2.1.

Probe Design

A low-loss light delivery and ultrasound detection system is required to enable detection of weak phtotoacoustic signals. In 2008, Maslov ⁹ presented a probe with two prisms. This probe successfully achieved focusing gain, minimal path length in water/tissue, and negligible acoustic attenuation in fused quartz. However, owing to acoustic impedance mismatch, acoustic losses at water-fused quartz interfaces can be high. Additionally, since longitudinal waves are converted to shear waves at liquid–solid interfaces, wave-mode conversion at a prism–fluid $45\text{-}\mathrm{deg}$ interface produces further losses. In 2009, Z. Xie, ¹⁰ designed a probe technique that enabled laser scanning in a photoacoustic microscopy system. They used an unfocused transducer obliquely positioned to receive photoacoustic signals. As a result of the long path length in water and the lack of any focal gain, the acoustic and diffractive losses are not negligible, and a low-frequency transducer was used, which results in a sacrifice in depth resolution.

In this study, we designed a unique probe with minimal loss, which is illustrated in Fig. 1 . This probe is somewhat similar to the light delivery system described by Maslov ⁹ and is adapted from our previous work.¹⁴ Our ultrasound transducer faces an oblique $10\text{-}\mathrm{mm}$ fused silica prism such that photoacoustic signals directed upward to the prism’s diagonal will be deflected to the transducer.¹⁴ Figure 1 presents the calculation of acoustic reflectivity of prism versus incident acoustic signal angle. The calculation result shows the reflectivity is almost 100% for incident acoustic signal angles within $45\pm 13$ . $4\mathrm{deg}$ (the angular acceptance angles of the transducer). Therefore, all of the acoustic energy is preserved at the prism interface. In our design, optical index-matching fluid (Catalog No. 19569, Cargille Labs, Cedar Grove, New Jersey) is used as ultrasonic coupling, and hence allows for top-down laser illumination to be directed to the tissue surface without optical refractive path variation. The loss caused by acoustic attenuation in index-matching fluid is only slightly more than water.¹⁴ Our probe therefore exhibits minimal loss, which will be critical for sensitive photoacoustic imaging.

Fig. 1

(a) Minimal-loss probe design (OL: objective lens; P: prism; IMF: index-matching fluid; UST: ultrasound transducer; M: $\sim 25\text{-}\mu \mathrm{m}$ -thick Saran Wrap membrane; F: focal point of both light and ultrasound). Ultrasound transducer faces a downfacing optical prism such that photoacoustic signals directed upward to the prism’s diagonal will be deflected to the transducer. An optical index-matching fluid ensures that light can be focused without refraction at the prism’s diagonal interface and provides a low acoustic loss medium for acoustic signal propagation. (b) Computed curve of acoustic power reflection coefficient $R_w$ (i.e., fraction of incident power reflected) versus incident angle of ultrasound signal on prism. This calculation was made using the acoustic impedances of the fused silica and index-matching fluid media, as based on the speed of sound in, and density of, these materials. Vertical dashed lines represent the acceptance angles of our transducer (about the optical or acoustic axis).

In our probe design, to focus the light to a diffraction-limited spot, we need to consider the optical aberration, especially due to the prism and matching fluid. In this study, we used the ray-tracing software Beam4 (Stellar Software, Inc., Berkley, California) to analyze the beam waist as it is focused through a $1\text{-}\mathrm{cm}$ slab (modeling the prism and index fluid). In Fig. 2 , the plot based on the Beam4 data is shown for a doublet lens ( $\mathrm{f}=18\mathrm{mm}$ at $532\mathrm{nm}$ ). While the input beam diameter to the lens increases, aberrations due to the prism-fluid layer become cubically worse, as measured by the full-width-half-maximum spot size of the beam at the focal waist (dashed line). On the other hand, in a diffraction-limited system without aberrations, the focal spot size decreases with increasing numerical aperture (or input beam diameter), as shown by the solid curve. Diffraction-limited performance may be nearly achieved in the neighborhood of the intersection of these curves.

Fig. 2

To study the effects of the prism-index-fluid layer of Fig. 1 on the potential optical focusing power of the system, we performed a ray-tracing simulation. We simulated a doublet lens ( $\mathrm{f}=18\mathrm{mm}$ , $\lambda =532\mathrm{nm}$ ) focusing a top-hat beam through a $1\text{-}\mathrm{cm}$ prism-index-fluid slab with air between the lens and the slab. The dashed curve is the spot size at the focus versus input beam diameter. The solid curve represents the spot size possible according to diffraction theory if no aberrating layer were present. The intersection of the curves represents a region where diffraction-limited performance is possible.

2.2.

Experimental Setup with Microchip Laser

The experimental setup for the photoacoustic imaging system employing our microchip laser is illustrated in Fig. 3 . The output mode diameter of the $808\text{-}\mathrm{nm}$ pump laser was $\sim 100\mu \mathrm{m}$ . Lens 1 $(\mathrm{f}=8\mathrm{mm})$ was used for pump collimation. The collimated pump light was focused onto the Nd: Cr: YAG microchip, which includes the gain medium and saturable absorber (SA) using lens 2 $(\mathrm{f}=11\mathrm{mm})$ . A long-pass filter 1 was used to block the $808\text{-}\mathrm{nm}$ pump light, while lenses 3 and 4 $(\mathrm{f}=10\mathrm{cm})$ were used to focus the laser onto a potassium titanium oxide phosphate (KTP) frequency-doubling crystal. The KTP crystal produces a green $532\text{-}\mathrm{nm}$ wavelength output, along with residual fundamental light. This $532\text{-}\mathrm{nm}$ wavelength exhibits strong hemoglobin absorption contrast. Although we did not need it for our imaging experiments, we demonstrated amplification of the $1\text{-}\mathrm{ns}$ microchip laser pulses to $60\mu \mathrm{J}$ using a large mode-area Yb-doped fiber amplifier. The pulse repetition rate was determined by the power of the microchip laser pump source at $808\mathrm{nm}$ and may exceed $10\mathrm{kHz}$ . After collimating the output light from the KTP frequency doubling crystal by lens 5 $(\mathrm{f}=30\mathrm{cm})$ , a glass slide was used to reflect a fraction of the light onto a photodiode, which was connected to an oscilloscope (DPO 7054, Digital Phosphor Oscilloscope, Tektronix) as a trigger signal input. Filter 2 was inserted before the scanning mirror system to block incident light at $1064\mathrm{nm}$ . The scanning mirror system (6230H, Cambridge Technology, Inc.) is driven by a function generator (AFG3101, Tektronix, Inc.). It can reach $300\mathrm{Hz}$ at mechanical angles of $\pm 5\mathrm{deg}$ , although we used much lower rates $\left(3\text{to}5\mathrm{Hz}\right)$ for our experiments at this time to ensure dense sampling. An objective lens ( $\mathrm{f}=18\mathrm{mm}$ , $\mathrm{NA}=0.15$ , K16033703, Mitutoyo Co.) was used to focus light through the prism and index-fluid and was positioned $\sim 11\mathrm{cm}$ away from the scanning mirrors. The minimal-loss probe with a $10\text{-}\mathrm{MHz}$ ultrasound transducer ( $\mathrm{f}=19\mathrm{mm}$ , $6\text{-}\mathrm{mm}$ active element, $\mathrm{f}\#=3.17,$ CD International, Inc.) was used for photoacoustic signal detection. The photoacoustic signals were amplified by $54\mathrm{dB}$ using an ultrasound pulse-receiver (5900PR, Olympus NDT, Inc.). The signals from the photodiode and ultrasound pulse-receiver and the position-feedback signals from the scanning mirror system were recorded by the oscilloscope. The data acquired by the oscilloscope were analyzed by a MATLAB program (Mathworks, Inc.). Because the pulse-repetition intervals of the passively Q-switched lasers are somewhat variable, it is important to know where each laser pulse occurs relative to the scanning mirror motion trajectory. To form an image, we register the position of photoacoustic A-scan lines based on the time-location of photo-diode-recorded laser pulses relative to the mirror position-feedback signals. A calibration curve relating laser-spot location and mirror position (assessed by the feedback signal, which is a function of voltage) is formed by stepping the mirror angle to several fixed intervals and then using a micron-precision translation stage to move the position of a carbon-fiber target to the light spot to maximize the photoacoustic signal. Figure 3 shows a photograph of the scanning mirror, objective lens, and probe.

Fig. 3

(a) Experimental setup for the photoacoustic imaging system employing our microchip laser (ML: microchip laser; DL: diode pump laser; L1: $8\text{-}\mathrm{mm}$ lens; L2: $11\text{-}\mathrm{mm}$ lens; A: Nd:Cr:YAG microchip; F1: long-pass filter blocking $808\text{-}\mathrm{nm}$ pump light; L3: $10\text{-}\mathrm{cm}$ lens; M: mirror; L4: $10\text{-}\mathrm{cm}$ lens; KTP: KTP frequency-doubling crystal; L5: $30\text{-}\mathrm{cm}$ lens; G: glass; PD: photodiode; F2: short-pass filter blocking $1064\text{-}\mathrm{nm}$ light; SM: scanning mirror; OL: objective lens; Pr: imaging probe; UST: ultrasound transducer). A diode-pumped Nd: Cr: YAG micro-chip laser is frequency-doubled by a KTP frequency-doubling crystal to $532\text{-}\mathrm{nm}$ output. (b) Photograph of the fast-scanning mirror, objective lens, and imaging probe.

2.3.

Passively Q-Switched Fiber Laser Experimental Setup

Figure 4 shows the experimental setup for the photoacoustic imaging system employing a novel passively Q-switched fiber laser. As described in previous work by Pan, ¹² the active medium in our fiber laser was a Yb-doped single-mode double-clad (SM DC) fiber with core and inner cladding diameters of 6 and $125\mu \mathrm{m}$ and core and inner cladding numerical apertures of 0.15 and 0.46, respectively. In our study, however, the length of Yb-doped SM DC fiber was around $5\mathrm{m}$ instead of $3\mathrm{m}$ as in previous work, which resulted in a slightly longer pulse duration of $\sim 250\mathrm{ns}$ . The $976\text{-}\mathrm{nm}$ pump light from the fiber-coupled diode pump laser ( $25\text{-}\mathrm{W}$ maximum output power) was coupled to the Yb fiber by dichroic mirror 1 with 96% transmission at $976\mathrm{nm}$ and 99% reflectance at $1075\mathrm{nm}$ . The SA used in this experiment was a Cr:YAG crystal with transmission of 30% at $1075\mathrm{nm}$ . One broadband high-reflectance dielectric mirror and the perpendicularly cleaved fiber end face of the laser medium composed a laser resonator. The other end of the fiber was angle cleaved to avoid feedback. Lens 1 $(\mathrm{f}=11\mathrm{mm})$ was used for pump light collimation. Lens 2 $(\mathrm{f}=8\mathrm{mm})$ was used to couple the pump light into the active medium. Lens 3 $(\mathrm{f}=6.24\mathrm{mm})$ was a collimation lens at the other end of the fiber. Lens 4 and lens 5 $(\mathrm{f}=15\mathrm{mm})$ were used to focus the laser onto the SA. In our experiment, as shown in Fig. 4, the $1075\text{-}\mathrm{nm}$ output pulses reflected from the dichroic mirror were frequency-doubled with a KTP crystal. Lens 6 $(\mathrm{f}=10\mathrm{cm})$ was used to focus the laser onto the KTP crystal. Before collimating the output light from the KTP crystal by lens 7 $(\mathrm{f}=15\mathrm{cm})$ , dichroic mirror 2 was inserted to transmit $1075\text{-}\mathrm{nm}$ and reflect $537\text{-}\mathrm{nm}$ light. The $1075\text{-}\mathrm{nm}$ light was detected by a photodiode, which was connected to an oscilloscope for trigger signal input. The pulse-repetition rate was determined by the power of the fiber laser pump source at $975\mathrm{nm}$ and was $100\mathrm{kHz}$ for $4.1\text{-}\mathrm{W}$ coupled pump power in our experiments. Repetition rates from this laser system up to $300\mathrm{kHz}$ are demonstrated in Ref. 12. Figure 4 shows the fiber laser signal. The minimal-loss probe with a $10\text{-}\mathrm{MHz}$ ultrasound transducer was used for photoacoustic signal detection.

Fig. 4

(a) Experimental setup for the photoacoustic imaging system employing a fiber laser (PL: pump laser; Fb: fiber; L1: $11\text{-}\mathrm{mm}$ lens; DM1: dichroic mirror reflecting $1075\text{-}\mathrm{nm}$ light; L2: $8\text{-}\mathrm{mm}$ lens; DF: Yb-doped single mode fiber; L3: $6.24\text{-}\mathrm{mm}$ lens; L4: $15\text{-}\mathrm{mm}$ lens; SA: saturable absorber; L5: $15\text{-}\mathrm{mm}$ lens; HRM: high-reflectivity mirror; L6: $10\text{-}\mathrm{cm}$ lens; KTP: KTP frequency-doubling crystal; DM2: dichroic mirror reflecting $537\text{-}\mathrm{nm}$ light, L7: $15\text{-}\mathrm{cm}$ lens; SM: scanning mirror; OL: objective lens; Pr: imaging probe; UST: ultrasound transducer). A Yb-doped fiber of $\sim 6.5\text{-}\mu \mathrm{m}$ mode diameter with an external saturable absorber is pumped by a $976\text{-}\mathrm{nm}$ diode laser. The $1075\text{-}\mathrm{nm}$ output pulses are coupled out of the laser using a dichroic mirror and then frequency-doubled with a KTP crystal. The minimal-loss probe with a $10\text{-}\mathrm{MHz}$ ultrasound transducer was used for photoacoustic signal detection. (b) Fiber laser signals as measured by a fast photodiode and oscilloscope: $250\text{-}\mathrm{ns}$ pulse duration, $100\text{-}\mathrm{kHz}$ repetition rate.

3.1.

Microchip Laser Experimental Results

Without a fiber-amplifier stage, the microchip laser produced $\sim 1.84\text{-}\mu \mathrm{J}$ , $0.6\text{-}\mathrm{ns}$ , $1064\text{-}\mathrm{nm}$ laser pulses with $7.4\text{-}\mathrm{kHz}$ repetition rate at $\sim 1.8\text{-}\mathrm{W}$ pump power. After frequency doubling and filtering, we obtained $0.21\text{-}\mu \mathrm{J}$ , $0.6\text{-}\mathrm{ns}$ , $532\text{-}\mathrm{nm}$ laser pulses, measured after the scanning mirrors and before the objective lens. With our objective lens, the calculated light spot size at the focus is $\sim 6\mu \mathrm{m}$ . A phantom study was performed using $7.5\text{-}\mu \mathrm{m}$ carbon fiber targets positioned $1.5\mathrm{mm}$ below the probe membrane. To mimic the turbid medium of human tissue (which has a reduced scattering coefficient ${\mu }_s^{\prime }$ of around $10{\mathrm{cm}}^{-1}$ ) at $0.5\text{-}\mathrm{mm}$ imaging depths, a solution of Intralipid was created with ${\mu }_s^{\prime }=3.33{\mathrm{cm}}^{-1}$ to be used at the $1.5\text{-}\mathrm{mm}$ depth. This solution was based on a reference Intralipid concentration of 20% with measured ${\mu }_s^{\prime }$ of $334{\mathrm{cm}}^{-1}$ . Oblique-incidence reflectometry was used to perform this scattering measurement.¹⁵ It can be seen from Fig. 5 that the measured photoacoustic signal amplitudes were around $1.9\mathrm{V}$ , $1.2\mathrm{V}$ , and $0.9\mathrm{V}$ , in clear water, turbid media with ${\mu }_s^{\prime }=3.33{\mathrm{cm}}^{-1}$ , and turbid media with ${\mu }_s^{\prime }=6.67{\mathrm{cm}}^{-1}$ , respectively. Photoacoustic signal FWHMs (full width at half maximums) were $8\pm 2\mu \mathrm{m}$ , $11\pm 2\mu \mathrm{m}$ , and $7\pm 2\mu \mathrm{m}$ in the aforementioned media. The mean photoacoustic signal FWHM in our study (for media ranging from water to turbid media with ${\mu }_s^{\prime }$ up to $7.94{\mathrm{cm}}^{-1}$ ) is $9\pm 2\mu \mathrm{m}$ . The measured widths of curves shown in Fig. 5 are partly due to the width of the carbon fibers themselves. To estimate the true optical resolution, we computed the convolution of a 2-D Gaussian beam with a carbon fiber (simulated as a 2-D rectangular absorption region). We plot the true FWHM of the Gaussian beam versus the FWHM of the convolution result in Fig. 6 , which shows that the corresponding lateral Gaussian probe spot resolution is $7\pm 2\mu \mathrm{m}$ , which is within error of the predicted spot size of $6\mu \mathrm{m}$ .

Fig. 5

Photoacoustic signals (circles) at $1.5\text{-}\mathrm{mm}$ depth below the probe membrane with Gaussian fits (lines) (a) in clear water, $\mathrm{FWHM}=7.9\mu \mathrm{m}$ , $\text{amplitude}=1.9\mathrm{V}$ ; (b) in turbid medium with ${\mu }_s^{\prime }=3.33{\mathrm{cm}}^{-1}$ , $\mathrm{FWHM}=11.3\mu \mathrm{m}$ , $\text{amplitude}=1.2\mathrm{V}$ ; and (c) in turbid medium with ${\mu }_s^{\prime }=6.67{\mathrm{cm}}^{-1}$ , $\mathrm{FWHM}=7.5\mu \mathrm{m}$ , $\text{amplitude}=0.9\mathrm{V}$ .

Fig. 6

Optical resolution as a function of full width at half maximum (FWHM) of measured photoacoustic signal. This plot was generated by taking the convolution of a 2-D Gaussian with a rectangle (simulating the beam spot and carbon fiber, respectively). The photoacoustic FWHM is larger than the corresponding beam FWHM due to the nonzero width $\left(7\text{to}8\mu \mathrm{m}\right)$ of the carbon fiber targets; the true optical resolution of the system is found to be $7\mu \mathrm{m}$ for the measured photoacoustic FWHM of $9\mu \mathrm{m}$ .

3.2.

Fiber Laser Experimental Results

In our study, the passively Q-switched fiber laser produced $\sim 13\text{-}\mu \mathrm{J}$ , $250\text{-}\mathrm{ns}$ , $1075\text{-}\mathrm{nm}$ laser pulses with $100\text{-}\mathrm{kHz}$ repetition rates at $4.1\text{-}\mathrm{W}$ coupled pump power. The output mode diameter of the single-mode Ytterbium-doped fiber is around $6.5\mu \mathrm{m}$ . After intentional attenuation (via a neutral density filter) and measuring before the objective lens, we obtain a mere $0.06\text{-}\mu \mathrm{J}$ , $250\text{-}\mathrm{ns}$ pulse of $537\text{-}\mathrm{nm}$ laser light. Using an objective lens $(\mathrm{f}=18\mathrm{mm})$ , the calculated light spot size at the focus is $\sim 10\mu \mathrm{m}$ for the fiber laser system. Due to lower peak power, the signal-to-noise ratio of the photoacoustic signals was lower for the fiber laser than for the microchip laser, and hence we chose to use an edge-spread target, since more samples could be obtained than from the carbon fiber target. To measure the resolution of the system, we plot the edge-spread function of the photoacoustic signal, shown in Fig. 7 , as the optical spot is scanned relative to a graphite foil sample. From the plot, we can see that the photoacoustic resolution is $15\pm 5\mu \mathrm{m}$ , which is significantly better than the ultrasonic lateral resolution of $\sim 580\mu \mathrm{m}$ . Measurements were performed in both water and Intralipid with similar resolution results; however, the results acquired in water are much less noisy and are those displayed in Fig. 7.

Fig. 7

Edge-spread function of photoacoustic signal from graphite foil; lateral 10% to 90% PA resolution is $15\pm 5\mu \mathrm{m}$ .