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We demonstrate mid-infrared spectroscopic imaging using a unique optical fiber probe consisting of an array of Ge waveguide cores embedded in a silica fiber matrix. Biological tissue slices are characterized to illustrate its potential endoscopic uses. The fiber probe based transmission measurements show excellent agreement with the result obtained from standard Fourier Transform Infrared spectroscopy transmission measurements in the wavelength range of 3289.8 nm to 3383.3 nm, where fat and muscle tissues could be spectroscopically distinguished.
© 2014 Optical Society of America
1. Introduction
Infrared spectroscopic imaging is a powerful technique for characterizing the distribution of chemical species within biological structures [1–5]. Infrared light is absorbed less than visible light in most biological tissues, allowing for characterization to greater depths [6–8]. Molecular vibrations associated with organic functional groups are prominent in the mid-infrared region of the spectrum, providing a distinct chemical fingerprint [9–11]. With the goal of chemically characterizing specimens with restricted optical access, development of spectroscopic infrared endoscopes has been pursued in the past two decades [12]. Near-infrared fluorescence imaging [13,14] and near-infrared Raman spectroscopic analyses [15,16] of mice and human soft tissues have been demonstrated using infrared fiber-optic bundles with core diameter ranging from ~300 μm to 500 μm.
Mid-infrared endoscopes have also been developed to deliver laser light (1.8 – 2 μm) [17] and transfer infrared images (2 – 14 μm) [18] to assist in surgery. Advances in Fourier Transform Infrared (FT-IR) fiber-optic techniques have enabled mid-infrared spectroscopy studies on tissues [19–21]. In these studies, the infrared irradiation from a FT-IR spectrometer is coupled into fiber probes, typically silver halide or chalcogenide. Although their optical losses on the order ~5 dB/m in the 3 - 10 μm region of the spectrum [22,23] are low enough, their large pixel size ranging from ~100 μm to 700 μm severely limits the attainable optical resolution.
In this letter, we demonstrate the proof-of-concept of mid-infrared spectroscopic imaging enabled by a unique high resolution imaging fiber probe composed of an array of semiconductor waveguides in a silica fiber matrix. Biological samples with mid-infrared sensitive chemical signatures have been characterized in-vitro to demonstrate the potential of this array for high resolution endoscopy.
2. Fabrication and imaging setup
Previously, we proposed and demonstrated near-field infrared image transfer through a fiber probe with an array of waveguides [24]. The probe used in that experiment was fabricated by infiltrating mid-infrared transparent semiconductor germanium (Ge) into the hollow cores of a silica microstructured optical fiber through a high pressure chemical fluid deposition technique [25]. The materials processing parameters are published elsewhere [26].
In our design, infrared light is highly confined inside the Ge due to its much higher refractive index (n ~4), as compared with silica (n ~1.4) in wavelength of 3 - 4 μm. Therefore, each Ge waveguide acts as a single pixel and the absorbing silica matrix provides the isolation between these pixels. Figure 1(a) is a cross-sectional optical micrograph of the as-deposited fiber probe after optical grade polishing. The diameter of each pixel is 13 μm. We define one of the two probe ends as the input side and the other end as the output side. Both sides are polished after deposition using colloidal silica to minimize coupling loss of light into and out of the fiber. A layer of 100 nm gold + 10 nm chromium (adhesion layer) is then deposited on one side of the fiber using an electron beam evaporator (Kurt Lesker Lab-18). Finally, each waveguide is exposed by milling the covering metal film into a circular aperture with ~7 μm in diameter using Focused Ion Beam (FIB) technique. Figure 1(b) shows the optical micrograph of the fabricated metal mask on the end of the fiber. The metal films function as a reflective layer for mid-infrared light since silica is still transparent (although transmission is low) at wavelength range of 3289.8 nm – 3383.3 nm, within which the spectra are collected. Experimentally we have observed that when the metal film is deposited on the output side, the pixels in the final images are better defined. This is due to the resolution limitation of our collective objective so that closely-packed 13 μm diameter pixels are not well resolved. Therefore, we set the film deposited end as the output side so that the pixels are better resolved. The 168-waveguide array fabricated this way resembles the multichannel detector array in infrared imaging, but with much higher resolution which is evaluated by both the size of the apertures (7 μm) and their center-center spacing of 16 μm.
Fig. 1 Cross-sectional optical micrographs of (a) the input side of the fiber probe after polishing and (b) output side of the fiber probe with deposited Au/Cr film after FIB milling to expose the Ge cores.
A continuous wave (CW) periodically-poled lithium niobate (PPLN) crystal based optical parametric oscillator (OPO) is used as the mid-infrared source. The PPLN chip being used has four channels with poling periodicity of 29.75 μm, 30.00 μm, 30.25 μm and 30.50 μm respectively. The OPO is pumped via a Nd:YLF CW laser (Quantronix 4216D) operating at 1053 nm. Output wavelength has been calibrated and it can be tuned through changing the temperature of the PPLN crystal, which is placed in a homemade furnace connected with a temperature controller (LakeShore 331). The tunable range is from 3.21 μm to 3.47 μm with a step of 0.20 nm (corresponds to 0.1 °C temperature step) in precision. The wavelength tuning speed is about 50 seconds (time required for temperature to stabilize) per 2 nm tuning step. The OPO can output 50 mW-300 mW power of light depending on the pump laser power.
The infrared spectroscopic imaging capability of the fiber probe is demonstrated using the experimental setup schematically illustrated in Fig. 2. The OPO output power is attenuated to ~4 mW using a neutral density filter with OD ~1.1 and then the light is focused by a 3-5 μm antireflection coated, 40 mm focal length ZnSe lens and uniformly illuminated on the biological tissues. The infrared transmission image of the tissues is then collected and transferred through the fiber probe, which is mounted on an XYZ translation stage. Transmitted signals are then collected by a 0.25 numerical aperture ZnSe objective and focused onto a 3-5 μm InSb CCD camera (FLIR SC7000). The sensitivity of the camera evaluated by the Noise Equivalent Temperature Difference (NETD) factor is about 25 mK.
Fig. 2 Schematic of the imaging system including the OPO light source and the fiber probe imaging setup.
3. Tissue preparation and characterization
In order to demonstrate that the probe has the capability of spectroscopically identifying different tissue components, 40 μm thick beef muscle and fat slices are sectioned by a high precision clinical cryostat (Leica CM1950) as biological tissues. The tissues are first placed on a piece of microscope slide and dried for several days to eliminate water content, since water molecules strongly absorb mid-infrared light. Although the coherent laser light has higher optical power than the thermal light source in conventional FT-IR technique and the sensitivity of our camera is high enough for measuring transmission of tissue samples containing water, as the measurement going (using both FT-IR and fiber based technique), the water may evaporate and cause error in the measured spectra. Therefore, the tissues prepared in this work are dried before any measurements.
FTIR-spectroscopy (Bruker Hyperion 3000 FT-IR Microscope) transmission measurement is then performed through the dried tissues from 2880 cm−1 to 3115 cm−1 (3210 nm to 3472 nm) on the red and blue dotted squares indicated in Figs. 3(a) and 3(b), respectively. The spectral resolution in the measurement is 1 cm−1. The transmission spectra of the tissues are normalized by the transmission of the glass slide in this wavelength range. In order to get enough signal, the aperture size of the FT-IR is chosen as 40 μm × 40 μm.
Fig. 3 Optical micrographs of the sectioned (a) beef fat and (b) muscle slices under transmission mode. The red and blue dotted squares indicate the areas where FT-IR transmission measurements are performed through.
After the measurement, the tissues are then carefully removed from the slide onto the input side of the fiber probe using a razor blade. The FT-IR measured tissue regions are placed on enough numbers of waveguide cores and several cores are intentionally left uncovered so that they can serve as the reference for normalizing the transmission through the tissue regions. Figure 4(a) shows the configuration of the tissues on the input side of the fiber probe. The dashed rhombuses correspond to the regions indicated in Figs. 3(a) and 3(b).
Fig. 4 (a) Optical micrographs of the tissues being located on the input side of the fiber probe. (b) Transmitted light image of the output side of the probe. The brighter purple spots are uncovered pixels. The red (fat) and blue (muscle) dashed rhombuses indicate the regions where transmission data is measured by the fiber probe.
At each wavelength, an infrared image is recorded on the computer as raw data. The images are then collected from wavelength of 3289.8 nm to 3383.3 nm, where our OPO has stable output and, in addition, the two tissue components have spectroscopic contrast determined from FT-IR. The spectral resolution is chosen as ~2 nm, and the exposure time for each image is 20 ms. The contrast between these two components arises from their different absorption behavior in this wavelength range. Figure 4(b) shows the infrared image collected at the output side of the probe at a wavelength of 3380.8 nm where both the fat and the muscle tissues have very high optical absorption. The brighter purple spots in the image are the light transmitted through the uncovered pixels. The transmission spectra of both muscle and fat regions are obtained through processing the raw data in MATLAB® software. In our probe design, each semiconductor waveguide individually measures the transmission spectrum of the local volume of the tissue. To be consistent with the FT-IR measurement, the eventual transmission values of both muscle and fat are obtained through averaging 4 × 4 pixels, which give the same area as the FT-IR aperture. The uncovered pixels are also averaged to provide a valid reference in order to normalize the transmission value of both the tissue components. To minimize the effect of the background noise of the CCD camera on the eventual transmission value, the background noise was fitted and subtracted from each IR image. The results have been compared with FT-IR transmission measurement on the same tissue regions.
4. Results and discussions
We find that our technique shows excellent agreement with the FT-IR measurement result. The solid red and blue curves in Fig. 5 are transmission spectra of the beef fat and the muscle components measured by the FT-IR. Transmission data measured by the fiber probe are shown by the red and blue dots. The measured regions are indicated by the red and blue dashed rhombuses in Fig. 4(a). At wavelengths below 3312 nm, both technique show that fat tissue has higher transmission than muscle, while at wavelengths above 3312 nm, the transmission of muscle tissue is higher.
Fig. 5 The comparison of beef muscle (blue) and fat tissue (red) transmissions measured by FT-IR (solid lines) and fiber probe (dots).
The spectroscopic contrast between these two tissue components is also reflected by the infrared images. Figures 6(a)-(d) are several images recorded at different wavelengths of light. The intensity variation with the wavelength of both tissue components can be easily seen. The intensity profile variation between individual pixels can be attributed to the variation of diameter of the apertures and the slight thickness variation of the tissues at different spatial positions. The in-coupling condition for each waveguide core cannot be guaranteed to be exactly the same, but should be similar because a long focal length lens is used to uniformly illuminate the light on the whole cross-section of the fiber probe. Even so, we have considered the slight difference in intensity from different apertures into our data analysis using the infrared image of the fiber itself before placing the biological tissues on it. The effect of changing the in-coupling condition for all the waveguide cores on the transmission spectrum is eliminated through the process of normalizing the spectrum using the reference cores.
Fig. 6 Infrared images collected at the output side of the fiber probe at different wavelengths showing the contrast between the two tissue components and the intensity variation with the wavelength. The red (fat) and blue (muscle) dashed rhombuses indicate the regions where transmission data is measured by the fiber probe.
The slight transmission spectra difference in Fig. 5 may be attributed to a slight scattering effect difference between our fiber-based technique and FT-IR technique. As the light transmits through the tissues, scattering effects may result in a pixel in one position capturing scattered light from another position, but same effect may happen in FT-IR technique as well. In conventional FT-IR transmission measurement geometry, a size-tunable aperture in the microscope is used to selectively collect transmission signal from the specimens. The scattering effect may result in the aperture capturing some portion of light from specimenareas that are beyond the aperture area. The waveguide cores in our fiber probe are functioning as the collective aperture in FT-IR spectrometer. But in our case, the scattering effect should play a smaller role since the specimens are placed in direct contact with the fiber, shortening the distance between specimen and the aperture.
Although the data acquisition time of the fiber technique (~40 minutes) is longer than FT-IR technique (~20 minutes, including fat, muscle and glass slide transmission measurements), the information obtained from the fiber technique can be valuable in several ways: FT-IR can only get an averaged transmission value over a certain specimen area, while the fiber technique can also get the local information, because each waveguide core is a probe and a pixel in the final image. In order to collect information from a smaller sample area using FT-IR technique, one needs to raster the sample stage and this will definitely take longer time and lower the signal power through a smaller aperture. One advantage of the fiber technique is that it can collect different sample information simultaneously from different regions with high spatial resolution and the data analysis is more versatile. Although high spatial resolution [27] and near field imaging [28] can be achieved in FT-IR technique, the primary advantage of developing fiber-based technique is that fiber-optics can lead to endoscopes for characterizing specimens with restricted optical access.
5. Future works
In this experiment, we are only using a straight fiber to demonstrate the proof-of-concept of the fiber probe. However, these pixels can be made smaller with sub-wavelength sizes by tapering the microstructured optical fiber using a fusion splicer [24] in order to further improve the spatial resolution and to provide a built-in pixel and image magnification. Optical loss is a problem that currently inhibits the maximum length of the fiber probe. The length of the probe that is used in this experiment is ~1 mm, and the loss for a single channel at the wavelength range used in this work is estimated to be 10-12 dB/cm based on our previous loss measurement on single core Ge waveguide with same diameter. The loss in these cores is dominated by scattering mechanism. Crystallizing the amorphous Ge that is used in this experiment to form a single crystal is one way to reduce the optical losses; we have recently been able to create long Si single crystals (~2 mm) by post thermal annealing and preliminary ongoing loss measurement results show low loss of ~0.47 dB/cm at 1550 nm, which is very close to the intrinsic material loss calculated for crystalline Si (~0.40 dB/cm). Intrinsic loss should be significantly lower at the mid-infrared wavelength (e.g. 10−2-10−3 dB/cm at 2.7-4 μm wavelength) due to the smaller extinction coefficient. These losses should therefore enable 1-2 meter long fibers in the future, which will be ideal for endoscopes. These investigations of the loss measurement will be reported in a future publication, but they already indicate that this reported avenue for infrared imaging spectroscopy is highly promising. Once this is done, with improved fiber quality, we will perform further in-vivo studies on more chemically defined specimens (cells, skins, molecules, self-assembled monolayers etc.) even in a reflection mode.
6. Conclusions
In conclusion, we have demonstrated the proof-of-concept of mid-infrared spectroscopic imaging using a unique semiconductor-filled optical fiber probe. Mid-infrared images and transmission spectra of biological tissues are recorded and measured from wavelength of 3289.8 nm to 3383.3 nm. The comparison between this technique and Fourier Transform Infrared spectroscopy transmission measurement shows excellent agreement. The probe has the advantage of imaging different tissue components simultaneously with high spatial resolution. By further reducing the optical losses, the probe can be made longer (~1 m length is required for endoscopy applications) and this could potentially promote the advent of next-generation infrared medical endoscope technology.
Acknowledgments
The authors acknowledge financial support from the National Science Foundation Grant No. DMR-1107894, and the Penn State Materials Research Science and Engineering Center Grant No. DMR-0820404. The authors also would like to thank Trevor Clark from PSU Materials Characterization Lab for his help with the FIB work and John Cantolina from PSU Huck Institutes of the Life Sciences for his help with the tissue slices sectioning.
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