1. Introduction

Bladder cancer is the second most common genitourinary malignancy worldwide [1,2]. Over 430,000 men and women globally and 74,000 in the United States alone are diagnosed with bladder cancer every year, with an incidence four times higher in men than in women. Bladder cancer ranks fourth in the estimated number of new cancer cases in the USA [3,4]. Bladder cancer is also recognized as one of the most notorious cancers in terms of its prognosis because upon transitioning to stage 3, it evolves very aggressively and ultimately kills the patient.

Most bladder cancers are epithelial cell tumors derived from epithelial cells. The most representative types of malignant epithelial tumors are urinary tract cell carcinoma, squamous cell carcinoma, and adenocarcinoma. Although most of these tumors originate from the epithelium, they soon invade the muscular layer of the bladder wall (see Fig. S1 of Supplement 1) and, as with other types of cancer, eventually spread to other organs. At present, the standard protocol for the diagnosis of bladder cancer is a cystoscopy-based visual examination followed by a biopsy. However, if muscular layer invasion is found, an MRI or CT scan of the abdomen and pelvis or urogram and X-ray chest imaging or CT scan are also conducted for disease staging and to identify cancer spread (metastasis).

The aggressiveness of bladder cancer requires that it be detected as early as possible, as with other types of aggressive cancers. However, convincing diagnostic information cannot be obtained using conventional cystoscopy because it relies on simple surface mapping technology achieved by a camera placed at the distal tip of a device (i.e., the instrument is essentially a urological endoscope). This limitation has been addressed in part by a photodynamic concept-based advanced detection method called blue light cystoscopy [5], which provides better discrimination of the cancer when used in conjunction with standard cystoscopy. However, this examination requires the administration of a contrast agent, so it cannot be accomplished as rapidly as a typical cystoscopic procedure. As a subsidiary method, a cytological examination of an excised specimen based on urinary biomarkers, such as nuclear matrix protein 22, has also been suggested [6]; these biomarkers provide better information when used in combination with urine cytology than when used alone. However, current clinical guidelines do not recommend using urinary biomarkers for detection and surveillance, because their detection accuracy is poor for low-grade cancers.

In the present study, we addressed these important clinical issues by engineering a new type of transurethral probe that can provide capillary-level vasculature information based on the high-contrast, label-free imaging capability of photoacoustic (PA) tomography (PAT) and has demonstrated the probe’s in vivo transurethral imaging capability based on a rabbit model. The protocol is based on the PAT technique [7–12]; therefore, it differs from conventional cystoscopy as it is a tomographic imaging device. Moreover, it features a high-resolution vasculature imaging capability because it adopts the optical-resolution (OR) PA endoscopy (PAE) concept of PAE [13–24]. In addition, our implementation of the proposed probe maintains its conventional endoscopic ultrasound (US) (EUS) imaging function [25,26], so it can provide tissue density distribution information over a depth far exceeding that of the OR-PAE mode.

Multiple studies have been conducted over the last decade to apply PAT technology to the diagnosis of urogenital diseases [27–42]. However, most of these studies mainly focused on cancers with a high incidence rate, such as prostate cancer [27–36], or on clinical issues that could be addressed using a relatively large PAE probe [21,37–42]. Our search of the literature did not uncover any reports that considered the physical implementation of a miniaturized PAE probe for imaging the urinary system following direct urethral introduction; only transurethral light delivery methods for prostate imaging have been reported thus far [29,30,34,36]. Moreover, although we limited our search to bladder cancer applications, only a few reports have addressed this issue using a transabdominal imaging approach [40–42]. In the identified literature, successful visualization of an approximate profile of a rat bladder was attained by applying abdominal optical illumination and subsequent PA signal detection. However, this and related approaches must overcome the next hurdle—achieving a clinically relevant penetration depth with PAT—because current PAT data indicate that signal detection at only a 3–4 cm level is possible in a human subject [32,33], whereas image penetration of at least 10 cm would likely be necessary to cover the entire human bladder.

Our working hypothesis upon initiating this study was that PAT images produced based on direct transurethral imaging could be more effective in diagnosis than those provided by transabdominal imaging. We thought that this is currently a true and feasible story because the OR-level image quality of the bladder microvasculature cannot be acquired with current transabdominal imaging approaches and the probe diameter (2.8 mm) that we achieved in this study is much smaller than the typical diameter of clinical cystoscopes, which ranges from 18F to 26F (a cystoscope with an instrument channel typically has a channel diameter of 9.5F). However, we also recognized that the ultimate success of this technique in real clinics would require that two issues be achieved or guaranteed. The first is that a dedicated catheter probe made of a very thin, soft material must be used to in order to avoid causing intolerable pain to the patient during the procedure. The second is that the device must provide diagnostically useful and distinct image information beyond the capacity of the current cystoscopy but without incurring excessive costs. Taking these requirements into account, we implemented the first prototype of our transurethral PAE-EUS probe, featuring a high-resolution vasculature imaging capability based on OR-PAE and a tapered distal structure similar to those of clinical intravascular US probes [43,44] and sheathed by a soft and flexible catheter material. In the present study, the imaging demonstration was limited to single-wavelength imaging. However, we have left options for equipping the implemented device with additional wavelengths to allow it to provide functional, molecular, and lymphovascular image information in the future.

2. Materials and methods

2.1. Transurethral photoacoustic and ultrasonic endoscopic probe

Figure 1(a) depicts the distal section of the prototype transurethral probe used in the present study. It features a GRIN lens (0.5 mm diameter, ∼0.1 pitch; GRINTECH) and dual US transducers (40 MHz, 0.6 mm × 0.5 mm × 0.2 mm, Blatek, USA) as core components for optical focusing (PA) and acoustic signal detection (PA and US). The overall configuration of the associated optical and acoustic elements is based on the probe design concept described in our previous report [24]. However, for a better fit with the intended application of urinary system imaging, we considered the following technical aspects when implementing the new probe:

 

Fig. 1. Transurethral PAE-EUS probe. (a) Schematic of the probe. Note that the Cartesian coordinate system denoted by (x, y, z) is a moving and rotating coordinate system, along with the scanning tip. (b) Photo showing the approximate flexibility of the tapered tip. (c) Structure of the tapered tip. (d) Optical layout. (e) 2D beam profile of the focused laser beam. (f) Simulated approximate 2D synthetic acoustic impulse response function of the dual transducers for the receiving mode (i.e., PA only) along the y–z plane. The overlaid green triangle represents the approximate location of the laser beam.

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We first reduced the size of the rigid distal section to a diameter of 1.6 mm and a length of 8 mm to increase the flexibility of the probe during introduction into the urethra, and we sheathed it with a customized catheter consisting of a 2.5-mm outer diameter (OD) perfluoroalkoxy tubing (for the body section) and a 2.8 mm OD Pellethane tubing with a wall thickness of 150 µm and a durometer value of 90A (for the imaging window section) [Fig. 1(b)]. The two-way acoustic transmittance of the applied Pellethane tubing was relatively lower (∼60%) than that of the Pebax 7033 tubing (durometer value: 70D) utilized for the imaging window of the previous probe [24]. However, we chose the Pellethane material for this transurethral probe because after multiple trials with in vivo animals, we found that the 90A durometer value, which is relatively softer than the previous, exhibits an appropriate hardness for the intended application, given that it does not develop crease marks when bent and still has a decent optical and acoustic transparency required for the imaging window. In Fig. S2 of Supplement 1, we present the acoustic transmittance measurement results of the two tubings. Second, as shown in Fig. 1(c), the catheter was implemented to have a distal structure with a tapered shape. The tapered distal tip acts as a guide during introduction into the urethra while also sealing in the acoustic matching medium that fills the catheter. We also increased the flexibility of the distal section (thereby reducing possible pain to the subject) by implementing a dedicated sealing plastic screw, which was employed as a kind of base frame for the distal tip, and installing a narrow tension spring along its axis. The tension spring provides a restoring force and its surface is coated with a soft and flexible silicone epoxy.

Given our choice of the rabbit urinary system as an animal model and the tubing diameter used, we set the optical working distance (WD) to 1.2 mm, as shown in the optical layout presented in Fig. 1(d)—the WD was defined as the distance from the surface of the stainless steel tubular housing with the 1.6 mm OD to the optical focus created by the GRIN lens. This provided a 6 µm FWHM beam diameter at the focus [Fig. 1(e)], according to a measurement obtained in air using a beam profiler. Here, it should be noted that, as the beam diameter was measured in air, the corresponding optical focus appeared closer (i.e., ∼1.3 mm) than the 1.76 mm distance predicted under the assumption of a water medium for the last immersion medium [see Fig. 1(d)]. In addition to the beam profiling, we also validated the actual transverse resolution of the OR-PAE mode by imaging a 13-µm diameter tungsten wire, and its value was confirmed to be less than 11 µm as the cross-sectional profile of the tungsten wire was recorded in a single angular step (in terms of the FWHM) in a B-scan image. The related imaging results are presented in Fig. S3 of Supplement 1.

Since the WD of the probe is determined by the gap distance (i.e., 1.16 mm) between the fiber tip and the entrance surface of the GRIN lens [Fig. 1(d)], it was critical to accurately control the distance. In this study, we organized the related installation method better than in our previous study [24] and achieved the desired WD by applying a precisely machined, dedicated metal holder for the GRIN lens during the assembly of the scanning tip (see Fig. S4 of Supplement 1). Figure 1(f) shows the relative position of the focused laser beam overlaid onto the 2D synthetic acoustic impulse response function of the dual transducers along the y–z plane. The impulse response function is applied only to the receiving mode (i.e., PA only) and was acquired from a numerical simulation using the k-Wave MATLAB tool box [45].

2.2. Imaging setup

We operated the imaging probe using the PAE-EUS universal driver that we recently reported [24]. As depicted in Fig. 2(a), the endoscope driver incorporates several key modules, including a laser source, US pulser, signal amplifier, US switch, stepping motor, oscillator, and delay generator. These modules are necessary to perform simultaneous PAE and EUS imaging in a co-registered manner in conjunction with an engaged imaging probe in a single-instrument case. Unlike previous PAE-EUS systems [14–17] that employ a commercial pulser receiver, our system achieves the required US pulse-echo imaging using a customized circuit that incorporates an RF switch in place of the conventional US switch [24]. Most specifications of the endoscope driver described in our previous report [24] were used in the current study, except that the previously used RF switch module was replaced with a new model (ZFSW2-33HDR-75+, 3W, Mini-Circuits, USA). We made this revision because the new model can handle greater power when transmitting an electric pulse for US pulsing.

 

Fig. 2. Experimental setup. (a) Peripheral systems. (b) Digital timing diagrams applied to trigger the sub-modules and peripheral systems presented in (a). (c) Photo illustrating the transurethral imaging procedure of a rabbit. (d) Illustration showing the rabbit urinary system in which the transurethral probe is introduced.

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Figure 2(b) shows the timing charts applied to trigger the key modules included in the universal driver and peripheral systems depicted in Fig. 2(a). The oscillator operates at a 16 kHz frequency, which corresponds to the step frequency of the stepping motor that actuates the scanning tip and was determined based on the available pulse energy of the employed laser. For every square pulse generated by the oscillator, the delay generator produces dual pulses with a 23.8 µs delay to trigger the data acquisition (DAQ) card (200 MHz, 12 bits, PCI-5124, National Instruments, USA) installed in a PC to record PA and US A-line signals according to the following procedure: almost simultaneously with the first pulse, the laser (SPOT-10-200-532, Elforlight, UK) installed in the universal driver is triggered and a 532 nm laser pulse is immediately fired. Subsequent PA signals detected by the dual transducers are then amplified by dual amplifiers (ZFL-500LN+, Mini-Circuits) and recorded by the DAQ computer. About 23.5 µs after the first pulse (i.e., near the rising edge of the second pulse that triggers the DAQ), an electric pulse with a 36 V amplitude is generated by the US pulser and sent to the dual transducers to fire an acoustic pulse. Subsequent US A-line signals are then recorded according to the same process explained for PA.

Here, the acoustic pulse firing moment comes ∼0.3 µs before the corresponding DAQ trigger for the US signal record because the generated acoustic pulse takes that amount of time to travel from the transducer to the imaging window. Note also that, unlike in our previous study [24], an increased pulsing voltage of 36 V was applied when operating this new probe. In the previous study [24], a 24 V pulsing voltage was sufficient to record the subsequent US signals without their saturation under the same dynamic range adopted for the PA imaging mode; however, we found empirically that the slightly increased 36 V pulsing voltage did not cause any signal saturation for this newly implemented catheter probe. We attribute the lesser tendency of this new catheter for signal saturation in the US imaging mode to the smaller inner diameter of the tubing employed (i.e., 2.5 mm compared with the previous 3.17 mm), which placed the membrane farther from the acoustic focal point created by the dual transducers.

The beam diameter of the endoscopic probe was 6 µm; however, we set the total number of scanning steps to 800 for one B-scan as a sort of tradeoff for operating the endoscopic system at a B-scan imaging speed of 20 Hz under the fixed 16 kHz A-line acquisition rate. These parameters yield a transverse displacement of ∼11 µm per unit step when assuming a 2.8 mm diameter virtual circle created by the trajectory of the laser beam along the surface of the imaging window. We applied a laser energy of ∼1.7 µJ per pulse, which was about a twofold increase of the previous value [24] and was achieved by optimizing the related optics; this was to avoid any possible PA signal acquisition failure caused by large distance variations in the urinary tract during a C-scan.

2.3. In vivo rabbit urinary tract imaging

Using this setup, we imaged the urinary tract of a New Zealand white rabbit (1.5–2.0 kg, male), as shown in Fig. 2(c), with the subject placed in a side-lying posture. Multiple trials revealed that it was essential to use a side-lying posture when performing a PAE-EUS pullback C-scan, because if the C-scan was attempted with the animal in a supine position, the probe was spontaneously ejected from the urinary tract immediately following introduction into the target organ. A kind of repulsive force seemed to be generated, probably due to the formation of a curved geometry in the rabbit’s urinary tract with the animal in the supine posture [Fig. 2(d)] and to the tight fit of the probe in the urinary tract. Thus, a side-lying posture was crucial for the acquisition of a 3D data set showing a continuous vasculature map over a full 360° angular field of view (FOV) and a large longitudinal interval. When we tried to image the urinary tract in a supine position, we also noted a tendency for the intended signals to be acquired only from a partial angular region, especially when the probe passed through the bladder region, because it was pressed against one side of the bladder wall rather than being located in a central position.

Prior to the transurethral imaging, we voided urine from the rabbit bladder using a clinical nelaton catheter, as urine might interfere with the intended signals and to situate the bladder wall at the WD. We then introduced the transurethral catheter into the urinary tract at a distance of ∼10 cm, which was assured by the grids marked on the surface of the catheter, and performed PAE-EUS imaging immediately. During the experiment, anesthesia was induced and maintained by the intramuscular injection of ketamine (35 mg/kg) and xylazine (6 mg/kg), and the animal was euthanized with CO₂ gas after multiple C-scans were performed. After euthanasia, the imaged urinary tract was surgically harvested for histological analysis. All animal experiments were performed in compliance with the protocols (YUMC-AEC2022-005) approved by the Institutional Animal Ethics Committee of Yeungnam University, College of Medicine.

3. Results

The urethral probe provided a vascular map of the urinary tract in incredible detail, as shown in Fig. 3(a). The image is a pseudo-color PA radial-maximum amplitude projection (RMAP) image, which can be understood as an unrolled presentation of the urinary tract, as though it had been cut open longitudinally. Two magnified images, which correspond to the dashed boxes in Fig. 3(a), are also presented in the upper row to highlight the high-resolution vasculature imaging capability of the developed imaging probe. The images were produced using the 3,900 B-scan image slices (scanning time: 3.3 minutes), and the pullback scan range and radial image FOV correspond to ∼10 cm and 260 A-line data points (2 mm), respectively. As presented in the magnified images, interesting vascular structures that resemble a natural scrubber appear in the penile urethra and bladder regions. We conjecture that the visualized vascular morphology, which seems to differ markedly from the typical pattern acquired from the rat colorectum [24], represents a unique feature of the urinary system, as opposed to the colorectal system [46]. A close analysis confirmed that a blood vessel with a diameter as small as 11 µm could be successfully resolved, even in the in vivo condition. Due to the limitations in the available space for the main figure, we present the related analysis results in Fig. S5 of Supplement 1.

 

Fig. 3. In vivo rabbit urinary tract imaging results. (a), (b) PA- and US-RMAP images of the rabbit urinary tract, respectively. The two images above (a) are magnified views of the dashed box regions in (a) (Visualization 1 & Visualization 2). The vertical axis corresponds to the angular FOV covering 360°, and the horizontal axis corresponds to the pullback length of ∼10 cm. MD, mid-dorsal; MV, mid-ventral; L, left; R, right. Note that in the RMAP images, PA and US signals are mapped on a logarithmic scale. In (b), “R” represents the acoustic reverberation signals generated by the imaging window. (c) Photo showing the excised urinary tract. Note that the photo was taken after the bladder was filled with water using a clinical nelaton catheter.

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Figure 3(b) shows a US-RMAP image corresponding to the PA-RMAP image. The US-RMAP image was acquired in a co-registered manner based on the simple pulse-echo imaging method rather than the Doppler mode. Thus, no blood vessel-like structures were visualized; only spotty patterns are shown in the image. The limited FOV of the US imaging mode, which was of a depth of less than 2 mm, also precluded the identification of any notable anatomic features that could be used to identify neighboring organs in the US-RMAP image. Therefore, we determined the anatomical location and configuration of the imaged urinary tract with respect to neighboring organs by carefully opening up the animal’s abdomen after euthanasia and identifying the approximate longitudinal locations of the urinary tract, as presented in Fig. S6 of Supplement 1. Figure 3(c) shows a photo of the excised urinary system with the approximate locations of the penile and pelvic urethra, as well as the bladder—note that the photo was taken after the bladder was filled with water to better present the anatomical shape of the excised organs. When producing the US-RMAP image, we did not remove the reverberation signals from the imaging window because, unlike in the previous version [24], the scanning tip was not located at the central position of the imaging window during its rotational scanning due to the highly curved structure of the urinary tract; thus, the reverberation patterns did not appear constantly throughout the acquired 3,900 B-scan images.

Figures 4(a) and 4(b)–4(e) show a volume-rendered image and four representative cross-sectional images, respectively, for the same data set used in Fig. 3 and plotted in a merged manner for the PA (red) and US (green) imaging modes. As shown in Figs. 4(b)–4(e), the PA imaging mode visualizes cross-sections of the blood vessels distributed in the epithelium of the urinary tract. Note that the cross-sections of capillaries appear to have a thorn shape because the radial resolution is lower than the transverse resolution in the OR-PAE imaging mode. In contrast, the US image mode mapped the tissue density distribution with a relatively larger imaging depth. Moreover, the presented images clearly show that the US signal-to-noise ratio (SNR) was greatly increased by the increased pulsing voltage of 36 V. However, the US profile of the imaging window appears to be much weaker than that of the previous case [24]. As stated earlier, we attribute the lowered signal levels of the imaging window to the reduced inner diameter of the employed tubing (i.e., 2.5 mm compared with the previous 3.17 mm), which shifted the membrane position farther from the acoustic focal point created by the dual transducers.

 

Fig. 4. Volume-rendered and cross-sectional images, along with histological analysis results. (a) Three-dimensionally rendered, merged PAE (red) and EUS (green) pseudo-color image of the rabbit urinary tract. The image corresponds to a range of over ∼10 cm with an image diameter of ∼6 mm, and the right-hand side of the image corresponds to the apex of the bladder. The horizontal and vertical scale bars represent 10 mm and 1 mm, respectively. (b)–(e) Representative merged PAE and EUS B-scan images selected from the vertical dashed line depicted in (a). Note that in these images, the profile of the imaging window (i.e., the features denoted by “IW”) did not appear always but appeared intermittently because it was mostly excluded along with the switching noise that is supposed to be located at the inner blank space of the plotted image region. In addition, the random US signals (green) that are distributed inner to the epithelial PA signals (red) and that appear in an irregular pattern are not true signals produced by the urinary tissues but merely artifacts, i.e., the switching noise generated by the RF switch or acoustic reverberations of the imaging window. (f)–(i) Typical histology images (H&E stain) harvested from the indicated locations. Scale bars, 1 mm. Note that the wall thickness and diameter of the specimen rather shrank from its original dimensions during the formalin fixation process.

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In addition to the above observation, the adoption of smaller diameter tubing caused some other unexpected results in the US imaging. First, unlike the previous case [24], in which a tubing with an ID of 3.17 mm was applied, there was overlap between the US echo signals from the Pellethane imaging window and the switching noise generated by the RF switch; more accurately, the very switching noise generated when the internal circuit state of the RF switch was changed from the pulsing mode to the receiving mode (i.e., the noise was generated near the falling edge of the control signal for the RF switch presented in the timing chart in Fig. 2(b)—for the operation principle of the RF switch, refer to Ref. [24]. That is, since the location of the imaging window was closer to the dual transducers, as compared to the previous case, the US echo signals of the Pellethane membrane were frequently mixed with the switching noise as the rotation center of the scanning tip fluctuated inside the imaging window during its scanning. Consequently, the US profile of the imaging window was not included consistently in the plotted image region because in many cases it was excluded along with the switching noise. As shown in Figs. 4(b)–4(e), the profile appeared intermittently over a partial angular region only when the imaging window was deformed in a non-circular way inside the curved urinary tract [Fig. 2(d)] and thereby when its signals decoupled from the locations of the mentioned switching noise in the acquired US A-line signals.

Moreover, due to membrane deformation, the reverberation pattern of the imaging window did not appear constantly but appeared randomly, as shown in the presented cross-sectional images. Thus, being linked to this effect, the random US signals (green) that are located inner to the epithelial PA signals (red) are not true signals produced by the urinary tissues but merely the acoustic reverberation artifacts generated by the imaging window. In addition, it should also be noted that there were no highly prominent acoustic echo signals from the surface (i.e., the epithelium) of the urinary tract. We attribute this to the softness of the epithelial tissues and their radial distances, which were located out of the focal plane created by the dual transducers. Further, the difference in the lumen wall diameters between the bladder and penile urethra regions presented in Figs. 4(b)–4(e) was not significant, unlike those shown in the surgical photo presented in Fig. 3(c), because the PAE and EUS images were acquired after the urine in the bladder was voided—this procedure was critical in terms of making the target tissue wall to be placed within the WD of the imaging probe. Visualization 1 and Visualization 2 are two rotating movies that correspond to the two magnified areas presented in Fig. 3(a).

4. Discussion

In this study, we demonstrated a PAE-based transurethral imaging of the urinary system of a live animal and presented a high-resolution vasculature map of the wall of the rabbit urinary system with a spatial resolution as low as 11 µm. To the best of our knowledge, this is the first in vivo urinary vasculature map ever acquired by any imaging modality—conventional cystoscopy cannot provide as detailed a vasculature map as our device did, and no other clinical tomographic imaging modality, including MR angiography, has demonstrated a comparable level of vasculature mapping capability over such a large area in a vertebrate urinary system. Previous PAT approaches using the rat urinary system relied on transabdominal imaging [40–42], which thus did not allow for the production of a detailed vasculature map. We attribute our high-resolution mapping of the vasculature over such a large area in a vertebrate urinary system to our adoption of the transurethral OR-PAE imaging mode as well as the tapered catheter design.

Indeed, the customized catheter was a crucial component in achieving the presented results. Pilot experiments performed prior to the current study revealed that the smooth introduction of the device tip into the urethra is not so simple due to the narrowness of the orifice and curved nature of the rabbit urinary tract. Thus, although we attempted related imaging multiple times with the conventional catheter design [24], which was, of course, based on the same 1.6 mm OD for the scanning tip, we could not even successfully introduce a related probe into the urinary tract of a rabbit. However, after we eventually reached the presented distal structure, which is soft, flexible, and tapered, we could easily introduce the probe up to the apex area of the bladder (∼10 cm) through the curved urethral tract without causing any recognizable complications. Moreover, through the implementation process of the presented transurethral PAE-EUS probe according to this study, we could work out a more reliable distal optics assembly protocol (Fig. S4 of Supplement 1). Based on these technical achievements, we expect that our developed transurethral OR-PAE technology could be successfully translated into clinics and thus ultimately provide adjunct information about carcinoma in situ lesions in human patients in future clinical use, as these lesions are flat and obscure, unlike papillary lesions, which grow into the bladder cavity and are readily visible.

We successfully acquired a vasculature map from almost all angular regions of the imaging probe with the reported image resolution. However, it may not be possible to acquire a vasculature image of similar quality when applying a probe with the same specifications to humans because of the much larger size of the human bladder. Given the typical volume of a human bladder, we could expect its walls to spread without wrinkles when inflated with a subsidiary tool, such as a medical balloon, to ∼200 cc (which corresponds to a ∼3.7 cm radius, assuming a spherical shape). In this case, based on a simple calculation considering the optical numerical aperture defined by the 3.7 cm WD, a comparable level of image resolution (i.e., ∼10 µm) would be expected using an aspheric lens with a ∼3 mm aperture. The cystoscopes used in clinics have a typical diameter of ∼6 mm, but we believe it would be possible to install a lens with a diameter this large. Of course, if the bladder were expanded as a sphere, a different scanning mechanism would need to be developed because the direction of the side-fired laser beam must also be controllable latitudinally (see Fig. S7 of Supplement 1). If the bladder were expanded cylindrically, the side-scanning mechanism demonstrated here would be applicable, but the probe would not be able to cover the apex and neck areas of the bladder. Regardless of which method is applied, we recognize that future studies must focus on developing a balloon catheter-based, self-internal 3D-scannable probe that resolves the eccentric positioning problem of the scanning tip.

In this study, the radial FOV of the US imaging mode was limited to a depth of only ∼2 mm, so we had difficulty identifying anatomical structures around the urinary tract based only on US image information [Figs. 3(b) and 4(a)]. We purposely set the radial FOV to achieve a 20 Hz B-scan imaging speed under the set A-line data length of 800 points, but we encountered a fundamental limitation in achieving an adequate US SNR for tissue regions deeper than ∼5 mm. This was because the pulser and receiver circuit in the applied universal driver were developed focusing on comparable signal levels of the PA and US, as discussed in our earlier report [24]. In addition to this issue, another possibility was that even narrower blood vessels might have not been detected or visualized because the total number of steps for one B-scan was only 800, which is insufficient when considering the beam diameter (∼6 µm) of the 532 nm laser beam. Although this issue might not be a critical factor when considering the typical diameter of capillaries, which ranges from 5 to 10 µm, following the Nyquist principle would clearly provide a better quality image. Thus, our future study will address these two issues. We will also endeavor to equip the transurethral probe with additional PA imaging capabilities to enhance the imaging depth and add multi-wavelength imaging capability to provide functional molecular image information and imaging of different anatomical structures distributed within and outside the bladder wall, such as the lymphatic system. We also envision the integration of endoscopic optical coherence tomography technology [47–50] into our device so that related probes can provide information regarding layered structures to add to the vasculature image.