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An all-optical 3D photoacoustic imaging probe that consists of an optical fiber probe for ultrasound detection and a bundle of hollow optical fibers for excitation of photoacoustic waves was developed. The fiber probe for ultrasound is based on a single-mode optical fiber with a thin polymer film attached to the output end surface that works as a Fabry Perot etalon. The input end of the hollow fiber bundle is aligned so that each fiber in the bundle is sequentially excited. A thin and flexible probe can be obtained because the probe system does not have a scanning mechanism at the distal end.
© 2013 Optical Society of America
1. Introduction
Photoacoustic imaging as a hybrid molecular imaging modality has been developed to visualize deep newly-formed vessels that lead to early detection of cancer of various organs. Recently, some groups have proposed thin photoacoustic endoscopes [1–4] for depth imaging of internal organs such as cardiovascular tissue, digestive tracts, and urogenital systems. In a conventional photoacoustic imaging system, piezoelectric materials such as PVDF (polyvinylidene fluoride) have been widely used for ultrasound detection. When implementing a transducer for an endoscopic imaging probe, the transducer should be very small. However, it is not easy to reduce the size while keeping the sensitivity high enough for high contrast imaging. In addition, the distal end of the sensing probe is solid in a certain length since the transducer and mechanical scanning system are enclosed.
As a promising alternative to a piezoelectric transducer, some devices that optically detect photo-induced ultrasonic waves by using Fabry-Perot etalons [5–7], fiber Bragg gratings [8], and polymer microring resonators [9, 10] have been reported. As for devices using Fabry-Perot etalons, a fiber optic probe with a water cavity sealed by a polymer-film diaphragm [11] and a single-mode fiber probe with a polymer film-based interferometer that is directly deposited on the end surface of the fiber [12] were fabricated. It was shown that the fiber probe had an acoustic sensing performance that was comparable to conventional PVDF transducers. More recently, a miniature photoacoustic fiber probe that utilizes a dual-clad optical fiber was proposed [13]. The single-mode core was used for sensing acoustic wave and the multimode core was used for delivery of excitation light and, by using a fiber with an slant end surface, a side-viewing probe was fabricated.
In this paper, we fabricate a fiber-optic Fabry-Perot acoustic sensor that uses a polymer film attached to the output end surface of an optical fiber. This sensor is appropriate for an endoscopic photoacoustic probe because it has no cavity or diaphragm that affects the stability of the detected signal. The polymer film itself detects the pressure change as a thickness variation [14]. By choosing a polymer film material that has appropriate thickness and elasticity, the sensor has high sensitivity to pressure changes.
In addition, we built a fiber-optic probe for photoacoustic 3D imaging by combining a fiber-optic acoustic probe with a bundle of hollow optical fibers for excitation of photoacoustic waves. Each element in the fiber bundle fired a laser beam in sequence, and we obtained a 3D image by reconstructing the depth images of each point. This probe can be thin and flexible without a solid part at the distal end because it does not need a mechanical scanner.
2. Fiber-optic acoustic probe
Figure 1 shows a schematic of the optical fiber probe system for ultrasound detection. This system consists of a 1550-nm distributed feedback (DFB) laser, an optical fiber circulator, and an InGaAs photo diode with a preamplifier. Usually acoustic sensing systems with a Fabry-Perot interferometer tend to be large and expensive since they need a highly stable light source and a highly sensitive detector. In our system, however, all components are those for optical communication systems that are compact and inexpensive although they have high stability and sensitivity; therefore, the system is feasible for clinical applications. The DFB laser operated in single mode with a linewidth of 0.5 nm and provided a high stability of ± 0.05 dB for 15 min. The single-mode-fiber (SMF)-based optical circulator showed an isolation of > 40 dB with an insertion loss of around 0.8 dB. The InGaAs photo diode equipped with preamplifier had a bandwidth of 30 kHz to 1 GHz and a maximum conversion gain of 700 V/W. Since all the components were connected with SMF patch cables, the system is compact and very stable in spite of the low cost.
As shown in Fig. 1, the light power Pi from the laser source was coupled into the circulator and reached the sensor probe, which was equipped with a polymer film on the end surface of the SMF. The reflected power from the probe Pr was modulated by the phase difference between P1 and P2 that are the reflections at the fiber/polymer interface and the polymer’s outer surface, respectively. Without ambient pressure, the reflected light power Pr was determined by the initial phase difference between P1 and P2, and this became the center operating point of the interferometer. When ultrasound pressure was applied to the end of the sensor probe, it induced a change in the film thickness that caused phase modulation, and then ultrasound wave was detected as an intensity change in Pr.
Next, we tried to confirm that the fiber probe system functioned as a photoacoustic detector using the experimental setup shown in Fig. 2. A Q-switched, second harmonic Nd:YAG laser with an operating wavelength of 532 nm was used as the light source to excite the photoacoustic signal. An aluminum sheet enclosed in gelatin was irradiated with laser pulses of 7 ns duration, pulse energy of 1.0 mJ, and repetition frequency of 10 Hz. To introduce the laser beam onto the Al sheet, a hollow optical fiber with an inner diameter of 0.7 mm was used.
The photoacoustic signal detected by the optical fiber probe is shown in Fig. 3. In this figure, the origin of the horizontal axis shows the time of the laser radiation. In the measured waveform, the clear signal appears at 3.8 μs, which coincides with the propagation time for the distance of 6 mm between the Al sheet and the end of the fiber probe; therefore, we confirmed that the detected signal originated from the photoacoustic wave induced by the laser beam. The Fourier spectrum of the photoacoustic signal shows that the center frequency is around 1.7 MHz and the bandwidth is nearly 1.8 MHz as shown in Fig. 4.
Generally, a photoacoustic signal induced by nanosecond pulsed laser has a wide bandwidth that ranges over several tens of MHz. However, the photoacoustic signals detected by an ultrasound probe depend on the frequency characteristics of the probe itself. In the proposed probe based on optical interference, the characteristics were determined by the thickness and elastic property of the sensing film. Therefore, we first investigated the correlation between the film thickness and the central frequency of the detected signal, and the results are shown in Fig. 5. In the experiment, the energy of the excitation laser pulses and the distance between the probe and sample were fixed, and signals were measured by the probes with various film thickness. The working point of the film-based Fabry-Perot interferometer was adjusted by trial and error. We repeatedly applied the film on the fiber and checked the response for acoustic wave to obtain a good waveform. As shown in Fig. 5, we confirmed that the center frequency became high for small film thicknesses and this was due to a shift in the resonance frequency. Figure 6 shows the correlation between center frequency and intensity of photoacoustic signals. Signal intensity decreased with increasing signal frequency because of the high attenuation in the medium. Considering the balance between signal intensity and high depth resolution obtained with high frequency, for photoacoustic imaging, we chose a film thickness of 7.5 μm, which provided a center frequency of around 5 MHz.
When using this probe for endoscopic imaging, the directivity affected the detectable area of the probe; therefore, we experimentally evaluated the directional characteristics of the probe. In the experiment, photoacoustic signals were induced at a distance of 6.5 mm from the probe with angles from −90° to + 90° in steps of 10°. Figure 7 shows the measured directional response for the RMS values of the detected signals. We confirmed that the probe could clearly detect all the induced signals and had a uniform sensitivity at 120 degrees in front of the probe. Figure 7 also shows that the probe detected signals induced from the position immediately lateral to it with around half the sensitivity of those of the center. This semi-omnidirectional property enables detection over a wide area in endoscopic imaging.
We performed experiments to compare the acoustic sensing performance of the optical fiber probe to a conventional PVDF hydrophone. In this experiment, we used a needle PVDF hydrophone with a sensing diameter of 0.5 mm and the frequency spectrum measured by the hydrophone is shown in Fig. 8. The center frequency and the bandwidth were 5 MHz and 10 MHz, respectively. The sensing film thickness of the fiber probe was adjusted to obtain the same frequency characteristic as that of the hydrophone and set each probe at the same distance from the sound source.
Figure 9 shows the photoacoustic signals detected by the optical fiber probe and the PVDF hydrophone in the same measurement system depicted in Fig. 2. The waveform acquired by the fiber optic probe agrees well with that of the hydrophone. Although usually noise equivalent pressure (NEP) values are used to evaluate characteristics of acoustic probes [12], here we used the signal-to-noise ratio (SNR) which does not require absolute pressure measurement. Since SNR is inversely proportional to NEP [15], evaluation of SNR is enough to compare performances of the acoustics probes. The SNRs of the optical fiber probe and hydrophone calculated by using the integral of the signal peak area are 14.8 dB and 14.5 dB, respectively. This result shows that the fiber optic probe enabled low-noise signal detection with a sensitivity that is comparable to that of the hydrophones.
3. Photoacoustic imaging
We first performed photoacoustic imaging using the fiber-optic acoustic probe combined with a single hollow optical fiber for excitation. We prepared two types of blood vessel phantoms that consist of silicon tubes with an inner diameter of 400 μm. One of them contained black ink, and the other contained 15% hemoglobin solution. As shown in Fig. 10, these phantoms were embedded in gelatin and it was immersed in water. The tubes were placed 5 mm beneath probe and the fluence of excitation beam was 180 mJ/cm2 hereafter. The acoustic probe bound with a hollow optical fiber scanned across the phantoms in 10 µm steps to obtain a B-mode image of the phantoms. Digital filtering and envelope demodulation were applied to the raw detected signals to construct the images, and MATLAB software was used for the calculation.
Figure 11 shows the constructed image of the blood vessel phantoms. The entire constructed area where the two phantoms could be seen at depths of 5.2 mm from the imaging probe is shown on the left. On the right there is an enlarged view of the region surrounding the phantoms, and one can see the difference in brightness between the hemoglobin phantom on the right and the black ink phantom on the left. The fiber probe detects small variations of absorption with high intensity despite existence of bubbles seen in Fig. 10 that cause scattering of the excitation beam. This result shows the capability of the combination of the fiber acoustic probe and the hollow optical fiber for excitation for photoacoustic imaging.
Next, we propose a 3D photoacoustic imaging system using a hollow fiber bundle for excitation and an acoustic fiber probe for ultrasound detection. Figure 12 shows a schematic of the proposed imaging system. An acoustic fiber probe was installed at the center of the bundled hollow fiber. The hollow fibers in the bundle were arranged in a line at the input end; therefore, one can excite the hollow fibers sequentially by linearly scanning the laser beam at the input end. Then, the 3D image was reconstructed by combining all the data obtained from each fiber whose locations have been associated with the order at the input end. The image resolution obtained was the same as the diameter of the hollow optical fiber since the output beam from the hollow optical fiber was almost parallel due to the extremely small NA (< 0.05) of the hollow fiber. This is one of the advantages of using hollow optical fiber for excitation: one does not need a lens array at the output end. However, when focusing of excitation beam becomes necessary to perform imaging of highly diffusive, biological samples, we can attach lensed caps to each fiber elements [16]. From another measurement using a single phantom tube, we evaluate the image resolution of this system. It was around 200 μm in lateral direction that is larger than the focused beam spot that was around 100 μm. This is due to the beam divergence and scattering in the sample. In depth direction, the resolution was around 250 μm at a frequency of 6 MHz.
The omnidirectional reception property shown in Fig. 7 enabled detection of the acoustic waves excited by all the fiber pixels with a single fiber probe. The distal end of the system did not have a scanning mechanism, and a very thin probing system that can be inserted into a small-diameter endoscope will be available in the future.
To show the feasibility of the system shown in Fig. 12, we performed 3D photoacoustic imaging using a bundle of 37 hollow fibers with an inner diameter of 320 μm whose distal end is shown in Fig. 13. The diameter of the output end was 3.2 mm, and the minimum bending radius of the probe was around 50 mm. At the input end of the bundle, a 532-nm Nd:YAG laser beam was sequentially launched into the fiber by using a motorized linear stage. In this experiment, an acoustic fiber probe was separately placed beside the sample. We first measured the acoustic wave from the flat sample face to correct for the effect for the variation in the fibers’ transmission losses. The intensities of the acoustic wave excited by each fiber were hexagonally arranged, and the correction coefficients were derived for each pixel to obtain a uniform intensity image.
Figure 14(a) shows the sample with silicon tubing with an inner diameter of 1 mm, and the bore was filled with 20% concentration hemoglobin solution. The imaging area is shown by the dashed circle. Figure 14(b) shows the raw image of the sample’s surface where the correction described above was applied. Then, as a smoothing process, the raw image was rearranged to 13 x 13 pixels by applying a 3 x 3 smoothing filter to obtain the image shown in Fig. 15(a). Finally, image interpolation was applied to obtain a smooth and clear image as shown in Fig. 15(b).
To reconstruct the 3D image, the same smoothing and interpolation processes were applied to seven B-mode images that correspond to the rows of the fibers at the bundle’s output end. Figure 16 shows the 3D image reconstructed by using ImageJ software. The silicon tube sample with hemoglobin solution in the bore clearly appeared in the image, and the feasibility of the proposed system with a combination of a hollow fiber bundle for excitation and a photoacoustic fiber probe is shown. The image resolution was around 250 μm in depth direction at 6 MHz that was the same as the result in the B-mode imaging shown above. The resolution in lateral direction is limited by the distance between the fiber elements. The observed resolution was around 500 μm that was larger than the outer diameter of the hollow fiber that was 430 μm. This was mainly due to beam divergence from the hollow fiber although it was small and scattering in the sample medium.
4. Conclusion
To construct an all-optical photoacoustic probe without a scanning mechanism at the distal end, we first fabricated an optical fiber probe for ultrasound detection. The probe consists of a single-mode optical fiber with a thin polymer film attached to the output end surface for detection of acoustic waves. Fiber-coupled optical components for optical communication, such as a DFB laser diode and optical circulator, were used to construct a stable and low-cost Fabry Perot interferometer. As a result of evaluating its operating characteristics, we confirmed that the probe had almost omnidirectional reception and a high SNR that is equivalent to those of common PVDF hydrophones. The photoacoustic fiber probe also successfully took B-mode images of the blood vessel phantoms when combined with a single optical fiber for excitation of photoacoustic waves. We fabricated a 3D photoacoustic imaging system that consists of a photoacoustic fiber probe and a bundle of hollow optical fibers. Owing to the extremely small NA of the hollow fiber, an image resolution that is the same as the diameter of the hollow optical fiber was obtained. In addition, without any scanning mechanism at the distal end, 3D imaging can be performed by subsequently exciting the hollow fibers at the input end of the hollow fiber bundle. After some image processing, a 3D image of the blood vessel phantom with an inner diameter of 1 mm was successfully reconstructed. The image resolution was around 250 μm in depth direction and 500 μm in lateral direction. We are working on fabrication of the fiber acoustic probe accepting higher frequency acoustic wave to improve resolution in depth direction. In addition, by using hollow optical fibers with a much smaller diameter, higher resolution imaging in lateral direction will be performed. These small diameter probes can be inserted into the working channel of the endoscope or small-bore catheter although transmission losses of small-bore hollow fibers should be further reduced.
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