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Abstract
Acoustic resolution photoacoustic microscopy (AR-PAM) has gained much attention in the past two decades due to its high contrast, scalable resolution, and relatively higher imaging depth. Multimode optical fibers (MMF) are extensively used to transfer light to AR-PAM imaging scan-head from the laser source. Typically, peak-power-compensation (PPC) is used to reduce the effect of pulse-to-pulse peak-power variation in generated photoacoustic (PA) signals. In MMF, the output intensity profile fluctuates due to the coherent nature of light and mode exchange caused by variations in the bending of the fibers during scanning. Therefore, using a photodiode (PD) to capture a portion of the total power of pulses as a measure of illuminated light on the sample may not be appropriate for accurate PPC. In this study, we have investigated the accuracy of PPC in fiber-guided and free-space AR-PAM systems. Experiments were conducted in the transparent and highly scattering medium. Based on obtained results for the MMF-based system, to apply PPC to the generated PA signals, tightly focused light confocal with the acoustic focus in a transparent medium must be used. In the clear medium and highly focused illumination, enhancement of about 45% was obtained in the homogeneity of an optically homogeneous sample image. In addition, it is shown that, as an alternative, free-space propagation of the laser pulses results in more accurate PPC in both transparent and highly scattering mediums. In free-space light transmission, enhancement of 25-75% was obtained in the homogeneity of the optically homogeneous sample image.
© 2022 Optica Publishing Group under the terms of the Optica Open Access Publishing Agreement
1. Introduction
Photoacoustic Imaging (PAI) has attracted a lot of attention in the last two decades in biological imaging. This hybrid imaging method has many applications, including structural and hemodynamics imaging of living tissue [1–4]. Utilizing non-ionizing optical radiation to stimulate the sample and capture generated ultrasound waves (also known as PA waves) has turned this method into high contrast, high resolution, and deep-tissue imaging modality. One of the most important types of PAI is Acoustic-Resolution Photoacoustic Microscopy (AR-PAM) [5–8], in which weakly focused light is used to illuminate the sample and a focused ultrasonic transducer (UT) to detect PA waves. By 2D scanning the imaging head or the sample, 3D photoacoustic data is captured. To deliver the light from the laser source to the scanning head, either free-space [9,10] or fiber-guided [11–13] method is used. Although free-space-based light delivery allows minimum loss in the light transmission from the laser source to the scanning head, fiber-based light delivery provides flexibility and a compact system. In a fiber-based system, multimode fibers (MMF) are primarily used to transmit light to the scan-head due to the feasibility and capability of carrying high energy pulses compared to single mode fibers [14–16]. Since most light sources are not very stable, a fast photodiode (PD) is used to sample a small portion of the laser output power and compensate for variations in detected PA signals [11,12]. A few things should be considered when using coherent light and MMF for light delivery. First, due to the high coherency of the laser light and the possibility of optical interference within the MMF, the MMF output will be in the form of a speckle pattern [17–23]. Second, due to the change in MMF bending during scanning, higher-order modes are excited and exchanged in the MMF, and the output profile will not necessarily be Gaussian [17,24–26]. Figure 1 is the photograph of an MMF output under different bending angles (up to ${30^\circ }$).
Fig. 1. The photograph was taken from the output intensity profile of an MMF at a distance of 10 cm from the MMF output surface at different bending conditions (0 to ${30^\circ }$).
In qualitative imaging, the lack of compensation for peak-power variations will have a destructive effect on obtained images in the form of noises similar to salt-pepper noise. Figure 2 is the result of simulation for PA imaging without PPC. Figure 2(a) is the image of the USAF 1951 test target, which is considered as the sample for imaging, (b) is the acquired PA image without PPC, and (c) is obtained by applying gradient filter on (b). Figure 2(b) indicates that the captured image suffers from severe intensity fluctuations in the presence of peak-power variations. This simulation is done in LabVIEW environment.
Fig. 2. Simulation of the effect of peak-power variations on acquired PA images. (a) is the considered sample (USAF 1951 test target) for imaging, (b) the captured photoacoustic image without PPC, and (c) the gradient of (b).
In the case of quantitative imaging, such as when blood oxygen levels need to be measured, the inability to compensate for pulse energy variations in the PA signals will decrease the accuracy of measurements. In this type of imaging, the accuracy of the concentration measurement is crucial in the examination and diagnosis process and cannot be ignored [27,28]. To reduce the effects of pulse energy fluctuations in fiber-based systems, averaging of photoacoustic signals can be used instead of PPC. There is, however, a problem with reduction in the imaging/scanning speeds of the photoacoustic systems. Fast scanning systems have made significant progress in recent years, allowing higher frame rates [29–30]. Live samples can now be examined online via these systems to evaluate PAI potential for clinical use. Averaging signals in these systems will substantially reduce the imaging speed, limiting their use in studying living tissues. Free-space light transmission can reduce some of the problems caused by MMF transmission. The accuracy of PPC in AR-PAM systems has not been discussed in the literature. Therefore, in this study, experiments have been designed and performed to evaluate the effect of PPC and its accuracy on PA images by transmitting light using MMF and free-space configuration.
2. Materials and methods
In this study, an AR-PAM system has been used for the experiments. More details on the imaging system used can be found here [31]. In summary, a Q-switched laser (MSL-AO-532-500µJ, CNI Lasers) with a wavelength of 532 nm, pulse width of 8 ns, and repetition rate of 1-5 kHz was used as the illumination source. According to the manufacturer's documentation, the output intensity of the laser was Gaussian to a reasonable degree. Otherwise, a combination of two lenses and a pinhole would have to be used to clean the beam. A fast photodiode (PD1) (Silicon PIN Detector, 818-BB-22, Newport) and a beam splitter (BS) were placed at the output of the laser to sample the energy of laser pulses needed to compensate for pulse energy fluctuation in generated PA waves as well as the elimination of pulses’ timing jitter noise. To transfer laser pulses to the scan-head of the system, the output of the laser is coupled to multimode fiber (MMF) (M29L05 - Ø600 µm, 0.39 NA, Thorlabs) with the numerical aperture (NA) of 0.39 using a coupling lens (CL). Another fast photodiode (PD2), a pinhole (PH), and a BS were placed at the output of the MMF to investigate pulses in the scan-head. Two lenses were used to collimate and focus light on the sample through an opto-acoustic combiner (OAC) comprised of two right-angle prisms (RAP). Generated photoacoustic (PA) waves after reflection from the boundary between the RAPs were captured by an ultrasonic transducer (UT) (V214-BB-RM, 50 MHz, Olympus) with a central frequency of 50 MHz. Then, these signals were amplified with two cascaded amplifiers (ZFL-500-LN, 0.1-500 MHz, 24 dB, Mini-Circuits) with a total gain of 48 dB, visualized by an oscilloscope, and captured by a data acquisition (DAQ) card. Also, an acoustic lens was attached to the bottom of OAC to acquire PA signals from a focused line. At any location, by illuminating the sample and capturing generated PA signals, 1D depth-resolved signals can be captured. By 2D scanning the imaging head using a mechanical XYZ scanner, 3D PA signals matrix can be constructed. The scanner consists of two parts, the fast and slow axis. The fast axis scans the scan-head, while the slow axis simultaneously scans the sample and the water tank. The maximum speed of the fast axis is 10 mm/s, and the slow axis is 10 µm/s. The scanner acceleration is usually set to 100 mm/s2. A schematic of the fiber-based imaging system is shown in Fig. 3(a). In addition, the capability of using free-space light transmission is embedded in the setup, as shown in Fig. 3(b). An adjustable mirror attached to the fast scanning axis (scan-head) was used to deliver light directly to the sample.
Fig. 3. Block diagram of the AR-PAM imaging system. (a) fiber-based and (b) free-space light transmission. Abbreviations: personal computer (PC), data acquisition card (DAQ), oscilloscope (OSC), fast photodetector (PD1,2), polarizers (P), beam splitter (BS), coupling lens (CL), multimode fiber (MMF), adjustable mirror (AM), objective lens (OL), opto-acoustic combiner (OAC), ultrasonic transducer (UT), water tank (WT), amplifiers (Amp). Fast and slow scanning axes have been shown by blue and red dashed boxes, respectively.
An analysis program prepared in Visual Studio environment written by C-Sharp (C#) language was utilized for data processing. Parallel and multi-thread processing was applied to increase analysis performance. Further, a control circuit consisting of an AVR microcontroller was used to synchronize, trigger, and control the scanner, laser pulse generation, and DAQ. To investigate the fluctuation of illumination patterns due to scanning and mode exchange in MMF, the PH was placed in front of PD2 to capture a small spatial portion of the light pattern. The captured signals were compared with signals captured by PD1. To capture the illumination pattern on the sample, a CCD sensor (5Mpixels, 1.4 µm pixel size) was placed instead of the sample in the acoustic focus. Also, the similarity of illuminated pulses’ power and the corresponding generated PA signals for an optically homogenous sample was studied to have a quantitative analysis. Between the CCD/Sample and the water tank was filled with clear deionized water as the clear medium and tissue phantom with the thickness of 1.5 mm as the highly scattering medium. Tissue phantom was prepared by adding Agarose (CAS-No: 9012-36-6, Merck) (1%) as the matrix and TiO2 (Product 224227, Titanium(IV) oxide, rutile, Sigma Aldrich) particles (0.1%) as the scattering centers to deionized water in consecutive heating, steering, and sonication cycles [32–35].
3. Results and discussion
To show the system's imaging performance, we imaged a mice ear (male, 25 grams) in vivo using the fiber-based configuration. The mice was anesthetized using a cocktail of Ketamine and Xylazine according to the instructions of the Faculty of Cognitive Sciences, Shahid Beheshti University. Depilatory cream was applied to one of the ears to remove hairs. Then, the mice was placed under the water tank, and the ultrasonic gel was applied between the ear and water tank. After imaging, the ear was cleaned, and the mice was allowed to recover. Figure 4 illustrates the photograph and the maximum amplitude projected (MAP) PA image of the mice ear. Here, scan area was 5.4 mm * 3.8 mm, fast scan velocity = 10 mm/s, slow scan velocity = 10 µm/s, and acceleration = 100 mm/${s^2}$.
To evaluate intensity profile fluctuation before and after the MMF, we placed PD1 at the laser output and PD2 at the output of the fiber, as illustrated in Fig. 3(a). PD1 samples a small portion of laser Gaussian output. A PH (diameter of 500 µm) was placed in front of PD2 to sample a small spatial portion of the light which is always assumed to be Gaussian. Therefore, if there was no intensity profile change while transmitting light by the MMF, fluctuation of signals captured by two PDs should follow the same behavior. Figure 5 indicates two signals captured simultaneously using two PDs during scanning two areas of about 5 mm * 1 mm (laser repetition rate = 1kHz, data acquisition rate = 10 data/s, fast scan velocity = 10 mm/s, slow scan velocity = 10 µm/s, acceleration = 100 mm/${s^2}$) while the system stopped scanning for 165 seconds between the two scan areas. The contrast of fluctuations was calculated using formula $C = \; \sigma /\mu = \; \sqrt {\Sigma {{({{x_i} - \; \mu } )}^2}/N\; } /\mu \; $ for three indicated regions by dashed lines to have a quantitative measure of fluctuations. In this formula, σ is the population standard deviation, µ is the mean, xi is each value from the population, and N is the size of the population. Two points can result from Fig. 5: two graphs do not have the same behavior and contrast at different states of the scan, and by starting the scanning, the output signal of the MMF shows different and fast fluctuations relative to the input signal. We conclude that small vibrations of the MMF are caused by vibration of the system and fans in the steady-state, along with vibrations resulting from mechanical scanning changing the output profile of the MMF. Therefore, when quantitative imaging is required, as in the case of measuring the oxygen level of blood, PPC can have negative consequences.
Fig. 5. (a) captured signal by PD1 at the laser's output and (b) captured signal by PD2 at the output of the MMF. The scan-head states (scanning / no scanning) are separated by red dashed lines. The signals captured by PD1,2 are gray, the upper RMS envelope purple, the lower RMS envelope green, and the averaged signal black.
It may be suggested that this profile change may not be effective when the sample is exposed to focused light. To evaluate the illumination profile on the sample, we placed a CCD sensor as the sample under the scan-head and captured the illumination profile directly. Figure 6 shows captured profiles at different degrees of optical focus and in the clear and highly scattering medium. For the clear medium, by moving away from the tightly optical focus, the size of speckle grains caused by interference of coherent light inside MMF increases [18,19]. Due to changes in fiber bending caused by scanning, the illumination pattern showed random translations. Figure 6(a) indicates that the size of speckle patterns in the case of tightly focused light is significantly smaller than the lateral resolution of the system [31]. Therefore, changing the distribution of speckle grains should not affect the average illumination intensity. We expect PPC to have a constructive impact only when tightly focused light is used to illuminate the sample.
Fig. 6. Illumination pattern on the sample captured by replacing the sample with a CCD. (a-d) clear medium, (e-h) highly scattering medium. (a, e) tightly optical focus, (b, f) 5 mm out of focus, (c, g) 10 mm out of focus, (d, h) 15 mm out of focus. The simplified experimental setups for focused and out-of-focus illumination have been shown in (i) and (j), where green and purple show the light and acoustic fields, respectively.
In the highly scattering medium, the size of speckle patterns remained the same as expected from coherent light scattering in highly scattering environments [18,19]. However, there were islands with different average spatial intensities. With the presence of these islands and the fact that their locations were changing during scanning, it seems that PPC may not have a constructive impact in highly scattering mediums such as biological tissues.
A quantitative experiment can be conducted to evaluate the effect of these changes directly on generated photoacoustic signals. Photoacoustic signals are generated according to the equation ${p_0} = \; \eta \Gamma {\mu _a}F$ [8], where p0 is the generated initial photoacoustic wave (pa), $\eta $ the heat conversion efficiency, Γ the dimensionless Grüneisen coefficient, µa the optical absorption coefficient (cm-1), and F the local light fluence(J/cm2). If an optically homogenous sample is chosen for imaging, the equation as mentioned earlier can be written as ${p_0} = \; \alpha F$, where α is a constant. Therefore, if MMF causes no effective intensity distribution change, photoacoustic waves should follow the same fluctuation as signals acquired by PD1. Figure 7 indicates corresponding PA and PD signals obtained by imaging a black tape. The experiments were conducted at different degrees of optical focus in the clear and highly scattering medium. As is evident in the graphs, only in clear medium and when tightly optical focus is used to illuminate the sample, PA and PD signals follow the same behavior for the homogenous sample. These results were also expected from illumination patterns captured by the CCD.
Fig. 7. Investigation of similarity of generated PA signals for an optically homogeneous sample and corresponding signals captured by PD1 at the output/input of the laser/MMF. (a-c) clear and (d-f) highly scattering medium. In (a, d), highly focused illumination was used, (b, e) 5 mm out of focus, and (e, f) 10 mm out of focus.
Table 1 indicates the contrast of the PA image with and without PPC. According to the results, enhancement of 45% due to PPC at the homogeneity of the PA image was obtained in clear medium and highly focused illumination.
Table 1. The contrast of captured PA signals for an optically homogeneous sample with and without PPC in MMF-based propagation. Enhancement of compensation was calculated from changes in the contrast of PA signals after PPC
To apply PPC accurately and maintain the spatial distribution of the laser’s Gaussian profile, we propose using free-space light transmission instead of a fiber-based configuration. The suggested experimental setup is shown in Fig. 3(b). In a clear and highly scattering medium, the PA image of the black tape was captured, and the similarity and contrast of laser peak-powers and corresponding PA signals were compared. Figure 8 illustrates corresponding PA and PD signals. According to the results, PPC in the clear medium greatly enhances compensation accuracy.
Fig. 8. Investigation of similarity of generated PA signals for an optically homogeneous sample and corresponding signals captured by PD1 at the laser's output. (a-c) clear and (d-f) highly scattering medium. In (a, d), highly focused illumination was used, (b, e) 5 mm out of focus, and (e, f) 10 mm out of focus.
Similarly, for highly and weakly focused illumination, PPC is also effective for improving the accuracy of PA signals in a highly scattering medium. Table 2 shows the contrast of the PA image with and without PPC. As a result, PPC enhanced the homogeneity of the PA images by 72-75% in clear and by 25-27% in highly scattering medium.
Table 2. The contrast of captured PA signals for an optically homogeneous sample with and without PPC in free-space propagation. Enhancement of compensation was calculated from changes in the contrast of PA signals after PPC
4. Conclusion
This study investigated the accuracy of peak-power compensation (PPC) in fiber-based and free-space light transmission Acoustic-Resolution Photoacoustic Microscopy (AR-PAM). The accuracy of PPC was studied by comparing fluctuation of intensity profiles at the input and output of the used multimode fiber (MMF). Moreover, during imaging of an optically homogeneous black tape sample, the optical intensity distribution on the sample plane, which a CCD had replaced, and the similarity of behavior between photoacoustic and laser signals was investigated. In MMF-based configuration, it was found that PPC is accurate only when the medium is clear and optical illumination has a high degree of focus on the sample. In this case, an enhancement of 45% in the homogeneity of an optically homogenous sample image was obtained. Therefore, in quantitative experiments such as oxygen level measurements in multi-wavelength systems using fiber-based systems, PPC could be applied only in clear mediums or superficial biological imaging with tightly focused light. We proposed that free-space light transmission could be used to maintain the spatial intensity profile of illumination during scanning. Results for free-space configuration indicate that PPC dramatically enhances the accuracy of compensating PA signals in the clear medium in both focused (confocal) and out-of-focus illuminations. In this configuration, enhancement of 72-75% was obtained. The enhancement was 25-27% in a highly scattering medium, and a focused or weakly focused illumination should be used for deep imaging.
Acknowledgments
The authors would like to thank Prof. Valery Tuchin from Saratov State University, Saratov, Russia, for his helpful comments in the preparation of phantoms while conducting experiments in highly scattering medium as well as Dr. Zeinab Chenari from Shahid Beheshti University, Tehran, Iran, for her comments in the preparation of the results.
Disclosures
The authors declare no conflicts of interest related to this article.
Data availability
The data supporting this study's findings are available from the corresponding author on request.
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