersive microstructure fiber (Crystal Fiber). The microstructure fiber was 100 meters long, had a dispersion of +35 ps nm-1 km-1, a pitch of Λ=1.72 µm, and an air-hole diameter of 0.65 µm. Figure 1 shows a schematic of the source and the Raman continuum output. The Raman-soliton continuum had 5.5 W of total power and a spectral width of 318 nm (at 20 dB). The spectral range from 1090 to 1370 nm was flat to ±5 dB and contained 2.3 W of power. This output was filtered using a special WDM coupler to remove the pump wavelength and to achieve a smooth, Gaussian-like spectrum in the 1300 nm wavelength range. Figure 2(a) shows a typical spectrum before and after spectral shaping by the WDM coupler. The output power after the coupler was ~330 mW and the spectrum was Gaussian-shaped with a bandwidth of ~140 nm, corresponding to a theoretical resolution of 5 µm in free space. The all-fiber light source contains no bulk optical components, requires no alignment, and is turnkey, compact and robust.
The intensity noise of the light source system was characterized using an RF spectrum analyzer. Figure 2(b) shows the noise spectrum of the Raman continuum generated by the source, the detection system noise, and the calculated shot noise level. Excess intensity noise is caused in part by feedback from the fiber-grating-based design of the Yb pump laser and reflection from the splices between the microstructure fiber and the pump laser and WDM coupler. Dual-balanced detection was used in the OCT system to reduce excess intensity noise.
Fig. 1. Left: Schematic of Raman continuum source. SMF: Single-mode fiber. Right: Output spectrum of the Raman continuum source. The output was 5.5W total, with 2.3 W in the spectral range from 1090 to 1370 nm. The output was filtered using a special WDM coupler to remove the pump wavelength and shape the spectrum.
Fig. 2. (a) Typical output spectrum of the Raman fiber light source before and after spectral shaping. (b) RF noise spectrum. High frequency oscillations are caused by spurious reflections from the fiber splices. Dual-balanced detection is used to reduce excess noise.
Figure 3 shows the schematic of the experimental setup for OCT imaging. The system consists of a dual-balanced interferometer with broadband 80/20 and 50/50 fiber couplers to optimize power coupled back to the detectors. The reference arm was scanned using a reflective delay scanner at a velocity of 6.2 m/s and 1600 Hz repetition rate. Polarization controllers were used in both the sample arm and the reference arm. Dual-balanced detection with two InGaAs photodiodes (D1 and D2) was used to reduce excess intensity noise in the light source. Since the interference signal occurs 180 degrees out of phase at the two detectors, subtracting the two signals adds the heterodyne interference signal but subtracts excess noise. The input to photodiode D2 is attenuated to match the input to photodiode D1. It was important to match the path lengths of the two arms of the double-detector receiver in order to achieve optimum noise reduction [18]. The interference signal was electronically bandpass filtered, logarithmically demodulated, low-pass filtered, and digitized. Detection was performed at a Doppler frequency of 9.8 MHz with a bandwidth of 1.2 MHz.
Fig. 3. Schematic of high-speed, ultrahigh resolution OCT system using an all-fiber Raman continuum light source. The system uses broadband 80/20 and 50/50 couplers to optimize the power on the sample and coupled back to the detectors. Dispersion was matched in the sample and reference arms of the interferometer in order to maintain high axial resolution.
In vivo imaging was performed using an XY galvanometer scanning probe. Near-infrared achromatic lenses were used to minimize chromatic aberration. The fiber collimating lens had a focal length of fc=10 mm, followed by a Hopkins relay pair of lenses with focal lengths f1=25 mm and f2=40 mm, and an objective of focal length fo=25 mm. A pair of galvanometer-controlled scanning mirrors was used between the fiber collimating lens and the Hopkins relay to perform high-speed transverse scanning. The focused spot size was 2ω0=18 µm. To minimize the effect of wavefront aberration, dehydration, and to achieve better index matching, the tissues were irrigated with saline and covered with a thin cover glass.
3. Results and discussion
The performance of the OCT system using the fiber Raman continuum light source was characterized using an isolated reflection from a single mirror. To maintain axial resolution, the dispersion in the interferometer sample and reference arms was carefully matched. Appropriate thickness glass blanks of fused silica, SFL6, and LakN22 were inserted into the reference arm to balance the dispersion of the achromatic lenses in the imaging probe. Dispersion mismatch was monitored by taking the Fourier transform of the interferometer fringe signal. Figure 4(a) and (b) show the interference signal after bandpass filtering and the logarithmic demodulated signal after lowpass filtering. The measured axial resolution was 6.3 µm in air, corresponding to 4.8 µm in tissue. The detected optical spectrum was measured to be 110 nm by Fourier transformation of the interferometric signal. This reduction in bandwidth may be the result of wavelength dependence of the fiber couplers as well as wavelength variations in the sensitivity of the InGaAs photodiodes.
The system sensitivity was measured using the minimum visible intensity of a reflection from a mirror and was measured to be 95 dB with an incident power of 20 mW. The theoretical sensitivity for this incident power and detection bandwidth is 103 dB. Parasitic losses in the sample arm optics, fiber couplers, and fiber connectors account for ~4 dB of the reduction in sensitivity. Incomplete noise cancellation from mismatches in the fiber couplers and detectors may account for the additional 4 dB reduction in sensitivity.
Fig. 4. (a) Point spread function measured using an isolated 3.0 OD attenuated reflection from a mirror. The axial resolution is 6.3 µm in air, corresponding to <5 µm in tissue. (b) Logarithmically demodulated signal showing low sidelobes.
We demonstrated ultrahigh resolution, high-speed, in vivo OCT imaging in human skin as well as in the hamster cheek pouch, a well-established model for studies of cancer progression. Imaging was performed using an XY scanning probe. Figure 5 shows high-speed, ultrahigh resolution in vivo images of human skin with 4.8 µm axial resolution. The image had 500 transverse pixels and 1000 axial pixels and covered an area of 2.25 mm by 1.8 mm. Imaging was performed at a rate of 3.2 frames per second. The axial dimension was scaled by 1.38 to account for the approximate index of refraction of skin [19]. Distinct morphological features such as keratinized stratum corneum, the junction between the dermis and epithelium, and sweat ducts are clearly visualized in the ultrahigh resolution images.
Fig. 5. In vivo high-speed OCT images of human skin. The stratum corneum layer (SC), epithelium (E), dermis, dermal-epidermal junction (inset left) and a spiraling sweat duct (inset right) can clearly be seen. Images were acquired at 3.2 frames per second. (~18 µm×4.8 µm transverse x axial resolution; 500×1000 pixels; 2.25 mm×1.8 mm).
Fig. 6. High-speed, ultrahigh resolution in vivo OCT image of Syrian hamster cheek pouch. The keratinized epithelial layer (e), muscular layers (m), and two prominent vessels (v) are clearly visible. A microscope cover glass (c) was placed over the top of the cheek pouch. (~18 µm×4.8 µm transverse x axial resolution; 500×1000 pixels; 2.4 mm×1.4 mm).
Figure 6 shows an in vivo image of hamster cheek pouch. Imaging was performed with the animal under anesthesia and in accordance with approved protocol reviewed by the M.I.T. Committee on Animal Care (CAC). The ultrahigh resolution images had 500 transverse pixels and 1000 axial pixels and covered an area of 2.4 mm by 1.4 mm. The ultrahigh resolution OCT image exhibits structure associated with the normal hamster cheek pouch. A thin, keratinized layer which is highly backscattering is present near the surface, followed by an epithelial layer (e), muscular (m) layers, and connective tissue. Two prominent blood vessels (v) are also clearly visible within the muscular layers and connective tissue. Image depths are comparable to those achieved with other 1.3 µm light sources.
To demonstrate the high-speed imaging capability of this ultrahigh resolution OCT system, in vivo three-dimensional imaging of human skin and hamster cheek pouch was performed. High-speed imaging enables volume datasets to be acquired, which permits tracking of morphological features through the imaging volume. Each three-dimensional data set consisted of sequentially acquired transverse OCT images, each containing 500 transverse pixels and 1000 axial pixels. Volume imaging of human skin consisted of 50 transverse images with 15 µm spacing between frames, spanning a transverse dimension of 0.750 mm. Imaging of hamster cheek pouch consisted of 50 transverse images with 20 µm spacing between frames, spanning a transverse dimension of 1.0 mm. The data was acquired at 3.2 frames per second, for a total acquisition time of ~15 seconds for 500×50×1000=2,500,000 data points or voxels. The three-dimensional dataset can be displayed in various orthogonal planes or rendered using image processing and rendering software.
Figure 7 shows an example of an animation showing three-dimensional imaging of the hamster cheek pouch. In one animation, Fig. 7(a), the three-dimensional data is viewed in sequential transverse slices, corresponding to the normal OCT view. The animation enables features such as the epithelial layer, muscular layers, connective tissue, as well as two large blood vessels to be tracked through successive image planes. In a second animation, Fig. 7(b), the three-dimensional data is viewed using en face slices at different depths, perpendicular to the normal OCT view, corresponding to the view typically provided by confocal microscopy. The en face slices enable larger scale features such as folds in the epithelium to be better visualized than in the transverse slices. Figure 8 shows a rendered volume of the hamster cheek pouch constructed from the three-dimensional data set. The animation shows the tissue volume as viewed from arbitrary virtual perspectives.
Fig. 7. Three-dimensional volume imaging using in vivo ultrahigh resolution OCT images acquired at 3.2 frames per second. (a). (2.3 MB) Movie showing sequential transverse slices of hamster cheek pouch with two large vessels, corresponding to the normal OCT view (6.0 MB version). (b) (2.4 MB) Movie showing en face slices at different depths, perpendicular to the OCT view, allowing features such as epithelial folds and vessels to be visualized (8.3 MB version). The rendered volumes cover 2.4 mm×1.0 mm×1.8 mm.
Fig. 8. (2.4 MB) Movie showing rendered volume of hamster cheek pouch constructed from a three-dimensional dataset. The tissue volume can be viewed from arbitrary virtual perspectives. The rendered volume covers 2.4 mm×1.0 mm by approximately 1.8 mm.
Figure 9 shows examples of three-dimensional imaging in human skin. In Fig. 9(a), the volume data is viewed in en face slices at different depths through the human fingerpad. Structures such as ridges in the stratum corneum associated with the fingerprint are clearly visible in the animation. Sweat ducts can also be clearly seen spiraling along the ridges of the fingerpad through the stratum corneum and into the epithelium. Three-dimensional rendering and segmentation can be performed as shown in Fig. 9(b), enabling the density of sweat ducts as well as individual ducts to be assessed (inset). Other structures such as the rete ridges in the junction between the epithelium and dermis can also be visualized.
Fig. 9. (a) (2.2 MB) Movie showing en face slices of human fingerpad created from cross-sectional images. Spiraling sweat ducts are seen along the fingerpad ridges (2.25×0.75×1.8 mm) (8.3 MB version). (b) (2.3 MB) Movie showing 3D rendering and segmentation of sweat ducts (6.2 MB version). [Media 7] Duct density as well as individual ducts can be assessed (1.7 MB). [Media 9]
4. Conclusion
In conclusion, we have demonstrated a new, high performance light source for high-speed, ultrahigh resolution OCT imaging. Bandwidths of 140 nm in the 1300 nm wavelength range are achieved, yielding axial resolutions of <5 µm in tissue. Output powers of 330 mW enable high-speed, OCT imaging in vivo. The axial resolution of <5 µm is comparable to that achieved using much more complex femtosecond solid state sources such as the Cr:Forsterite laser. Output powers are comparable to or better than those available from femtosecond lasers, and more than one order of magnitude higher than superluminescent diode sources.
Since this light source is all-fiber, it requires no alignment and provides completely turnkey operation. The source is extremely compact, measuring only 25×25×20 cm, and can easily be integrated into portable OCT systems. Excess noise is relatively high, but can be reduced using dual-balanced detection. Further reduction in noise should be possible by reducing parasitic feedback effects in the fiber splices and by techniques such as active seeding of the Raman process. Broader output bandwidths should be achievable using higher pump power Yb-fiber lasers and different fiber geometries. The bandwidth will ultimately be limited by the water absorption of microstructure fibers which reduces the effective interaction length of the Raman scattering. The current cost of the source is still relatively high due to the cost of the microstructure fiber. However, since demand for microstructure fiber is increasing, the price of these fibers is decreasing and can be expected to approach that of other specialty fibers. The cost of continuous-wave Yb-fiber lasers has also dramatically decreased in recent years. The cost of fiber Raman continuum light sources will therefore be significantly lower than that of bulk solid-state femtosecond laser-based light sources. The high performance and ease of use of the fiber Raman continuum source promises to enable a wide range of new ultrahigh resolution, high-speed, OCT imaging applications.
Acknowledgments
We thank V. Sharma, P. Herz, V. Srinivasan, and N. Nishizawa for assistance and helpful discussions. This research was sponsored by National Institutes of Health R01-CA75289-06 and R01-EY11289-18, National Science Foundation ECS-01-19452 and BES-0119494, Air Force Office of Scientific Research Medical Free Electron Laser Program F49620-01-1-0186 and F49620-01-01-0084, the Poduska Family Foundation Fund for Innovative Research in Cancer, and the philanthropy of Mr. G. Andlinger. Y. Chen acknowledges the Cancer Research and Prevention Foundation. T. H. Ko acknowledges the Whitaker Foundation.
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