2.1.

Experimental Setup

The details of the UOM experimental setup (Fig. 1 ) based on CFPI ( $50\text{-}\mathrm{cm}$ cavity length, $0.1\text{-}{\mathrm{mm}}^2\mathrm{sr}$ etendue, and $\sim 20$ finesse) have been described earlier. The CFPI is used as a real-time optical filter for filtering multiply scattered light modulated by high frequency ultrasound pulses propagating in biological tissue. Figure 1 shows the configuration of the sample S that is imaged, as well as the oblique configuration of the light illumination and ultrasound insonification of the sample S. The axis of propagation of the ultrasonic beam is along the $z$ axis. The light beam propagates about $35\mathrm{deg}$ with respect to the ultrasonic beam. The ultrasonic and optical beams were focused to the same spot below the sample surface. We used a high frequency focused ultrasound transducer with a central frequency of $30\mathrm{MHz}$ (Panametrics, $4.25\text{-}\mu \mathrm{s}$ fused silica delay lines, $6.35\text{-}\mathrm{mm}$ element sizes, $5.5\text{-}\mathrm{mm}$ focal lengths, and 80% estimated fractional bandwidths). This focused transducer is modified from a commercial broadband delay-line transducer (V213-BB-RM, Panametrics, Waltham, Massachusetts) by grinding a concave spherical lens directly in its quartz delay line. This concave lens works as a focusing acoustic lens. The lateral resolution of UOM in biological tissues is mainly determined by the center frequency and the numerical aperture (NA) of the ultrasound transducer, and the axial resolution is primarily determined by the frequency bandwidth of the ultrasound transducer. The center frequency also limits the imaging depth due to the frequency-dependent ultrasonic attenuation, $0.7–3\mathrm{dB}∕\left(\mathrm{cm.}\mathrm{MHz}\right)$ in human skin. Our earlier experiments, using an ultrasound transducer with a $30\text{-}\mathrm{MHz}$ central frequency, showed that an axial resolution of about $60\mu \mathrm{m}$ and a lateral resolution of $70\mu \mathrm{m}$ could be achieved in a depth range of about $2.5\mathrm{mm}$ in chicken breast tissue.²⁹ As in the earlier work, in these experiments the transducer was driven by a square bipolar pulse with a period of $34\mathrm{ns}$ . This bipolar pulse was generated by a trigger generator (DG535 Stanford Research, Sunnyvale, California) and amplified (Amplifier Research, 75A250). The focal ultrasound pressure measured by a needle hydrophone (ONDA HNV-0200) was about $3\mathrm{MPa}$ , within the ultrasound safety limit at these frequencies for tissues without well-defined gas bodies.³¹

Fig. 1

(a) Schematic of the experimental setup: APD, avalanche photodiode; BS, beamsplitter; CF, collecting fiber; CFPI, confocal Fabry-Perot interferometer; CO, coupling optics; DAQ, data acquisition board; FO, focusing optics; M1 and M2, mirrors inside the cavity; PD, photodetector; PDG, pulse delay generator; PZT, piezoelectric transducer; PZT-D, piezoelectric transducer driver; RF Amp, RF amplifier; S, sample; SH, shutter; TG, trigger generator; and UT, ultrasonic transducer.

In this experiment, the laser light (Coherent, Verdi, $532\text{-}\mathrm{nm}$ wavelength) was focused onto a spot $\sim 100\mu \mathrm{m}$ in diameter below the surface of an otherwise scattering-free sample. The optical power delivered to the sample was about $50\mathrm{mW}$ . However, for in vivo experiments, the duration of the light exposure to the living subject could be reduced to only a few $\mu \mathrm{s}$ for each cycle of ultrasound propagation through the region of interest, to meet the ANSI safety limit for average power.³² Diffusely reflected light was collected by a multimode optical fiber with a $600\text{-}\mu \mathrm{m}$ core diameter. The sample was mounted on a three-axis motorized translational stage and controlled by the LabView program. The ultrasound transducer (UT) and the sample S were immersed in water for acoustic coupling, and thus the light focusing optics (FO) and the collecting fiber (CF) were also immersed in the same water tank. The collected light was coupled into the CFPI using coupling optics (CO), and was operated in transmission mode. Details about the subsequent ultrasound-modulated optical signal detection using the CFPI were discussed in our earlier work.²⁹ The photodiode (PD) (New Focus, model 2031, New Focus, San Jose, California) was used in a cavity alignment procedure. The signal acquired by an avalanche photodiode (APD, Advanced Photonix, Ann Arbor, Michigan) during the ultrasound propagation through the sample represented the distribution of the ultrasound-modulated optical intensity along the ultrasonic axis. Therefore, it yielded a 1-D image (A-line) along the $z$ direction. In each operational cycle, the resonant frequency of the CFPI was tuned first, and then data from 4000 ultrasound pulses were acquired in one second. Averaging of $3\text{to}5\text{cycles}$ was usually needed to obtain a satisfactory signal-to-noise ratio (SNR) for each 1-D image. 2-D images were obtained by scanning the sample along the $z$ direction and acquiring each corresponding 1-D image.

After each B-scan along the $x$ axis, the sample was returned to its original $x$ coordinate and further translated by one scanning step (usually $0.1\mathrm{mm}$ ) along the $y$ axis to start a new B-scan. Thus several B-scans were obtained by raster scanning the sample in the $x\text{-}y$ plane. To obtain high contrast blood vessel imaging, each B-scan was processed by subtracting each A-line in the B-scan from an A-line obtained from the transparent region of the respective B-scan. We further performed maximum intensity projection (MIP) on each of the processed B-scan images along the $z$ axis. The final MIP image was obtained by combining projections from all B-scans.

2.2.

Phantom Preparation

Ultrasound-modulated optical imaging is primarily aimed at imaging intrinsic optical contrast of soft biological tissue. Therefore, we prepared several objects of different absorption coefficients (including a nearly transparent object), and measured the ultrasound-modulated optical signal with our UOM setup, presented in Fig. 1. The four objects were made of 15% porcine gelatin, with different amounts of Trypan Blue dye, so that the optical absorption coefficients at $532\mathrm{nm}$ were 145, 72, 36, and $0.1{\mathrm{cm}}^{-1}$ (whole blood absorption at $532\mathrm{nm}$ is around $120{\mathrm{cm}}^{-1}$ ). The approximate dimensions of all four objects were 1, 2, and $1\mathrm{mm}$ in the $x$ , $y$ , and $z$ directions. These objects were transparent enough for the ultrasound and absorptive for the light. The background tissue phantom was made from gelatin, water, and 1% Intralipid (20% Liposyn II, Intravenous Fat Emulsion Hospira, Incorporated, Lake Forest, Illinois) with a $1\text{-}\mathrm{mm}$ optical transport mean free path. The tissue phantom was an optically scattering slab $10\mathrm{cm}$ wide in the $x$ and $y$ axes (the ultrasound scanning axes, or the B-scan and C-scan directions, respectively) and $5\mathrm{mm}$ thick in the $z$ axis (ultrasound propagation direction). The optical reduced scattering and absorption coefficients of the scattering slab (background medium) were ${\mu }_s^{\prime }=10{\mathrm{cm}}^{-1}$ and ${\mu }_a=0.1{\mathrm{cm}}^{-1}$ , respectively. As shown in Fig. 2a, a piece of stainless steel needle ( $254\text{-}\mu \mathrm{m}$ outer diameter) and all four objects were placed inside the scattering slab at a depth of $z=3\mathrm{mm}$ . Figure 2b shows a photograph of the prepared phantom. Since ultrasound waves were strongly reflected by the needle, it helped in aligning all four objects precisely at the ultrasound focal point. We used a pulser (GE Panametrics, 5072PR) in this alignment procedure.

Fig. 2

Dependence of ultrasound-modulated optical signal on optical properties. (a) Photograph of three optically absorbing objects ( ${\mu }_a=145{\mathrm{cm}}^{-1}$ , ${\mu }_a=72{\mathrm{cm}}^{-1}$ , ${\mu }_a=36{\mathrm{cm}}^{-1}$ , and all three having the same ${\mu }_s^{\prime }=0.5{\mathrm{cm}}^{-1}$ ) and one transparent object ( ${\mu }_a=0.1{\mathrm{cm}}^{-1}$ , ${\mu }_s^{\prime }=0.5{\mathrm{cm}}^{-1}$ ) placed vertically below a piece of needle, respectively. These four objects are marked as O-1, O-2, O-3, and O-4, respectively. (b) Photograph of the prepared phantom sample. (c) Background subtracted B-scan UOM image of the sample shown in (a). (d) Respective 1-D profiles (A-lines) along the $z$ axis through the center of all four objects. (e) Normalized A-lines, acquired from the original (not background subtracted) data, through the centers of all four objects and the background medium BG. (f) Plot of maximum contrast of all three absorbing objects.

We obtained several ear and skin tissue samples from a mouse and a rat sacrificed during other experiments in our laboratory. As shown in Figs. 3a and 4a , a piece of stainless steel needle ( $254\text{-}\mu \mathrm{m}$ outer diameter) and these tissue samples were then placed inside an optically clear medium (made from gelatin and water) at a depth of $z=2.5\mathrm{mm}$ (the ultrasound propagation direction). This clear medium was $10\mathrm{cm}$ wide in the $x$ and $y$ axes (the ultrasound scanning axes, or the B-scan and C-scan directions, respectively) and $5\mathrm{mm}$ thick in the $z$ direction.

Fig. 3

Imaging intact blood vasculature in a rat ear. (a) Photograph of blood vasculature inside the rat ear. (b) A-line, A-1, represents temporal dependence of ultrasound-modulated light intensity during the propagation of ultrasound pulse through the ear tissue sample. A dot placed below this curve in the figure represents the position of the blood vessel along the ultrasound propagation direction. A-2 represents a typical A-line not passing through the blood vessel. (c) through (f) Four consecutive B-scan images that were obtained $0.2\mathrm{mm}$ apart laterally at $y=0.2\mathrm{mm}$ , $y=0.4\mathrm{mm}$ , $y=0.6\mathrm{mm}$ , and $y=0.8\mathrm{mm}$ , respectively. (g) MIP image of (a) obtained using UOM.

Fig. 4

Imaging blood vasculature with mouse skin intact. (a) Photograph of the exterior of the skin sample obtained from the mouse. (b) The vascular distribution intact with the skin sample photographed from the underside. (c) and (d) Two consecutive B-scan images that were obtained $0.2\mathrm{mm}$ apart laterally at $y=1.5\mathrm{mm}$ and $y=1.7\mathrm{mm}$ , respectively. (e) MIP image of (a) obtained using UOM.

3.1.

Ultrasound-Modulated Optical Signal Dependence on Optical Properties

The contrast in ultrasound-modulated optical imaging is primarily based on optical properties of the biological tissue. Although both scattering and absorption influence the strength of the signal, variations of the ultrasound-modulated intensity are more sensitive to absorption.³³ Therefore, it is of primary importance to explore the dependence of the ultrasound-modulated optical signal on optical properties using the UOM system shown in Fig. 1. As described in the previous section, we prepared an optically scattering sample and embedded three optically absorbing objects ( ${\mu }_a=145{\mathrm{cm}}^{-1}$ , ${\mu }_a=72{\mathrm{cm}}^{-1}$ , and ${\mu }_a=36{\mathrm{cm}}^{-1}$ , with all three having the same ${\mu }_s^{\prime }=0.5{\mathrm{cm}}^{-1}$ ) and one nearly transparent object ( ${\mu }_a=0.1{\mathrm{cm}}^{-1}$ , ${\mu }_s^{\prime }=0.5{\mathrm{cm}}^{-1}$ ), $3.0\mathrm{mm}$ below the surface of the sample. Photographs of the objects and the prepared phantom are shown in Figs. 2a and 2b, respectively. Figure 2c shows a background subtracted 2-D (B-scan) image of all four objects shown in Fig. 2a, obtained by scanning the sample in the $x$ direction. Respective 1-D profiles (A-lines) along the $z$ axis through the center of all four objects are plotted in Fig. 2d. Figures 2c and 2d demonstrate that the image contrasts of all three absorbing objects are negative, whereas the image contrast of the transparent object is positive. Further, the image contrast of all three absorbing objects decreases with a decrease in absorption coefficient.

Figure 2e presents normalized 1-D profiles (A-lines), acquired from the original data, along the $z$ axis through the center of all four objects and the background medium BG. These A-lines represent the temporal dependence of the ultrasound modulated light intensity during ultrasound-pulse propagation through the sample. For each A-line, the time of propagation is multiplied by $1500{\mathrm{ms}}^{-1}$ , the approximate speed of sound in the sample, to be converted to distance along the $z$ axis, where the origin of each A-line corresponds to the trigger for the signal acquisition from the APD. For the given acoustic parameters, these results demonstrate that the characteristics of the ultrasound-modulated optical signal were influenced by the size and the absorption coefficient of the object and by the scattering properties of the objects and surrounding medium. As shown in Fig. 2e, both the width and height of the signal drop decrease with a decrease in the absorption coefficient of the object. This decrease in the width of the signal drop can be attributed to increased diffusion of photons (into the object) with a decrease in the absorption coefficient of the object. The decrease in the height of the signal drop with an increase in the absorption coefficient of the object can be directly related to an increase in the absorption of photons. In addition, the comparison between A-lines through the centers of the transparent object ( ${\mu }_s^{\prime }=0.5{\mathrm{cm}}^{-1}$ and ${\mu }_a=0.1{\mathrm{cm}}^{-1}$ ) and the background medium ( ${\mu }_s^{\prime }=10{\mathrm{cm}}^{-1}$ and ${\mu }_a=0.1{\mathrm{cm}}^{-1}$ ) shows that the shape of the ultrasound-modulated optical signal (A-line) is quite sensitive to the local (confined by the ultrasound focal volume) optical scattering properties relative to the scattering properties of the background medium. Figure 2c also shows that the intensity of modulated light passing through the transparent object is more than that through the background medium. Since the transparent object and the background have the same ${\mu }_a$ , this increase in the intensity and the change in the shape of the A-line passing through the center of the transparent object can be attributed to the lower ${\mu }_s^{\prime }$ of the transparent object than that of the background.

We further measured the contrast at the center of all three absorbing objects in Fig. 2c, and plotted the results in Fig. 2f. This measurement demonstrates that the contrast of each object was nearly linearly related to the absorption coefficient of the object. Simple theoretical considerations show that at the limit, where the absorption length (reciprocal of the absorption coefficient) is small compared to the path length of the light through the absorbing object, the dependence of the ultrasound-modulated signal intensity on the object absorption is nearly linear.³³ More importantly, the measurable absorptions are clearly much lower than the blood absorption ( ${\mu }_a=120{\mathrm{cm}}^{-1}$ at $532\text{-}\mathrm{nm}$ wavelength), which implies that this technology has great potential in detecting small blood vessels and other functional parameters associated with blood, such as oxygenation saturation.

3.2.

Ex Vivo Blood Vessel Imaging Using Tissue Samples