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Abstract
Simultaneous photoacoustic (PA) and ultrasound (US) imaging provides rich optical and acoustic contrasts with high sensitivity, specificity, and resolution, making it a promising tool for diagnosing and assessing various diseases. However, the resolution and penetration depth tend to be contradictory due to the increased attenuation of high-frequency ultrasound. To address this issue, we present simultaneous dual-modal PA/US microscopy with an optimized acoustic combiner that can maintain high resolution while improving the penetration of ultrasound imaging. A low-frequency ultrasound transducer is used for acoustic transmission, and a high-frequency transducer is used for PA and US detection. An acoustic beam combiner is utilized to merge the transmitting and receiving acoustic beams with a predetermined ratio. By combining the two different transducers, harmonic US imaging and high-frequency photoacoustic microscopy are implemented. In vivo experiments on the mouse brain demonstrate the simultaneous PA and US imaging ability. The harmonic US imaging of the mouse eye reveals finer iris and lens boundary structures than conventional US imaging, providing a high-resolution anatomical reference for co-registered PA imaging.
© 2023 Optica Publishing Group under the terms of the Optica Open Access Publishing Agreement
1. Introduction
PA and US microscopy both detect acoustic signals, allowing them to be readily integrated into a single system. Photoacoustic microscopy (PAM) maps the optical absorption coefficient [1–15], while ultrasound microscopy reveals acoustic impedance mismatch [15–19]. By combining these two modalities, integrated PA/US microscopy can identify co-registered structural, functional, and molecular features with high resolution and sensitivity, making it an invaluable biomedical imaging tool.
Dual-modal PA/US microscopy has been developed by two separate approaches. The first is to use a single-element transducer for both PA imaging (PAI) and US imaging (USI) [20–22], with a transmit/receive switch used to alternate between the two modes [23–28]. If a low-frequency transducer is employed, the spatial resolution will be poor [29,30]. Utilizing a high-frequency transducer can improve spatial resolution, though strong ultrasound attenuation may lead to poor image quality in deep tissue [31,32]. Especially, the acoustic waves are attenuated in a round trip, severely limiting the use of high-frequency transducers. The second approach involves the use of a dual-element transducer, allowing for two different ultrasound bands to be employed for the PAI and USI, thus providing enhanced imaging ability. For instance, a dual-element transducer with both low and high frequencies can be used to acquire tissue harmonic imaging, achieving both deep penetration and high resolution. Harmonic USI, which is based on the acoustic nonlinear effect in tissue, offers higher resolution than fundamental frequency USI and is more suitable for identifying PA features. A fundamental frequency ultrasound beam can pass through tissue and undergo acoustic distortion, resulting in nonlinearity that generates super-harmonic echoes with higher frequency. High-resolution harmonic USI can be obtained by using a high-pass filter to suppress the fundamental echoes, leaving only the second or higher harmonic frequency components. Utilizing a dual-band US probe thus reduces attenuation due to the relatively deep penetration of the low-frequency transducer, while the improved image quality of the high-frequency transducer is provided through harmonic USI [33–35]. A remaining challenge is that the dual-element transducer is difficult to fabricate and not readily available at various frequency ranges [36,37].
Here, we develop new simultaneous dual-modal PA/harmonic US microscopy by combining two different ultrasound transducers. A low-frequency transducer is used for ultrasound transmission, while a high-frequency transducer is utilized for US and PA receiving. Tissue harmonic USI can achieve high resolution and deep penetration, offering anatomical references for PA features. The dual-band ultrasound beams are merged through an acoustic beam combiner, the acoustic transmittance/reflectance ratio of which is adjusted to achieve sufficient US excitation and sensitive PA/US detection. We characterize the imaging quality in phantom experiments and demonstrate the dual-modal imaging ability in vivo. In the mouse eye, harmonic ultrasound imaging allows us to identify fine anatomical features, such as iris and lens boundaries, providing a clear reference for the co-registered vascular features. Imaging of subcutaneous tumors reveals co-registered vessels, tumor boundaries, and tumor thickness.
2. Method
2.1 Dual-modal PA/harmonic US microscopy system
Figure 1(a) illustrates a schematic of the dual-modal PA/US microscopy system. For PA excitation, a nanosecond pulsed laser (532-nm wavelength, VPFL-G-20, Spectra-Physics) is employed as the light source, set to a pulse repetition rate of 4 kHz and a pulse width of 7 ns. A neutral density filter (NDF) is utilized to adjust the laser intensity. The optical beam is delivered to the imaging probe via a 1-meter multi-mode fiber (MMF, M43L01, Thorlabs Inc). In the probe, the optical beam is collimated by an achromatic lens (ACL, AC064-013-A, Thorlabs Inc), reflected by a mirror, then condensed by an achromatic objective (ACO, #33-202, Edmund Optics). The laser beam is reflected by an optical/acoustic combiner (OAC), passes through an acoustic lens (AL, AL#45-697, Edmund Optics), and illuminates the sample. The optical illumination spot at the focal depth is approximately 126 µm in diameter.
Fig. 1. (a) Schematic of simultaneous dual-modal PA/US microscopy. AC, acoustic combiner; ACL, achromatic lens; ACO, achromatic objective; AL, acoustic lens; FC, fiber coupler; MMF, multi-mode fiber; NDF, neutral density filter; OAC, Optical/acoustic combiner; UT, ultrasound transducer. (b) Time sequence for PAI and USI. DAQ, data acquisition card; Tx, transmission, Rx, reception.
For US transmission, an arbitrary function generator sends a sinusoidal signal to a 43-dB power amplifier (LZY022+, Mini-Circuits). The amplified signal drives a low-frequency transducer (V212-BC-RM, 20-MHz center frequency, 78% bandwidth, Olympus NDT) for US transmission. The ultrasound waves are reflected by an acoustic combiner (AC), then pass through an OAC and an acoustic lens, before being delivered to the sample.
The induced PA and US waves are collected by the acoustic lens, then transmit through the two combiners (OAC and AC) and detected by a high-frequency transducer (V214-BC-RM, 50-MHz center frequency, 78% bandwidth, Olympus NDT). The PA and US waves are amplified by 48 dB and acquired by a data acquisition card (ATS9360, Alazar Tech Inc). The probe is mounted to a three-axis translation stage for raster scanning, with the system sequentially acquiring one PA A-line and one US A-line at each scanning spot [38,39, 40]. The imaging sequence is controlled using an FPGA card (PCIe-7852, National Instruments). Figure 1(b) shows the sequence of the simultaneous PAI/USI. The laser is triggered first, then the DAQ is triggered after $\Delta {t_1}$ to acquire the PA signal. 8-µs after the laser trigger, the function generator is triggered to transmit a low-frequency US wave, after which a US signal is detected by the high-frequency transducer and the DAQ.
As shown in Fig. 2(a), the AC is composed of an aluminum-coated prism, a coupling medium layer, and an uncoated prism. To achieve both effective excitation and sensitive detection, different coupling medium layers and several different thicknesses of aluminum coating are investigated to determine the acoustic transmission/reflectance ratios. As depicted in Fig. 2(a), the transmitted and reflected photoacoustic waves are detected with two transducers. To increase the acoustic reflectivity, the aluminum coating is set to 1.5 µm and the acoustic reflectivities are measured at three different coupling media. The acoustic reflection/transmission ratio is defined as the ratio of the reflected/transmitted acoustic energy to the total energy before passing through the AC. As shown in Fig. 2(b), the acoustic reflectivity of the epoxy resin layer is lower than that of water or silicone oil, with the acoustic energy reflectivity of water and silicone oil both around 10%. For long-term use without drying out, silicone oil is selected as the coupling medium. To further increase the acoustic reflectivity, the coating thickness in AC is increased. Different thicknesses of aluminum coating are tested, as indicated by the red dots in Fig. 2(c). However, when the coating is too thick, around 5 µm, the transmission/reflection ratio is about 2 and the acoustic attenuation is too high, as seen in Fig. 2(c). Thus, a transmission/reflection ratio of ∼3 is expected to offer high-power US excitation, sensitive PA/US detection, and less acoustic attenuation, as represented by the blue line in Fig. 2(c). The uncoated AC attenuates the acoustic energy by roughly half compared to just one OAC. The acoustic attenuation increases with the coating thickness, and the 1.5-µm aluminum coating is close to the blue line, providing low acoustic attenuation. Therefore, 1.5-µm aluminum coating with silicone oil medium in the AC is chosen.
Fig. 2. (a) Schematic of AC and setup for reflectivity test; (b) Acoustic reflectivity of different coupling medium layers. The aluminum layer is 1.5-µm thick; (c) Acoustic reflectivity and transmission ratio with different thicknesses of aluminum coatings.
2.2 System characterization
We quantified the spatial resolutions by imaging a 6-µm-diameter carbon fiber submerged in water. Figure 3(a) shows a cross-sectional profile of the carbon fiber, with a lateral resolution of 45 µm, measured from the full width at half maximum (FWHM) [41]. The axial resolution was calculated from the FWHM of a PA envelope, which was derived from the absolute value of a Hilbert-transformed A-line signal [41–43]. As shown in Fig. 3(b), the axial resolution was 42 µm. Furthermore, imaging a tungsten wire in chicken breast tissue in Fig. 3(c) yielded a PA maximum penetration depth (MPD) of 2.94 mm.
Fig. 3. Spatial resolution quantification of simultaneous dual-modal PA/US microscopy. (a) PA lateral resolution, 45.0 µm. (b) PA axial resolution, 42.0 µm. (c) PA max penetration depth, 2.94 mm. scale bar: 1 mm. (d) Lateral resolution of fundamental and harmonic USI. (e) US axial resolution, 81 µm.
At USI, a series of 2-cycle sine waves (25 MHz) are transmitted to the low-frequency ultrasound transducer for transmission, and the backscattered waves are detected by both the low and high-frequency transducers. Figure 3(d) reveals a cross-sectional profile of the carbon fiber. Due to nonlinearity, the detected ultrasound signal is composed of fundamental and harmonic frequencies. The fundamental signal can be obtained using a digital low-pass filter (128th order, 28-MHz cutoff frequency). The FWHM of the fundamental US profile is 62 µm. To obtain the harmonic signal, a digital high-pass filter (128th order, 25-MHz cutoff frequency) is used, and the FWHM of the harmonic profile is 34.4 µm, which is nearly 55% of the fundamental one. Figure 3(e) shows that, in the axial direction, the FWHM of a US A-line signal (without filtering) is 81 µm.
3. Results
3.1 PAI/USI of the mouse brain
We used simultaneous dual-modal PA/US microscopy to investigate the mouse brain. The protocol of animal experiments was approved by the animal ethical committee of the City University of Hong Kong. The mouse brain was covered with ultrasound gel and placed in a water tank to ensure an optimal environment for the experiments. The pulsed laser beam was delivered to the brain for PAI and the pulse energy is ∼4.5 µJ, resulting in a fluence of around 16-mJ/cm2 on the tissue surface (assuming the optical focus had a depth of 500 µm below the tissue surface). This is below the 20-mJ/cm2 ANSI laser safety limit [44]. A PA image was collected by the high-frequency ultrasound transducer. A PA maximum amplitude projection (MAP) image of the brain vessels of the region outlined by the black dashed box in Fig. 4(a, i) is shown in Fig. 4(a,ii). For USI, ultrasound waves were emitted from the low-frequency transducer to the mouse brain and the backscattered US waves were detected by the high-frequency transducer. Figure 4(b) displays an US image of the skull, which was taken with a step size of 2.5 µm and 2000 × 2000 A-lines in the lateral direction. The US and PA images are automatically co-registered. Figure 4(c) shows an overlaid PA/US image. Two representative PA/US B-scan images along the dashed line in Fig. 4(c) are shown in Fig. 4(d). The US B-scan image shows hypoechoic areas of the interfrontal suture and the coronal suture, and the PA B-scan image shows the co-registered superior sagittal sinus and interior cerebral vein. It demonstrates complementary contrasts in the PAI and USI.
Fig. 4. (a, i) photograph of a mouse brain. (a, ii) PA image of the brain vessels of the region outlined by the black dashed box in (a, i); (b) US image of the skull; (c) overlaid PA/US image; (d) PA/US B-scan images along the dashed line in (c); scale bar: 500 µm.
3.2 PAI/harmonic USI of the mouse eye
Using the designed probe, we can implement simultaneous PAI and tissue harmonic USI. For the laser pulse energy in PAI, the maximum permissible exposure (MPE) on ocular exposure for a single laser pulse is MPESP = 5.0× CE × 10−7 = 8.34 × 10−4 J/cm2, where CE is a correction factor calculated as 1668 according to the NA = 0.25 of our system. The repetitive pulse limit was then calculated as MPERP = ntotal−0.25 · MPESP = 2.64 × 10−5 J/cm2, where ntotal was the total number of pulses during imaging (1000 × 1000 pixels in the x and y direction respectively). For a typical human pupil of 7 mm, the maximum permissible single laser pulse energy was calculated as MPERP × pupil area = 10.2 µJ, which is higher than the pulse energy used in our experiment [6,44,45]. Figure 5(a) shows a PA MAP image of the eye of the region outlined by the black dashed box in (a, i) in the iris. In harmonic USI, the low-frequency transducer emits a sequence of two cycles of 25-MHz sine waves, which are then detected by the designed probe. Figure 5(b) displays the fundamental US image of the eye after low-pass filtering, while Fig. 5(c) shows the tissue harmonic US image achieved through high-pass filtering. Compared to the fundamental image, the tissue harmonic image offers higher resolution and reveals finer structures of the iris [46–49]. The overlaid PA/US image is shown in Fig. 5(d). Figure 5(e) shows overlaid PA/US B-scan images along the white dashed line in Fig. 5(d). Distinct structures can be observed in Fig. 5(e), such as the pupil in the PA image, and the cornea, iris, and lens in the US image (Visualization 1). Figure 5(f) shows the PA, fundamental and harmonic US profiles along the yellow dashed curve in Fig. 5(d). Because of the complementary contrasts, the local peaks of the harmonic US curve represent the iris structure and are approximately aligned with the local valleys of the PA curve. The local peaks of the PA curve represent the blood vessels and are nearly aligned with the valleys of the harmonic US curve. Due to poor resolution, the fundamental US curve does not show obvious features complementary with the PA result. Thus, compared to the fundamental USI, the harmonic USI offers a higher resolution reference for PAI.
Fig. 5. (a, i) photograph of the mouse eye. (a, ii) PA MAP image of the eye of the region outlined by the black dashed box in (a, i); (b) a fundamental and (c) a harmonic US image of the mouse eye; (d) overlaid PA/US image; (e) PA/US B-scan images along the white dashed line in d; (f) PA, fundamental US, and harmonic US profiles along the yellow dashed curve in d. The scale bars are 200 µm.
3.3 PAI/harmonic USI of the tumor
We conducted imaging of a 4T1 tumor using simultaneous dual-modal PA/US microscopy. The mouse used was a BALB/c 4-5 weeks old with the tumor subcutaneously implanted in its leg. PA and US excitations were similar to those used in eye imaging. During each raster scan, 1000 × 1000 PA and US A-lines were acquired, with the step size in the lateral direction measuring 7.5 µm. Figure 6(a,ii) shows a PA MAP image of the tumor of the region outlined by the black dashed box in Fig. 6(a, i) on the 7th day after implantation. Blood vessels in and around the tumor are revealed. Figure 6(b) and 6(c) illustrate the fundamental and harmonic US images of the tumor, with the latter producing images with finer structures. Figure 6(d) displays an overlaid PA/US image. Figure 6(e, i) plots a representative overlaid PA/US B-scan image, showing the tumor to be hypoechoic in the harmonic USI. An approximated tumor boundary was thus delineated. To determine the tumor thickness, we plot an overlaid US/PA side view projection image in Fig. 6(e, ii). Measuring the tumor thickness from a harmonic US image yielded a result of 1.7mm. The advantage of using a low-frequency ultrasound for this assessment is that it can penetrate deeply and clearly outline the lower boundary of the tumor. The advantage of using a low-frequency ultrasound for this assessment is that it can penetrate deeply and clearly outline the lower boundary of the tumor. Moreover, the potential of using a simultaneous dual-modal PAI/USI system for tumor assessment is promising, as it can simultaneously image the vascular structures and 3D tumor boundaries.
Fig. 6. (a, i) photograph of the mouse leg. (a, ii) PA MAP image of the tumor of the region outlined by the black dashed box in (a, i); (b) fundamental US image, (c) harmonic US image of the tumor, (d) overlaid PA/US image, (e) overlaid PA/US B-scan image (i) along the white dashed line in (d), and overlaid side view projection of PA/US image (ii) in the white dashed region in (d). The scale bars are 1 mm.
4. Conclusion
We present simultaneous dual-modal PA/US microscopy to reveal co-registered optical and acoustic properties in vivo. This is achieved by combining two ultrasound transducers in an acoustic combiner for high-frequency PAI and harmonic USI. The design of the probe achieves deep-penetration and high-resolution USI. To ensure both sufficient ultrasonic excitation and sensitive PA/US detection, we have optimized the transmittance and reflectance ratio of the AC. We characterize the imaging quality in phantom experiments and demonstrate the dual-modal imaging ability in the mouse brain. In the eye imaging, harmonic USI can reveal fine anatomical features, such as the iris and the lens boundaries, which provides a clear reference for the co-registered vascular features. The subcutaneous tumor experiment can image vessels and tumor boundaries in 3D, which demonstrates a potential application of the simultaneous dual-modal PA/US microscopy.
Funding
Research Grants Council of the Hong Kong Special Administrative Region (9042941); National Natural Science Foundation of China (62135006, 81627805, 81930048).
Disclosures
Lidai Wang has a financial interest in PATech Limited, which, however, did not support this work. All authors declare no competing interests.
Data availability
Data underlying the results presented in this paper are not publicly available at this time but may be obtained from the authors upon reasonable request.
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