Ultra-high resolution optical coherence tomography (OCT) imaging is demonstrated simultaneously at 840 nm and 1230 nm central wavelength using an off-the-shelf turn-key supercontinuum light source. Spectral filtering of the light source emission results in a double peak spectrum with average powers exceeding 100 mW and bandwidths exceeding 200 nm for each wavelength band. A free-space OCT setup optimized to support both wavelengths in parallel is introduced. OCT imaging of biological tissue ex vivo and in vivo is demonstrated with axial resolutions measured to be <2 µm and <4 µm at 840 nm and 1230 nm, respectively. This measuring scheme is used to extract spectroscopic features with outstanding spatial resolution enabling enhanced image contrast.
©2007 Optical Society of America
Optical coherence tomography (OCT) is a non-invasive cross-sectional imaging modality with high spatial resolution power . Since the first application of OCT imaging for in vivo ophthalmologic examination of pathologic changes of the retina  and for non-invasive access to the dimensions of the anterior chamber of the eye  it has been established in a multiplicity of biomedical applications  and for material characterization .
To use OCT as an instrument for “virtual histology”, micrometer-scale resolution is a fundamental prerequisite. In OCT the axial resolution is determined by the coherence length of the light source and consequently is inversely proportional to its bandwidth. Most current OCT systems employ compact, low cost and easy to use superluminescent diodes (SLDs) at 800 nm or 1300 nm central wavelength which enable imaging resolutions in the range of 10 µm. Kerrlens mode locked solid state lasers allow for axial resolutions below 5 µm and opened the field of ultra-high resolution optical coherence tomography (UHR-OCT) [6, 7]. Additional extracavity spectral broadening in highly nonlinear fibers allowed for sub-micrometer resolution OCT at 800 nm and sub-2-micrometer resolution at 1.13 and 1.38 µm wavelength [8–10] thereby pushing the axial image resolution close to the physical limit .
For this reason, techniques exceeding image resolution to further improve the significance of OCT attracted the interest of the scientific community over the last years. Techniques such as Doppler-OCT, polarization-sensitive OCT, and second harmonic OCT are applied to additionally detect blood flow, birefringence and optical nonlinearities, respectively [12–14]. A number of techniques referred to as spectroscopic OCT (SOCT) were introduced to exploit the spectroscopic response of the sample for contrast enhancement. The evaluation of spectroscopic content of the backscattered light in combination with UHR-OCT was shown to distinctly enhance contrast and provide additional information on the composition and function of normal or pathological tissue [15–17]. The spectral distribution of the backscattered light can be extracted from interference patterns of OCT signals. The spectral center of gravity  or the spectral autocorrelation  is used as a measure for these SOCT methods. Imaging the same cross section by use of different wavelength bands is a related approach to extract spectral information. Sequential imaging at different wavelengths was applied to characterize light scattering coefficients of biological tissue at 830 nm and 1300 nm  and to localize a near infrared dye within tissue . Simultaneous OCT imaging at two operating wavelengths as a prerequisite for in vivo imaging was demonstrated using a single interferometer setup [20, 21] and two interferometers in parallel . Such imaging systems were applied for differential absorption OCT [23, 24]. Dual-wavelength OCT systems used so far for these investigations apply a pair of SLD light sources with coherence lengths exceeding 10 µm.
In this paper, we report on simultaneous dual-band ultra-high resolution OCT imaging. The filtered spectrum from an off-the-shelf all fiber integrated supercontinuum (SC) source achieves >200 nm bandwidth at two wavelength bands centered at 830 nm and 1230 nm, hence enabling OCT imaging with 1.7 µm and 3.8 µm axial resolutions, respectively. A free-space OCT system based on a Mach-Zehnder interferometer is designed to support both wavelength bands in parallel with sensitivities of <-95 dB. To demonstrate the capability of the introduced enhancement of UHR-OCT we recorded images of biological tissue in vivo and in vitro. In particular, the illustration of the inner structure of tissue is improved by the combination of the high resolution power at 840 nm and the high penetration depth at 1230 nm as well as by speckle reduction via frequency compounding. Evident contrast enhancement is demonstrated based on the specific wavelength dependency of tissue scattering coefficients.
2. Experimental setup and system performance
The supercontinuum light source SC500 (Fianium Ltd., Southampton, UK) we used for dual-band OCT compromises a passively mode-locked Yb-doped fiber laser, a high power cladding pumped fiber amplifier and a highly nonlinear photonic crystal fiber (PCF). The broadband output of the source is delivered in a collimated beam with an output power of up to 2.0 W, depending on the adjustable pump power of the fiber amplifier. The SC spectrum of the light source measured at 1.8 W output power is represented in Fig. 1 (black curve) on a linear and logarithmic scale. Here, the intensity peak at 1060 nm represents the residual pump laser emission which is not converted into SC radiation within the PCF.
A Gaussian spectral shape resulting in a Gaussian coherence function would be ideal for OCT imaging. Spectral modulations cause side lobes in the corresponding coherence functions thereby reducing sensitivity and resolution. To optimize the spectral shape of the light source we used wavelength dependent attenuation within a double-pass prism sequence with adjustable razor blades and two end mirrors as a very flexible approach (Fig. 2(a)). Here, the cut-off wavelengths of the spectral bands around 840 nm and 1230 nm are adjusted by the positions of the razor blades and the first end mirror. The residual pump peak is blocked by a short pass filter inserted in front of the second end mirror. In Fig. 1 the filtered continuum spectrum which was used for dual-band OCT imaging (gray curve) is compared to the unfiltered spectrum (black curve) on a linear and logarithmic scale. The full-width at half-maximum bandwidths of the spectral bands measured after passing the prism sequence were 223 and 211 nm centered at 840 and 1230 nm, respectively. The measured power is 107 mW in the 830 nm band and 119 mW in the 1230 nm band, sufficiently high for OCT imaging. Hence, no effort was made to reduce the overall power loss of 85% within the prism sequence.
A Keplerian telescope beam expander using NIR achromatic lenses in combination with a 50 µm diameter pinhole was applied to expand the beam and for spatial filtering. The output beam was coupled into the time-domain OCT system depicted in Fig. 2(b) for simultaneous dual-band UHR-OCT using this light source. The optical components of the Mach-Zehnder interferometer setup were chosen to support the full spectral bandwidth of the filtered SC source ranging from 650 nm to 1450 nm; in-house fabricated metallic beamsplitters (BSs) were used. The reflectivities of BS1 and BS2 were chosen to be 98% and 95%, respectively. The high reflectivity of BS2 reduces the incident power on the sample to an acceptable exposure level for in vivo imaging, thereby achieving efficient utilization of the light backscattered within the sample. The rate of 95% is comparable to that achieved with optical circulators, which are not available with the appropriate bandwidth. The splitting ratio of BS4 used to recombine the sample and reference arm light has to be symmetric to efficiently cancel random intensity noise of the light source by balanced detection. It was optimized to 0.50±0.03 within the full spectral range of the light source. The absorption loss of this beamsplitter was measured to be 30%, thereby reducing the fraction of the backscattered sample arm light which is utilized for signal detection in our setup to about 65%.
An NIR achromatic lens with 35 mm focal length was used to focus the light in the sample, resulting in a lateral resolution of approximately 7 µm at 840 nm and 11 µm at 1230 nm. To balance additional chromatic dispersion of the thicker lens in the reference arm LaKN22 and SF6 glass plates of appropriate thicknesses were introduced into the sample arm. Dichroic mirrors with a reflection band from 1050 nm to 1350 nm were used to separate the two spectral bands before detection. The interference pattern of the light returning from the sample and reference arm are separately detected for both wavelength bands by balanced photoreceiver heads (Femto Messtechnik GmbH, Berlin, Germany) based on silicon- (840 nm band) and InGaAs-photodiodes (1240 nm band). The detected signals are amplified and bandpass filtered before being sampled by a multi-channel transient recorder (Saturn Transient recorder, AMO Gmbh, Aachen, Germany). Optical path length scanning of the reference beam is realized by a multipass translating retroreflector based on a piezoelectric stage with a scan depth of 2 mm, a repetition rate of 50 Hz and a scanning speed of 235 mm/s. This results in signal carrier frequencies of 560 kHz at 840 nm and 367 kHz at 1230 nm.
The measured spectra of the filtered bands used for OCT imaging are presented in Figs. 3(a) and 3(b) (gray curves). The axial point spread functions (PSFs) for the two wavelength bands were characterized by measuring a single reflection on a glass substrate using a neutral density filter within the sample arm and corresponding dispersion matching in the reference arm. The PSFs are measured simultaneously and displayed on a linear scale in Figs. 3(c) and 3(d). The resulting axial resolutions in air are 1.7 and 3.8 µm at 840 and 1230 nm central wavelength, respectively; corresponding resolutions within tissue are 1.3 and 2.8 µm. The inverse Fourier transforms of the PSFs are given as black curves in Figs. 3(a) and 3(b) confirming that almost the full bandwidth of the light source is transferred through the interferometer setup. Demodulated PSFs on a logarithmic scale are presented in Figs. 3(e) and 3(f), which were scaled to reflect the applied sample arm attenuation of -74 dB. Sensitivities of the OCT setup were determined as the minimum detectable reflection level below the sample arm power to -98 dB at 840 nm for 3 mW sample arm power and -96 dB at 1230 nm for 4 mW sample arm power.
3. Dual-band UHR-OCT imaging
Dual-band UHR-OCT imaging of a human nail fold in vivo is demonstrated in Fig. 4. Image dimensions are 2.5mm×1.5 mm. Index matching using glycerine and a glass slide was applied to reduce the entrance signal. Images obtained at 840 nm and 1230 nm are displayed in Fig. 4(a) and (b). Within the tomograms the higher scattering stratum corneum (S) can be distinguished from the epidermis (E) and the dermal-epidermal junction is clearly delineated by the low scattering band of the basement membrane (BM). The different layers of the nail plate (N) and the blood vessels (BV) within the dermis are illustrated with high contrast.
As a second example UHR-OCT imaging of the limbus cornea of a rabbit eye ex vivo is demonstrated in Fig. 5 with image dimensions of 3.5 mm×1.5 mm. Figures 5(a) and 5(b) represent the tomograms obtained at 840 nm and 1230 nm, respectively. On the left hand side of the tomograms the typical appearance of the cornea is imaged including the epithelium (Ep), the stroma (St), Descement’s membrane (DM) and endothelium (En). The transition zone between stroma (St) and sclera (Sc) is given with high contrast in the center of the image.
When comparing fine structures in the single wavelength images like the basement membrane or blood vessels within the dermis and the conjunctiva it is obvious that the resolution achieved at 840 nm represented in Figs. 4(a) and 5(a) is higher than that at 1230 nm given in Figs. 4(b) and 5(b), as expected from the PSFs. On the other hand, the higher penetration depths achieved at 1230 nm is also evident from these images.
OCT like any coherent imaging technique suffers from speckle noise. Because of the uncorrelated nature of the speckle pattern generated by separated wavelength bands speckle reduction can be achieved by frequency compounding to effectively increase image quality . We use pixelwise averaging of the OCT intensity recorded at 840 nm and 1230 nm to achieve frequency compounded images. This method is applied in Fig. 4(c) and Fig. 5(c). Here, the tissue layers, e.g. epidermis, dermis (Fig. 4(c)) and conjunctiva (Fig. 5(c)), appear with enhanced homogeneity. This allows for a better discrimination of structural information from the speckle pattern as observed for example for the blood vessels within the conjunctiva (Fig. 5(c)). Noteworthy, this step also combines the superior resolution of the image measured at 840 nm with the higher penetration depth that is achieved at 1230 nm.
Image contrast in OCT is typically restricted to the variation of backscattered intensity of different types of tissue. Numerous studies of wavelength dependent scattering and absorption in tissue document that discrimination of different tissue types based on spectroscopic properties is able to enhance image contrast in OCT [18, 22, 26–28]. The difference in backscattered intensity observed at the two wavelengths applied in dual-band OCT is a spectroscopic measure, that can be utilized to characterize the sample beyond backscattered intensity of a single wavelength band. For example, when comparing the single wavelength images in Fig. 5, scattering within the conjunctiva which covers the sclera is higher at 840 nm (a) while scattering within the sclera is higher at 1230 nm (b). This allows for clear differentiation of both tissues. We use a hue-saturation-value mapping - a color space commonly used in computer graphics applications - to represent the spectroscopic information content of the dual-band OCT images. Constructed images are presented in Figs. 4(d) and 5(d), where the value-attribute was set as the average OCT signal amplitude of both images. The saturation-attribute is fixed to a level of 0.4 to ensure equal color saturation for all colors set by the hue-attribute. To identify the hue-value the datasets of both wavelength bands were rendered with normalized intensities, averaged over the full image and additionally filtered by a two dimensional adaptiveWiener-filter, as commonly used for image processing tasks. These extra images are smoothed by a 7x7 filter and than the pixelwise difference is mapped to the hue-attribute. The color scale was adjusted to map structures with higher scattering in the 1230 nm domain in a reddish tone, while structures which are dominant in the 840 nm image are colored in blue.
Recently, SC light sources characterized by two separated spectral bands were introduced [29, 30]. It was demonstrated that both wavelength bands can be used separately for UHR-OCT achieving axial resolutions of approximately 3 µm and 6 µm at 800 nm and 1300 nm, respectively. Simultaneous detection of the wavelength bands - essential to apply dual-band OCT for in vivo imaging - was demonstrated to our knowledge only for SLD based light sources [20–22]. Here, the spectroscopic information which can be obtained by the use of a second wavelength was used to address specific research questions like the hydration within the cornea and observation of dye diffusion [23, 31].
To our opinion, the use of such imaging systems in terms of “virtual histology” is restricted due to the limited resolution of these systems. To non-invasively differentiate healthy tissue from invasive tumors and non-invasive precursors, e.g. the visualization of the integrity or nonintegrity of the basement membrane is of crucial importance. Therefore, the long-term objective of non-invasive tumor staging can only be reached using state of the art UHR-OCT systems in combination with other contrast enhancing techniques like SOCT.
When the spectral distribution for SOCT is extracted from the interference pattern of single band OCT signals by time-frequency decomposition any non-uniformity in the reference path motion will result in a Doppler shift of the spectra. When using the center of gravity of the spectrum as a measure in SOCT one has to correct for this artifact by the use of a reference interferometer at the cost of increased system complexity. The spectroscopic measure we used in this study is the difference in spatially resolved scattering intensity of the two wavelength bands. Like the bandwidth of the spectral autocorrelation , this measure is insensitive to non-uniform reference path scanning.
As only backscattered intensity is used for imaging by the method introduced here, only demodulated OCT data have to be recorded which essentially reduces data rates. Additionally, when compared to spectroscopic OCT using a single wavelength band, the complexity of data processing is significantly reduced. To visualize the spectroscopic OCT data, only basic graphic data processing is applied which should allow for easy real time implementation of image processing. The increased system complexity in consequence of the need for a second wavelength band is diminished by the use of the turn-key SC light source introduced here for UHR-OCT imaging.
Due to collinear propagation of both wavelength bands within the interferometer setup introduced here, accurate spatial superposition of the two OCT images recorded simultaneously is obtained without the need of corrections. This is clearly demonstrated in the frequency com- pounded images (Figs. 4(c) and 5(c)). Here, also no interfering effect of dispersion within the sample on the spatial superposition at deeper layers within tissue is observed. The mismatch in the coherence length of the two wavelength bands degrades to some extend the resolution of the compounded image when compared to the 840 nm image. For example the apparent layer thickness of the basement membrane imaged in Fig. 4(c) is reduced when compared to Fig. 4(a). However it is noteworthy, that the integrity of the basement membrane is comprehensible also in the compounded image.
The interpretation of the spectral information obtained by OCT imaging is challenging as the spectral composition of the backscattered light is dependent on the properties of all traversed tissue . For example, the overall scattering cross section in tissue is higher for shorter wavelengths resulting in an increasing redshift with tissue depth as observed for example in the dermal layer of the human nail fold in Fig. 4(d). It is necessary to consider the influence of scattering and absorption of the intervening layers on the measured differential spectrum to avoid misinterpretation of the spectroscopic information. Keeping this in mind the spectroscopic information obtained by dual-band OCT is capable to significantly enhance image contrast when compared to OCT based only on backscattered intensity. This is clearly demonstrated e.g. by the blueshift observed for the internal structure of the nail plate given in Fig. 4(d) and by the distinct discrimination of the conjunctiva from the stroma in the differential image of the rabbit eye given in Fig. 5(d).
Effective speckle reduction is observed for the wavelength compounded images because of the uncorrelated nature of speckle patterns generated by spectrally separated bands. The speckle reduction is associated with an increase in image contrast allowing better discrimination of structural features from the speckle noise as can be seen for example for the inner structure of the nail plate (Fig. 4(c)) and in the representation of the conjunctiva in the center of Fig. 5(c). The actuality that the speckle patterns are uncorrelated emphasizes the need for spatially filtering the differential data of the two images before mapping them to the hue value of the differential color images. Otherwise the speckle pattern of the single wavelengths will be transfered to the color distribution within the image. The representation of structural features that are given by the value-attribute are not compromised by this filtering step.
In summary, this manuscript presents dual-band UHR-OCT as a new method to achieve spatially resolved spectroscopic information in OCT imaging. Two OCT images are recorded simultaneously by use of two wavelength bands at 840 nm and 1230 nm which are supplied by a single turn-key supercontinuum light source. Here, axial resolutions of 1.3 µm and 2.8 µm are achieved within tissue. The difference in scattering intensity of the two wavelength bands is used as a measure of spectral modulation within the sample. Compared to spectroscopic OCT using a single wavelength band the requirements of data acquisition and data processing are substantially reduced. Imaging studies ex vivo and in vivo demonstrate, that this method is capable to clearly enhance the contrast between different tissue types without the need for exogenous contrast agents. A further obvious variation of the method introduced here is differential absorption OCT using functionalized near-infrared dyes as molecular contrast agents. To achieve higher imaging speeds for improved in vivo application dual-band UHR-OCT can directly be adapted to frequency-domain OCT without compromising the easy access to the spectroscopic response of the sample.
We gratefully acknowledge scientific contributions of Jaroslav Lazar and Amjad Naami. This work was financed by the Deutsche Forschungsgemeinschaft (Grant No. Ku 540/47-2).
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