IEEE SENSORS JOURNAL, VOL. 19, NO. 3, FEBRUARY 1, 2019 909

Miniature Probe for Forward-View Wide-FieldOptical-Resolution Photoacoustic Endoscopy

Guangyao Li, Zhendong Guo, and Sung-Liang Chen

Abstract— Photoacoustic endoscopy (PAE), an ideal approachto enable internal organ and intravascular imaging, has toadopt a tailored scanner or scanning mechanism for 2-D or3-D imaging. Forward-view endoscopy has found medical appli-cations in gastrointestinal endoscopy and cystoscopy. Althoughrecent works have demonstrated rotational scanning by usinga micromotor or a rotational joint, they are not applicable toforward-view PAE. A micro-electro-mechanical systems mirrorscanner has been used in forward-view PAE but suffers fromthe large probe diameter. Optical fiber bundles for either pho-toacoustic excitation or acoustic detection have been utilized forrealizing forward-view PAE, yet the approaches have drawbacksin low resolution and a tradeoff between wide field of view (FOV)and probe miniaturization. Hence, the implementation of aminiature PAE probe capable of forward-view imaging with wideFOV and high resolution has remained a challenge. Herein,we present a miniature PAE probe (2.4 mm in diameter)capable of forward-view imaging with wide FOV (>3.5 mm indiameter) and high resolution (7.7–10.4 µm) through the useof an imaging fiber bundle, a gradient-index (GRIN) objectivelens, and a fiber-tip Fabry–Perot ultrasound sensor with noscanner integrated into the probe. By incorporation of the GRINlens, the FOV much larger than the probe diameter and highresolution can thus be achieved. The proposed approach isrelatively simple and inexpensive. It may benefit applications ingastrointestinal endoscopy and cystoscopy by the forward-viewcapability with wide FOV and high resolution.

Index Terms— Endoscopes, fiber optics, optical fiber sensors,photoacoustic imaging.

I. INTRODUCTION

ENDOSCOPY has been widely used for industrial andmedical applications because of its ability to accessotherwise invisible sites. Clinically, the endoscope has beenan indispensable tool for visualization of internal structures,such as gastrointestinal tract imaging [1] and intravascularimaging [2]. Photoacoustic imaging is an emerging hybridimaging modality for visualizing optical absorption in biolog-ical tissue [3]. As a result of the significance of endoscopy,many attempts to develop photoacoustic endoscopy (PAE)have been made over the past decade [4]–[22]. PAE is ableto provide high-resolution three-dimensional (3D) images of

Manuscript received August 27, 2018; revised October 21, 2018; acceptedOctober 27, 2018. Date of publication October 30, 2018; date of currentversion January 11, 2019. This work was supported by the National NaturalScience Foundation of China under Grant 61775134. The associate editorcoordinating the review of this paper and approving it for publication wasProf. Agostino Iadicicco. (Corresponding author: Sung-Liang Chen).

The authors are with the University of Michigan-Shanghai Jiao TongUniversity Joint Institute, Shanghai Jiao Tong University, Shanghai200240, China (e-mail: lgy_sjt@sjtu.edu.cn; gzd1247891940@sjtu.edu.cn;sungliang.chen@sjtu.edu.cn).

Digital Object Identifier 10.1109/JSEN.2018.2878801

biological tissue by its endogenous or exogenous absorptioncontrasts [9], [10], [16]. Although the anatomical and func-tional imaging makes PAE valuable in medical diagnosis,unlike bulky photoacoustic imaging, the scanning mechanismsof PAE require special design and implementation, thus tech-nically challenging.

PAE probes basically consist of parts for light illumina-tion and acoustic detection. The photoacoustic A-line signalprovides a one-dimensional depth-resolved image. To enabletwo-dimensional (2D) and 3D images, scanning parts needto be further implemented in the probe, which is essential tofacilitate PAE imaging in most practical cases and clinicalsettings. A number of works have demonstrated scanningsolutions. The use of a micromotor to rotate a mirror hasbeen reported [4], [9], [10], [16]. Both the light illuminationand acoustic detection are steered by the mirror. However,the method requires skillful and sophisticated assembly andsuffers from the finite size of the micromotor, limiting thespread and the overall diameter of the probe. A rotationaljoint has been used in PAE [13], [14]. Since the scanningelement is implemented at the proximal end of the probe, itssize can be made down to 1.1 mm in diameter [13]. However,the scheme involves relatively complicated implementation ofthe proximal actuation unit and flexible catheter section [14].Moreover, the above two methods enable only side-view rota-tional scanning and could not allow forward-view scanning.

Forward-view endoscopy has been demonstrated invarious biomedical imaging modalities including whitelight [23], ultrasound [24], [25], fluorescence [23], [26], andoptical coherence tomography [26], [27], as well as med-ical applications in gastrointestinal endoscopy for tumorimaging [23]–[25] and cystoscopy [27]. Forward-view PAEhas also been investigated [5], [8], [11], [17]–[20]. A micro-electro-mechanical systems (MEMS) mirror scanner offersone possible route to forward-view scanning in PAE and hasbeen used in photoacoustic imaging probes [5], [8]. However,the large probe diameter of 11.5 mm highly restricts itsendoscopic applications. An optical fiber bundle for pho-toacoustic excitation has been used to obtain PAE imagesin the forward direction [11], [18], [20], yet the methodsuffers from narrow field of view (FOV) of ∼2 mm [20].On the other hand, an optical fiber bundle used for acousticdetection by synthesizing a 2D array of ultrasound sensorshas also been demonstrated for forward-view PAE [17], [19].However, the size of the PAE probe (if housed) is largerthan 5 mm [19], which hinders its insertion into the instru-ment channel (typically ∼3.7 mm in diameter) of a standard

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video endoscope [21], [28]. Besides, FOV of both the twomethods (using optical fiber bundles for either photoacousticexcitation or acoustic detection [11], [17]–[20]) is limitedby the PAE probe diameter (i.e., FOV < Probe diame-ter). Further, the two methods have relatively low resolutionof 50 µm–200 µm [11], [19]. Therefore, implementation ofa miniature PAE probe (3.5 mm in diameter), which is larger than the probediameter; (iv) High resolution (7.7–10.4 µm); (v) Relativelysimple assembly.

The probe consists of an imaging fiber bundle and agradient-index (GRIN) objective lens for photoacoustic excita-tion and a fiber-tip FP ultrasound sensor for acoustic detection.The novelty of this design for photoacoustic excitation andacoustic detection is illustrated as follows. For photoacousticexcitation: wide-angle excitation with a focused laser beamis realized by virtue of incorporation of the GRIN lens [36]instead of using an optical fiber bundle alone [11], [18], [20],and the components are in miniature size (Imaging fiberbundle: 950 µm in diameter; GRIN lens: 1 mm in diameter).For acoustic detection: as mentioned above, there are severaladvantages to use the fiber FP sensor for PAE. Here, wide-angle detection is enabled by the fiber FP sensor [22], whosediameter is ∼250 µm. Finally, scanning of laser beams canbe implemented at the proximal end of the probe. That is,no scanner integrated into the distal end of the probe is needed.Hence, the above-mentioned five advantages can be enjoyed.Specifically, (i) The overall diameter of the PAE probe canbe highly compact because of the miniature components usedand no scanner needed. (ii) The design allows forward-viewPAE imaging. (iii) Wide FOV is enabled by the synergy ofwide-angle photoacoustic excitation and wide-angle acousticdetection. Furthermore, FOV can be much larger than theprobe diameter because of the GRIN lens [36]. (iv) Highresolution is achieved by the focused laser beam. (v) Theassembly is relatively simple because integration of only thethree components is needed (details described later).

Fig. 1. (a) Schematic of the miniature forward-view PAE probe. As illus-trated, FOV can be much larger than the probe diameter and high resolutionis achieved by the incorporation of the GRIN lens. (b) Picture of the PAEprobe.

II. METHODS

A. Probe Design and Fabrication

The schematic diagram and the picture of the forward-viewPAE probe are shown in Fig. 1. The probe consists of animaging fiber bundle and a GRIN lens with a fiber-tip FPultrasound sensor attached. There is no additional scanner.To deliver the excitation laser pulses, the imaging fiber bundle(FIGH-30-850N, Fujikura, Japan) was employed. The wholeouter diameter of this imaging fiber bundle is 950±50 µm,while its image circle, consisting of ∼30,000 light guidingelements, has a diameter of 790±50 µm. The individual coreshave an average diameter of ∼3.0 µm with a core-to-corespacing (from center to center) of 4.5 µm. The minimumbending radius of the imaging fiber bundle is 50 mm. Theattractive characteristic of the imaging fiber bundle is thattens of thousands of cores preserve the spatial relationshipbetween the input (proximal) and the output (distal) sidesof the fiber when used for light delivery. Because of this,no micromotor or moving elements are needed in situ toaccomplish the scanning procedure. Instead, raster scanningis implemented at the proximal end of the fiber, and thusforward-view scanning can be realized at the distal end.

Although the imaging fiber bundle allows the removal ofscanning elements and offers several advantages, such aseasy assembly and miniaturization, the imaging fiber bundlealone is not suitable to be used in endoscopic applicationsdue to narrow FOV (less than the diameter of the imag-ing fiber bundle) and very short working distance (WD).

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To overcome these challenges, a GRIN lens (GT-IFRL-100-005-50-C1, pitch: ∼0.28, GRINTECH, Germany) was intro-duced and mounted on the end face of the imaging fiber bundlefor focusing the emitted light, which enables enlarged FOVand extended WD, as illustrated in Fig. 1(a). The GRIN lensused in this work has a length of 2.46 mm with an outerdiameter of 1 mm. It was fabricated by a non-toxic silver-basedglass material and designed to have magnification of –5.86.The AR coating on both end faces of the GRIN lens helps toreduce reflection loss, 98% wasdeposited. To protect the FP cavity, a parylene-C thin filmof ∼5 µm was coated to cover the whole fiber tip. Thefiber FP sensor has sensitivity of 4.5 mV/kPa over bandwidthof 15 MHz and NEP of ∼0.33 kPa, which were calibratedusing the methods described in our previous work [22].Finally, the SMF with the FP sensor was attached to the glasstube to form the PAE probe, as shown in Fig. 1. The overalldiameter of the PAE probe is 2.4 mm. As mentioned above,in contrast to conventional piezoelectric transducers, the fiberFP sensor offers several advantages such as wide acousticdetection angle, miniature size with size-independent sensi-tivity, and broad bandwidth [6], [21], [22], [32], [34]. Espe-cially, the wide-angle acoustic detection of the fiber FP sensoris highly desired in the proposed PAE probe considering thewide-angle illumination and lack of moving elements that canbe utilized to steer acoustic waves.

B. Imaging System

To conduct imaging experiments using the probe, an imag-ing system was built, as shown in Fig. 2. A diode pumpedsolid-state pulsed laser (FDSS532-Q3, CryLas, Germany) witha wavelength of 532 nm, pulse duration of 1.3 ns, and arepetition rate of 1 kHz was used. The laser beam emitted fromthe laser head was split by a beam splitter with a ratio of 10:90.The 10% laser beam was detected by a photodiode (DET10A,Thorlabs), which was used to trigger photoacoustic signalacquisition. The 90% laser beam was used for photoacousticexcitation. It was first attenuated by neutral-density filters toadjust the laser intensity. The laser beam was further expandedto be 10 mm in diameter and then reflected by a mirror.

Fig. 2. Schematic of the imaging system of the PAE probe. CW, continuouswave; ND, neutral density; PC, personal computer.

A coupling lens (AC080-016-A-ML, Thorlabs) was used tofocus the laser beam to couple it into the cores of the imagingfiber bundle. The laser energy per pulse before the imagingfiber bundle was measured as ∼570 nJ, and that after the GRINlens was ∼210 nJ. Therefore, the coupling efficiency (beforethe imaging fiber bundle and after the GRIN lens) was deter-mined to be ∼37%. The coupling lens is 8 mm in diameterand was mounted on a 3D motorized stage (M-L01.2A1 andM-110.1DG, Physik Instrumente (PI), Germany). The 3Dmotorized stage was employed for two purposes. First,the axial scanning along the z direction is to facilitate thelaser beam coupling into the cores of the imaging fiber bundle.Second, the 2D transverse scanning along the x-y direction isto scan the focused laser spots over the proximal end faceof the imaging fiber bundle during imaging acquisition. Thetransverse scanning speed of the motorized stage was setas 1 mm/s. Considering the laser repetition rate of 1 kHz,the scanning step size of 1 µm can be realized. The PAEprobe was submerged in water during imaging acquisition.

Photoacoustic signals were detected by the fiber FP sensor,which was interrogated by a tunable continuous-wave laser(HP 8168F, Agilent) with wavelengths ranging from 1450 nmto 1590 nm. An optical fiber circulator was used to accessthe input and output ports of the fiber FP sensor, as shownin Fig. 2. The output light from the circulator (port 3) wassplit by a 1 × 2 fiber coupler with a power ratio of 10:90.The 10% power was recorded by a power meter (2832-C,Newport), and the reflection spectrum of the FP cavity wasobtained. The wavelength corresponding to the peak of the firstderivative of the reflection spectrum was set as the optimal biaswavelength, which offers the best sensitivity of the fiber FPsensor. Besides, a feedback control method was used to ensurethe stability of the sensitivity [30]. On the other hand, the 90%power was detected by a photodetector (1811-FC-AC, NewFocus) to record the time-varying optical signal modulation,corresponding to the incoming acoustic wave on the FP cavity.The output signal of the photodetector was recorded by adigitizer (CSE1422, Gage) at a sampling rate of 200 MS/s,and the data was stored in a computer for image formation.

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Fig. 3. Zemax simulation to find out the enlarged FOV, WD, and lightfocusing, which is helpful to the design of the probe. (a) Zemax layout andoptical paths of the three light beams. (b) Simulated FOV and WD as afunction of D1, and the comparison of simulated and experimental WD of3 fabricated PAE probes. (c) Simulated NAf as a function of D1, and thecomparison of simulated and experimental NAf of 3 fabricated PAE probes.

C. Simulation

To find out the enlarged FOV, WD, and light focusingafter the GRIN lens by the PAE probe, ray tracing softwareZEMAX was used for simulation. The quantitative simulationis helpful to the design of the probe. Without loss of generality,three light beams with the same wavelength of 532 nmemitted from the center and edge positions of the image circleof the imaging fiber bundle were considered. As shown inFig. 3(a), the distance between the imaging fiber bundle andthe GRIN lens is denoted as D1, and the beam size right afterexiting the GRIN lens is denoted as D2. In this simulation,a numerical aperture (NA) of 0.39 of the imaging fiber bundlewas assumed, which determines the light emitting angle fromthe imaging fiber bundle. The NA of 0.39 was chosen to bethe same as that of the imaging fiber bundle used in this study.Another NA for light focusing, NAf, is defined as nD2/(2 f ),in which n is the refractive index of the medium (whichis water) where the light-exiting side of the GRIN lens issurrounded, and f is the WD, as shown in Fig. 3(a). Note thatthe medium between the imaging fiber bundle and the GRINlens is air. NAf is associated with the focused spot size and thus

TABLE I

COMPARISON OF SIMULATION AND EXPERIMENTALRESULTS OF FORWARD-VIEW PAE PROBES

lateral resolution. As can be seen in Fig. 3(a), after passingthrough the GRIN lens, the three light beams are focused withthe enlarged FOV and extended WD.

In this design, the distance D1 is one of the key parametersto determine the FOV, WD, and NAf. To quantitatively under-stand the effect of the distance D1 on the FOV, WD, and NAf,we performed this simulation by checking D1 at 0–100 µm,as shown in Figs. 3(b) and 3(c). When the imaging fiberbundle is in perfect contact with the GRIN lens, wide FOV of∼5 mm and long WD of ∼7 mm can be obtained, while NAfis ∼0.064. As D1 increases, FOV and WD gradually decreaseto ∼3 mm and ∼4 mm, respectively, at D1 = 100 µm, whileNAf increases to ∼0.11. That is, depending on the need ofendoscopic imaging applications, lateral resolution can befurther enhanced at the expense of FOV and WD.

To verify the simulation results, we made 3 forward-viewPAE probes with different D1, measured the WD and NAf, andcompared the experimental and simulation results, which isshown in Table 1 and Figs. 3(b) and 3(c). The D1 values of the3 probes were measured to be ∼14, ∼46, and ∼87 µm using amicroscope. The same D1 values were used in the simulationfor comparison. As shown in Table 1 and Figs. 3(b) and 3(c),both WD and NAf are in excellent agreement between simu-lation and experimental results.

III. RESULTS

A. Measurement of Resolution and FOV

We first measured the resolution and FOV of our probe toshow its imaging abilities. The measurement was conductedusing the PAE probe with the D1 of ∼20 µm. In order totest the uniformity of resolution, the laser beam was coupledinto (and emitted from) five different cores at the center andedge positions of the imaging fiber bundle (as illustrated inthe upper-left figure in Fig. 4) and then focused by the GRINlens. And then, a 6 µm carbon fiber was laterally scannedthrough the five focused laser spots. The WD was measuredto be ∼6 mm. Fig. 4 shows the measured results at thefive different positions. By checking their full width at halfmaximum (FWHM), lateral resolution of 7.7–10.4 µm can bedetermined. The nonuniformity of the resolution was within±15%. The resolution at the center achieved the best valueof 7.7 µm, while that at the edge degraded, more or less. Thiscan be explained in part by the astigmatism aberration of theGRIN lens, which is also revealed in the ZEMAX simulationin Fig. 3(a).

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Fig. 4. Measurement of resolutions and their uniformity of the PAE probe.Upper-left figure: Schematic of the chosen five different positions at the centerand edge positions of the imaging fiber bundle for laser coupling. Otherfigures: Measured resolutions corresponding to the five focused laser spotsby coupling the laser into the above-mentioned five different positions of theimaging fiber bundle and then being focused by the GRIN lens. The resultsshow that resolutions are 7.7–10.4 µm.

The axial resolution is mainly determined by the bandwidthof an ultrasound sensor (i.e., the fiber FP sensor in this work).The axial resolution of the fiber FP sensor has been measuredas 68 µm in our previous work [35].

To investigate the FOV of the probe, a phantom with a blackmesh grid pattern printed on a transparent film was imaged.The square mesh pattern has a period of ∼680 µm with alinewidth of ∼250 µm. The scanning of laser beams wasconducted at the proximal end of the imaging fiber bundlewith a scanning region of 0.83 mm × 0.83 mm, which waschosen to cover the whole image circle of the imaging fiberbundle. Considering a random distribution of individual fibersin the imaging fiber bundle, scanning step size of 1 µm wasused to ensure that the laser beam was successfully coupledto most of the individual fiber cores. The photoacoustic imageis shown in Fig. 5. As can be seen, the FOV is widerthan 3.5 mm.

This measured FOV is smaller than the simulation result.This discrepancy may be explained as follows: First, the sizesof the imaging fiber bundle and the GRIN lens are comparable,790 µm vs. 1 mm. Light emitted from the edges of the imagingfiber bundle cannot be entirely collected and focused by theGRIN lens (i.e., light leakage), leading to lower excitation laserenergy at the edges in the raster scanning region than the centerand thus narrower FOV. Second, the non-flat angular responseof the fiber FP sensor [22] also contributes to the differencebetween the simulation and the experimental results.

Fig. 5. Measurement of FOV of the PAE probe. Photoacoustic imaging ofthe square mesh pattern was acquired by the PAE probe. The image showsFOV >3.5 mm.

B. Imaging Demonstration

Phantom imaging was performed to show the performanceof the probe. In order to show that the probe is suitablefor endoscopy settings and forward-view imaging, we imageda phantom consisting of five hairs placed at the one sideof a plastic tube (inner diameter: 5 mm) to mimic internaltissues. The PAE probe was inserted from the other sideof the tube. Fig. 6(a) shows the picture of the PAE probeused in endoscopy settings. To enhance the signal-to-noiseratio (SNR), signal averaging of 16 photoacoustic A-lines wasapplied, and then a matched filter of 5–20 MHz was used.As shown in Fig. 6(b), the hairs were successfully imaged overwide FOV. That is, the capability of forward-view PAE andwide FOV of the probe is demonstrated. Another phantom with‘SJTU’ letters printed on a transparent film was also imaged.The font was “Times New Roman.” As shown in Fig. 6(c),the small features of the letters can be clearly visualized,showing the high-resolution capability of the probe.

IV. DISCUSSION

In this work, we demonstrated a miniature PAE probecapable of forward-view endoscopic imaging with wideFOV and high resolution, which is promising for med-ical endoscopy. Table 2 shows the comparison of differ-ent scanning mechanisms in existing PAE probes and ourPAE probe. As mentioned above, compared with the rota-tional scanning methods using a micromotor or a rotationaljoint [4], [9], [10], [13], [14], [16], the proposed approachhas the advantage of relatively simple assembly and enablesforward-view endoscopic imaging, which can be useful forsome applications such as gastrointestinal endoscopy andcystoscopy.

Forward-view PAE probes have been studied [5], [8], [11],[17]–[20]. However, the PAE probe using a MEMS scannersuffers from the large probe diameter of 11.5 mm [5], [8].The methods using a fiber bundle for either photoacoustic

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Fig. 6. Demonstration of forward-view PAE imaging. (a) Picture of the PAEprobe used in endoscopy settings. (b) PAE imaging of five hairs acquiredby the PAE probe. The wide FOV is indicated by the yellow dashed circle.(c) PAE imaging of ‘SJTU’ letters acquired by the PAE probe.

excitation or acoustic detection have the issues in nar-row FOV (smaller than the probe diameter) and relativelylow resolution [11], [17]–[20]. By contrast, our PAE probeachieved a high degree of miniaturization (2.4 mm in diameter)and thus, can readily fit in the working channel of a white-lightendoscope. By the incorporation of the GRIN lens, wideFOV of >3.5 mm was achieved, and unlike using an opticalfiber bundle alone [11], [17]–[20], FOV was larger than theprobe diameter. The proposed PAE probe achieved high lateralresolution similar to the best one reported among all otherexisting PAE probes with scanning capability (9.2 µm [16] vs.7.7–10.4 µm for this work), as shown in Table 2. Note thatthe lateral resolution can be further enhanced at the expenseof FOV and WD, as demonstrated in the simulation.

Besides using the imaging fiber bundle, a GRIN relay lenscan also be used to realize forward-view wide-field endoscopicimaging, as demonstrated in [36]. However, the GRIN relaylens is rigid and thus cannot be bent. By contrast, the imagingfiber bundle is flexible (with the minimum bending radiusof 50 mm) and may thus facilitate clinical applications ingastrointestinal endoscopy and cystoscopy. In this regard,the phantom imaging demonstrated in this work serves as apreliminary study, and further investigation to test the bendingability of the proposed PAE probe is required.

As mentioned previously, the lateral resolution was mea-sured using the PAE probe with the D1 of ∼20 µm.Based on the results in Fig. 3(b), when the D1 is ∼20 µm,

TABLE II

COMPARISON OF SCANNING MECHANISMS IN PAE

FOV is ∼ 4.3 mm. Given that the image circle of the imagingfiber bundle is 790 µm, the magnification from the exit ofthe imaging fiber bundle to the focal plane (after the GRINlens) can be estimated to be ∼5.44 [= 4.3/0.79]. Note thatthe imaging fiber bundle tightly confines the light due tothe high index contrast [37]. Thus, the mode field diameter(amplitude reduced to 1/e) can be estimated to be the sameas the core diameter, which is ∼3.0 µm. Then, the focusedbeam diameter (amplitude reduced to 1/e) can be calculatedto be 16.3 [=3 × 5.44] µm. Considering that the lateralresolution is determined by FWHM (amplitude reduced tohalf), the measured lateral resolution of 7.7–10.4 µm slightlysmaller than the calculated focused beam diameter of 16.3 µmis considered to be reasonable.

As mentioned, when the input laser energy was ∼570 nJ,the output laser energy (after the GRIN lens) of ∼210 nJcan be used for photoacoustic excitation. No obvious dam-age of the imaging fiber bundle was observed. The energylevel is considered to be high enough for optical-resolutionphotoacoustic microscopy [38]. However, when we furtherincreased the input laser energy, partial damage at the exit ofthe imaging fiber bundle (the side close to the GRIN lens) wasobserved. This may be explained by the property of inter-corecoupling of the imaging fiber bundle [37], [39], which maylead to too high laser energy density at the exit of the imagingfiber bundle and thus the damage. Therefore, the suggestedmaximum input laser energy of our PAE probe is ∼570 nJ.In acquiring Figs. 5 and 6(c), the input laser energy was belowthe suggested maximum input laser energy, and good-qualityimages were obtained. By contrast, the input laser energyused to acquire Fig. 6(b) was above the suggested maximuminput laser energy, resulting in partial damage of the imagingfiber bundle. Thus, the laser was not delivered successfully incertain individual fibers, and some discontinuities of the hairsin Fig. 6(b) were displayed.

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In this work, the wavelength of 532 nm was used. As amatter of fact, visible light can be transmitted by the imagingfiber bundle. Thus, functional information such as oxygensaturation may be acquired by the proposed PAE probe usingmultiple wavelengths for photoacoustic excitation [3].

Further developments are needed to improve the perfor-mance of the probe for clinical applications. The scanningspeed needs to be improved. Currently, the imaging speed ismainly limited by the 1 kHz repetition rate of the pulsed laserand the signal averaging of 16 photoacoustic A-lines due tothe mediocre sensitivity of our fiber FP sensor. The laser witha repetition rate up to 100 kHz is commercially available.Besides, a fiber-optic FP-based optical microresonator witha very low NEP of 5 Pa for ultrasound sensing has beenreported [34]. Thus, high-speed or even real-time imaging bythe proposed PAE probe is technically feasible.

In this work, pure water was used as the coupling mediumof ultrasound. In clinical applications, the target may be underscattering media. In this case, resolution and sensitivity ofPAE imaging will be degraded. The issue could be miti-gated by using the near-infrared wavelength for photoacousticexcitation.

In photoacoustic imaging, the imaging depth is highlydependent on the sensitivity of the imaging system(i.e., the SNR of photoacoustic signals). The imaging depthusing the fiber FP sensor (NEP of ∼0.33 kPa) has beenmeasured as 0.38 mm in biological tissue in our previouswork [35], where the laser pulse energy of ∼510 nJ withfocused spot size (lateral resolution) of 3.7 µm was used forphotoacoustic excitation, and the WD was ∼5.5 mm. In thiswork, the fiber FP sensor (with the same NEP of ∼0.33 kPa)was employed at similar WD of ∼6 mm, while the laser pulseenergy of ∼210 nJ with focused spot size (lateral resolution)of 7.7–10.4 µm was used. Because of lower laser fluence usedin this work, it is expected to have lower SNR and thus lowerimaging depth (i.e., 3.5 mm), (iv) High resolution(7.7–10.4 µm), (v) Relatively simple assembly. These featuresmay be useful for applications in gastrointestinal endoscopyand cystoscopy.

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Guangyao Li received the B.S. degree in opto-electronics information engineering from Xi’anTechnological University, Xi’an, China, and the M.S.degree in optical engineering from Tianjin Univer-sity, Tianjin, China. In 2015, he joined the Universityof Michigan-Shanghai Jiao Tong University JointInstitute, Shanghai Jiao Tong University, Shanghai,China, as a Graduate Student. His research interestsinclude the development of Fabry–Perot sensorsand photoacoustic imaging systems for biomedicalapplications.

Zhendong Guo received the B.S. degree in opto-electronics information engineering and the M.S.degree in optical engineering from Tianjin Uni-versity, China. He is currently pursuing the Ph.D.degree with the University of Michigan-ShanghaiJiao Tong University Joint Institute, Shanghai JiaoTong University, Shanghai, China. His researchinterests include the development of photoacousticimaging systems for biomedical applications.

Sung-Liang Chen received the Ph.D. degreein electrical engineering from the University ofMichigan, Ann Arbor, MI, USA, and post-doctoraltraining at the University of Michigan MedicalSchool, USA. He is currently an Assistant Profes-sor with the University of Michigan-Shanghai JiaoTong University Joint Institute, Shanghai Jiao TongUniversity, Shanghai, China. His research interestsinclude optical resonators for sensing applications,optical imaging systems, and biomedical photoa-coustic imaging. He was a recipient of the Thousand

Talents Plan from the Chinese Recruitment Program of Global Experts foryoung professionals and the Shanghai Pujiang Talent Award.

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