A tunable acoustic gradient (TAG) lens has been introduced in the fast axial scanning in optical imaging. However, it still needs additional imaging time for depth scanning. In this study, we split a single laser pulse into three sub-pulses and introduce them into three fibers with different lengths. The sub-pulses out of the fibers were combined thereafter. We then obtained a pulse train with a time interval of 120 ns. By controlling the fire time of the pulse train and the driving signal of the TAG lens, we can receive three focal spots in one A line data acquisition using a single input laser pulse. The depth of focus (DoF) of the system was measured to be 360 μm, which is three times of that of previous systems without the sacrifice of time resolution. A mouse ear and mouse cerebral vasculature were imaged in-vivo to demonstrate the feasibility of the extended DoF of our system.
© 2017 Optical Society of America under the terms of the OSA Open Access Publishing Agreement
Many kinds of three-dimensional optical imaging technologies have played significant roles in biomedical research, such as confocal microscopy , optical coherence tomography . As a promising tool for biomedical research, optical-resolution photoacoustic microscopy (OR-PAM) provides noninvasive visualization of optical absorption with high resolution and high sensitivity [3–7]. OR-PAM utilizes the time-of-flight information carried by the PA signals, which makes it own the ability of volumetric imaging with only two-dimensional (2D) scanning. It has been widely applied in the characterization of the microvascular system [8–10] that is crucial for the biomedical studies including cancer researches  and brain activities .
The conventional OR-PAM employs tightly focused Gaussian beam to achieve high lateral resolution . Since the focused Gaussian beam only has a narrow depth range in focus, OR-PAM suffers a rapidly degrading lateral resolution and signal-to-noise ratio out of its optical focus depth, little details in depth direction can be revealed . The non-diffraction properties of Bessel beam make it own large depth of field compared with Gaussian beam , while the artifacts introduced by the side lobes should be suppressed by non-linear method. By exploiting chromatic aberration and stimulated Raman scattering, the depth of field (DoF) could be improved to be more than 440 μm. However, this system cannot be used for multi-wavelength functional imaging . The mechanical raster scanning can enhance the DoF of PAM system by moving the objective or sample with motor-driven linear platform in depth direction , and it is commonly used to volumetric imaging. But this method limits the volumetric imaging speed with slow mechanical adjustment of the motor-driven linear platform.
In our previous work, we demonstrated that the TAG lens could be used in the OR-PAM for fast axial scanning . However, in our previous research, we still need additional imaging time for depth scanning, which makes this method not favorable in OR-PAM. Therefore, the ultrafast axial scanning of the TAG lens was not fully exploited in the OR-PAM.
In this manuscript, we report a novel method to address the problem. We employs a high-speed TAG lens and optical delay pathways to obtain multi-focus with one input laser pulse, to achieve large DoF. The DoF is estimated by a vertically tilted carbon fiber. A mouse ear and the mouse cerebral vasculature were imaged to demonstrate the feasibility of our method in biomedical imaging.
2.1 System setup
The multifocus photoacoustic microscope referred to as “MF-PAM” hereafter, is illustrated in Fig. 1(a). The system utilizes an Nd: YLF laser (IS8II-E, EdgeWave GmbH), which provides laser pulses with a pulse width of 9 ns and a repetition rate of 1 kHz at the wavelength of 523 nm. First, an iris with a diameter of 0.8 mm is used to reshape the laser beam. Then, the laser beam is expanded by a beam expander consists of planoconvex lenses L2 and L3. A pinhole with a diameter of 50 μm is well placed near the focus of the expander as a spatial filter. Then the laser beam is split into three beams, and coupled into three multimode fibers (core diameter: 50 μm) with a length of 1 m, 26 m and 51 m, respectively. The pulse energy of each split pulses at the front focal plane of the objective is controlled to about 80 nJ by neutral density filters placed at each optical path. Laser beams from the distal end of the three multimode fibers are coupled into a single-mode fiber together. The output of the single-mode fiber is transformed into collimated beam with a diameter of 0.65 mm by a fiber port (PAFAX-4-A, Thorlabs). Then the collimated beam is guided into homemade TAG lens vertically. A beam expander formed by planoconvex lenses L11 (f = 18 and L12 (f = 150 mm) with a magnification factor of 8.3 expands the beam which comes from TAG lens. And the TAG lens is conjugated to the back focal plane of the objective. Finally, the laser beam is tightly focused on the sample by an objective (5 × Mitutoyo, N. A. 0.14) and a home-made prism. The prism is also used as an acoustic lens (N. A. = 0.5) to detect photoacoustic signals by being glued on an ultrasonic transducer (central frequency 50 MHz, bandwidth ~39 MHz at −3 dB, V214-BB-RM, Olympus) . The ultrasonic detection is coaxial with the optical excitation. The detected photoacoustic signals are amplified by an amplifier (AU-1291, MITEQ) and digitized by a data acquisition card (ATS9350, Alazartech). The water tank is used to couple the photoacoustic signals. The sample is moved by a 3D scanning stage that is assembled by a 2D linear stage (ANT95-XY, Aerotech) and a lifting stage (M-Z01.5G0, PhysikInstrumente). In the B mode scanning, the linear stage moves at a constant speed, generating a position synchronization output (PSO) pulse every pixel pitch. A function generator (DS345, Stanford Research Systems) generates a sinusoidal signal with a frequency of 707 kHz to drive the TAG lens at an eigenmode. The Sync Output of the function generator, which provides the synchronizing TTL square wave, is connected to the clock input of a D-type flip-flop (SN74AUC1G74, Texas Instruments). PSO signal of the 2D scanner is sent to the data input of the flip-flop. The output of the D-type flip-flop is used as the trigger signal of the laser. Thus, the laser, scanner, and TAG lens were synchronized.
The home-made TAG lens used in our system consists of a cylindrical piezoelectric shell (PZT-8, Boston Piezo Optics) filled with a transparent silicone oil (100 cS, Sigma-Aldrich), with a refractive index of 1.403 and a speed of sound of 1000 m s−1. The length of cylindrical piezoelectric shell is 20 mm with inner diameter of 16 mm and outer diameter of 20 mm. The focusing power of the TAG lens will exhibit a sinusoidal oscillation at the frequency of the driving signal [18, 20]. Since the trigger signal of the laser is synchronized with TAG driving signal, different delay time of laser pulse relative to the TAG driving signal allows us to synchronize laser pulse with the desired vibration state of the lens. Figure 1(b) is the sequential chart of laser pulses and sinusoidal driving signal of TAG lens in one A-line data acquisition we measured. There are ~120 ns between laser pulse from MMF with a length of 1 m and that with a length of 26 m. Therefore, the laser pulses out of three multimode fibers synchronize with three vibration states of the TAG lens, respectively. And we finally achieved multifocus in each A line data acquisition.
It is worth noting that after 120 ns, when the second laser pulse is delivered at the depth ~120 μm away, the first ultrasonic wave have already propagated for about 180 μm, taking the ultrasound velocity as 1500 m/s in tissue and water. The bandwidth of our transducer is 39 MHz, which provides an axial resolution of 15 μm. Therefore, the impulse responses can be clearly separated. Figures 1 (c) and 1(d) are PA signals of the carbon fiber which is placed at the center focus of the MF-PAM in one A line when TAG lens is on and off, respectively. As we can see, when TAG lens is on, only one laser pulse is focused on carbon fiber, the other two laser pulses introduce very little response (indicated by blue and red arrows in Fig. 1(c). When TAG lens is off, three laser pulses focus on the carbon fiber as shown in Fig. 1(d). In the following experiments, when TAG lens is off, two optical paths were blocked by white paper.
2.2 System performance
A vertically tilted carbon fiber with a diameter of 6 μm was used to demonstrate the system performance of DoF. We used a glass slide, and several fragments of cover glass to make the carbon fiber vertically tilted. Each cover glass was 170 μm thick. And several pieces of glass were glued onto a glass slide to form a raised platform with a height of about 500 μm. As shown in Fig. 2, one end of a straightened carbon fiber was fixed on the glass slide, while the other end was fixed to the surface of the platform. 2D raster scan with a fixed height was performed on the vertically tilted carbon fiber to acquire the depth-dependent PA signal distribution of the carbon fiber. And the carbon fiber was kept linear during the scanning. We imaged the phantom with TAG lens on and off, respectively.
2.3 In vivo imaging
To demonstrate in vivo imaging capability of our system, the left ear of an 8-week-old male Balb/c mouse (Animal Biosafety Level 3 Laboratory, Wuhan, China) was imaged in vivo. Before imaging, the mouse was anesthetized by intraperitoneal injection of chloral hydrate (0.2 g/kg) and urethane (1 g/kg). To show the extended DoF of MF- PAM system, we made the mouse ear tilted in the Y direction. All procedures were carried out in accordance with the Institutional Animal Care and Use Committee of Hubei Province.
To fully demonstrate the performance of our multi-focus system, we performed in vivo cerebral vascular imaging on a mouse. A 20 g female Balb/c mouse (Animal Biosafety Level 3 Laboratory, Wuhan, China) was used for imaging. Before imaging, the mouse was anesthetized by intraperitoneal injection of chloral hydrate (0.2 g/kg) and urethane (1 g/kg). And then the head of the mouse was fixed on a brain stereotactic apparatus for further operation. Most of the skull was removed using a dental drill, forming a cranial opening window of 5 × 5 mm2. After a layer of artificial cerebrospinal fluid had been implemented on the exposed dura mater, ultrasonic gel was applied for acoustic coupling. All procedures were carried out in accordance with the Institutional Animal Care and Use Committee of Hubei Province.
3.1 System performance
Figures 3(a) and 3(b) are the max amplitude projection (MAP) images when TAG lens is on (5 Vp-p sinusoidal signal) and off, respectively. We can find out that more details in depth can be revealed when TAG lens is on. We carried out the YZ and XZ projections, Figs. 3(c) and 3(d) are YZ projections of the vertically tilted carbon fiber when TAG lens is on and off, respectively. Figures 3(e) - 3(h) show cross-sectional B images indicated by f1, f2, f3, f2’ in Figs. 3(a) and 3(b), respectively. The vertically tilted carbon fiber from bottom to top is in extended DoF of MF-PAM, while, only one laser pulse is in focus, the other two laser pulses (indicated by white arrow in Fig. 3(f)) are focused to up and bottom, respectively. By utilizing Gaussian fit to each line in Figs. 3(a) and 3(b), we acquired the FWHM of the profile of carbon fiber. Figures 3(i) and 3(j) show the variations of lateral resolution along the z direction when TAG lens is on and off, respectively. Since the lateral resolution at the focal plane is the best, three focal planes exist (f1, f2 and f3) in Fig. 3(i), while only one focus (f2’) exists in Fig. 3(j). The position in depth of each focus is indicated by the red dashed lines in Fig. 3(i) and 3(j). The depth difference between f1 and f2 is about 129 μm, and the depth difference is about 117 μm between f2 and f3. The distance between f1 and f2 is not the same with the distance between f2 and f3, which is introduced by the time delay of the D-type flip-flop. The time delay of the D-type flip-flop was measured to be about 10 ns, which can theoretically introduce a focus shift of about 10 μm. We defined the DoF as the depth range in which the FWHM of the carbon fiber broadened to twice the narrowest one. The lateral resolution was estimated to vary from 4 to 8 μm over a depth range of 360 μm. Therefore, we estimate that the DoF is about 360 μm when TAG lens is on. The DoF of the acoustic lens is measured to be about 450 μm, the decrease of detection efficiency at the margin of the DoF compared with that at the center in MF-PAM is 33%. Therefore, the ultrasound detection sensitivity within the depth range of 360 μm can be regarded as uniform or similar.
3.2 In vivo imaging of the mouse ear
Figure 4 shows in vivo PA images of the mouse ear. Figures 4(a) and 4(b) are depth-coding MAP images obtained when TAG lens is on and off, respectively. Figures 4(c) and 4(d) are close-up images of the areas indicated by the yellow dashed rectangles in Figs. 4(a) and 4(b), respectively; Figs. 4(e) and 4(f) are close-up MAP images of the areas indicated by the white dashed rectangles in Figs. 4(a) and 4(b), respectively. It can be seen that the vessels in Fig. 4(a) can be clearly resolved when the TAG lens is on. However, only part of the vessels are in focus in Fig. 4(b) when the TAG lens is off. The detailed comparison can be found in Fig. 4(c-f). To quantitatively demonstrate the advantage of our system, a vessel (indicated by the white dashed line in Figs. 4(a) and 4(b)) was chosen for width variations analysis, we chose a point every 9 μm in depth along the white dashed curve, and utilized Gaussian fit for each point. The corresponding width of the vessel was defined as the FWHM of the fitted curve. As shown in Fig. 4(g), the width from the defocused part of vessel broadened rapidly when TAG lens is off. Figure 4(h) which is corresponds to Fig. 4(g) shows the SNR variations of the vessel we chose. The SNR of the defocused part of vessel when TAG lens is off is quite low (8 −15 dB), and the SNR varies sharply (8 −22 dB), whereas the SNR retained relatively high and stable (21 −25 dB) when TAG lens is on. Therefore, our system has special value in quantitative analysis of the network of micro-vessels, with the homogenous spatial resolution and SNR.
3.3 In vivo imaging of the mouse cerebral vasculature
As shown in Fig. 5, about 3 mm × 3 mm area of the open-skull mouse cerebral vasculature was imaged. Figure 5(a) and Fig. 5(b) are depth-coding MAP images obtained when TAG lens is on and off, respectively. Benefited from the much larger DoF of the MF-PAM, more microvessels at different depths can be distinguished. Figures 5(c) and 5(e) are close-up images of the areas indicated by the yellow dashed rectangles in (a) and (b), respectively; Figs. 5(d) and 5(f) are close-up images of the areas indicated by the white dashed rectangles in Fig. 5(a) and Fig. 5(b), respectively. The vessels that are indicated by white arrows can be visualized by MF-PAM, but fuzzy or missing when TAG lens is off.
4. Discussion and conclusion
In summary, by using a TAG lens and optical delay pathways, we developed a multifocus photoacoustic microscope. The TAG lens is used to realize ultrafast axial scanning. The optical delay pathways output three laser beams with different delay time from single laser pulse, which gives us three focuses in each A line data acquisition. Compared with the system when TAG lens off, the DoF of MF-PAM has been improved to 360 μm, with a lateral resolution no worse than 8 μm. Any OR-PAM can be upgraded by following our design to extend the DoF without sacrifice of time resolution.
The TAG lens works for a wide frequency range, except that at some frequencies the PZT tube inside the lens resonates, which contributes to significant improvement in focusing capability. These resonance frequencies are dependent on the physical property and dimensions of the PZT tube, the lower limit can be down to several kHz and the upper limit can be several to tens of MHz according to the authors’ knowledge. Although we can only adjust the frequency of TAG lens to fit the repetition rate of the laser. It still requires several pulses from laser to get an A line with large DoF. This will be even more time consuming, since the laser will not always work in its maximum frequency. Our method split one pulse from laser to several pulses. This can be regarded as a burst of much higher frequency pulses to the laser. Therefore, adjusting the driving frequency to fit the laser is not the best choice.
It is worth noting that, The DoF of the acoustic lens we measured is about 450 μm, and the DoF of single focus system is about 120 μm (Fig. 2(b)). To maintain uniform or similar ultrasound detection sensitivity, 3 or 4 pulses in one A line can be adopted. Larger DoF can be achieved by using more pulses. But the DoF of the ultrasound transduce must be taken into consideration. Because the response of the ultrasound transducer will degraded out of its own DoF. As shown in Fig. 3(i), the delay of 120 ns can cause a focus shift of about 120 μm with current focusing power of our TAG lens, which can maintain relative uniform lateral resolution along the depth direction in our system and keep the ultrasound signal induced by each pulses far from each other. Longer delay line can cause a larger focus shift, but this will lead to non-uniform lateral resolution along the depth direction and non-uniform ultrasound detection sensitivity. Nevertheless, the chosen of number of the pulses and the length of the delay line is dependent on optical focus, the DoF of ultrasound transducer and the focusing power of TAG lens. One can also change these parameters in different applications to optimize the performance of the system. Normally, the original pulse energy of the laser is much higher than we needed in our imaging. Therefore, we can get enough pulse energy when we need more focuses in each A line.
The volumetric imaging speed of our current multi-focus system is three times of that of single-focus system with the same DoF, which can significantly improve the volumetric imaging speed with uniform lateral resolution along the depth direction. This method can be easily adopted in many point-scanning PAM to accelerate the 3D image acquisition and will be helpful in the observation of physiological and pathological processes in 3D.
973 project of China (Grant No. 2015CB755602); Science Fund for Creative Research Group of China (Grant No. 61421064); National Natural Science Foundation of China (NSFC) (Grants No. 91442201); Director Fund of WNLO.
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