1.1.

Optical Imaging Methods

Optical imaging techniques including reflectance and fluorescence imaging, confocal microscopy, multiphoton microscopy (MPM), and optical coherence tomography (OCT) have proven to be sensitive and specific methods of detection for different cancers.^7–18 A wide-field imaging technique, such as reflectance or fluorescence imaging, can first be used to identify suspicious areas of tissue that may contain cancer. These areas can then be further investigated using a high-resolution technique such as MPM, OCT, or optical coherence microscopy (OCM). Our endoscope combines the three modalities: wide-field narrow-band reflectance imaging, MPM, and OCM.

Reflectance imaging obtains contrast from tissue scatterers and absorbers, particularly hemoglobin. Narrow-band illumination in the blue or green wavelength region enables the identification of regions with increased vascularity, a hallmark of cancer.¹⁹ OCM provides depth-resolved, micron-scale resolution images of tissue microstructure using reflected near-infrared (NIR) light. OCM uses a high numerical aperture (NA) objective to achieve high lateral resolution while sacrificing the extended depth of focus typically achieved with OCT.

Two multiphoton processes of interest in MPM are two-photon excited fluorescence (2PEF) and second harmonic generation (SHG). 2PEF occurs when multiple photons are absorbed nearly simultaneously by a molecule. This molecule then emits a single photon at a wavelength shorter than the incident light yet longer than half the incident wavelength. Fluorescent molecules, including some with excitation or emission in the NIR such as lipopigments and porphyrins, are known to be altered in cancerous tissue. SHG is generated by non-centrosymmetric molecules, such as collagen, which emit light at exactly half the excitation light wavelength. Therefore, combining reflectance, OCM, and MPM in an endoscope may provide a highly sensitive and specific tool for the in vivo detection of cancer in its early stages. As such, it could serve as a minimally invasive alternative to surgical biopsy and histology. Additionally, the co-registered imaging of known features of cancerous tissue, such as increased vascularity, modified tissue and cellular microstructure, altered collagen structure, and changes in biomolecule concentration, can potentially provide additional information about the health of the tissue, further increasing the sensitivity and specificity over a single-modality system.

1.2.

Goals of this Study

With the intent of providing a tool to detect abnormal cells and tissue indicative of early stage cancer, we designed a rigid endoscope that can perform co-registered visible light reflectance imaging for navigation, and OCM and MPM for detailed tissue interrogation. For operational functionality, two additional channels are included in the endoscope, one for irrigation and another for biopsy forceps or other small-diameter tool use. This triple-modality endoscope is small enough for minimally invasive imaging and could be used as an optical screening method for various cancers in the future. For example, it could be utilized as a transvaginal salpingoscope to screen women at high risk for ovarian cancer. It could also be used during surgical laparoscopy to locate cancer margins or metastases, similar to previous studies.^20,21 The endoscope is designed so that the narrow-band, wide-field navigation modality could be used to safely insert the device and navigate into the peritoneal cavity and identify suspicious areas. Once a suspicious area is identified, the front face of the endoscope could be placed in contact with the tissue to perform high-resolution imaging with OCM or MPM. The goal of this study is to describe our endoscope and demonstrate proof of concept.

2.1.

Functional and Physical Requirements

In cooperation with our physician collaborators, a set of functional and physical requirements were derived for the endoscope intended to be used primarily as a transvaginal salpingoscope. These requirements are summarized in Table 1. The endoscope requires a forward-viewing reflected-light modality with a wide field of view (FOV) to enable insertion, navigation to the tissue of interest, and rapid identification of abnormalities. Tissue suspected to be abnormal would then be imaged at high resolution with OCM and MPM to confirm findings. Ideally, the navigation and high-resolution modalities should be co-registered. Additional requirements for clinical utility include an irrigation channel that can either deliver saline for cleaning the optics at the distal tip of the endoscope or deliver exogenous contrast agents for improved labeling of pathological tissue. Additionally, an accessory channel for biopsy tools enables the collection of a small tissue sample for conventional methods of diagnosis.

Table 1

Requirements for the endoscope.

The ability to reach organs of interest, such as the ovary, dictates the physical requirements for both the insertable distal section and the non-insertable proximal section (handle). For the insertable portion, we established basic requirements for rigidity, diameter, length, biocompatibility, and sterilizability. We chose to make the distal section completely rigid for familiarity and allow the physician to have an explicit reference for the endoscope’s location. Therapeutic endoscopes with diameters as large as 7 mm are often utilized. However, for diagnostic purposes, diameters of $<4\mathrm{mm}$ are preferred to minimize tissue trauma. This diameter is large enough to encapsulate the optics and the irrigation and biopsy accessory channels. The length of the insertable section should be large enough for the tip to reach the tissues of interest in a variety of patient anatomies. It was estimated that an insertable length of $\sim 25\mathrm{cm}$ would be sufficient for imaging of ovaries and fallopian tubes. All exposed components must be made from biocompatible materials and be sterilizable without losing performance.

2.2.

Optical Requirements

Clinically relevant requirements for FOV, working distance, resolution, imaging speed, illumination wavelength, and power were established for all of the optical modalities. A summary of the optical requirements is presented in Table 2.

Table 2

Summary of optical requirements.

2.3.

Reflected Light Modality (Navigation) Optical Requirements

The FOV for the navigation modality should be sufficient to identify the target tissue and avoid damage to adjacent tissues, ideally 60 deg full FOV, with 25 deg considered adequate. Working distance should be at least 5 to 20 mm, with a lateral resolution of better than $50\mu \mathrm{m}$ at 5 mm distance at the center of the field. These requirements allow adequate navigation to identify focal areas of disease that are abnormal clusters of 10 or more cells.²² Video-rate (i.e., $30\text{images}/\mathrm{s}$) imaging speed is required to reduce motion artifacts and to enable rapid identification of abnormal areas, so that most of the allowable $\le 30\min$ screening time can be spent with high-resolution modalities. The illumination source should provide high contrast for navigation purposes. For monochrome imaging, blue to green wavelengths are preferred to enhance blood contrast. Laser safety standards should comply as per ANSI Z136.1 standard.

2.4.

High-Resolution Modalities Optical Requirements

The FOV of the high-resolution modalities should be sufficient to image a representative amount of tissue (e.g., at least 10 cell diameters or about $100\mu \mathrm{m}$). The high-resolution mode is utilized when the endoscope is near or in contact with the tissue. A focal distance of $\sim 50\mu \mathrm{m}$ enables visualization of simple epithelia and acceptable tissue attenuation. Cellular resolution ($\text{lateral}\le 2\mu \mathrm{m}$ and axial $\le 10\mu \mathrm{m}$) is necessary to identify pathology. Although having the endoscope in contact with tissue mitigates motion artifact, rapid imaging, ideally at video rate, is still desired. This rapid imaging allows the simultaneous SHG and 2PEF imaging modalities to image over $3{\mathrm{cm}}^2$, and OCM to image over 3 cm linear within the allowable screening time. For OCM and MPM imaging modalities, illumination wavelengths in the NIR range are preferred due to relatively low attenuation and significant penetration for tissue imaging. An OCT source with a 1300 nm center wavelength is ideal for tissue penetration and availability of photonic components. A bandwidth of at least 75 nm is desired to achieve $10\mu \mathrm{m}$ axial resolution in tissue. A variety of wavelengths of MPM lasers are useful, with wavelengths between 1200 and 1500 nm providing advantages of low tissue attenuation as well as ability to excite NIR endogenous fluorophores or exogenous dyes. The MPM source must also provide short pulses with durations in the 100 to 200 fs range to achieve high-excitation efficiency but minimize non-linear processes in fiber while maintaining a high repetition rate for rapid image acquisition and should comply with the ANSI Z136.1 standard.

2.5.

Data Acquisition Requirements

MPM and OCM are techniques in which the image is reconstructed by scanning the FOV point by point or column by column, respectively. Navigation imaging could be performed with either point-by-point reconstruction or a two-dimensional snapshot. However, the former technique was chosen to maintain compatibility with MPM and OCM. As such, a distal two-dimensional scanning system is necessary. Given the need for fiber deliver of light through the endoscope, we have chosen to use a quartered piezo scanning system with cantilevered fiber, described in detail later.

A sufficiently fast detector and data acquisition (DAQ) method must be implemented to meet imaging sampling and speed requirements. Video-rate imaging with $100\times 100\text{pixels}$ per image in the MPM and reflectance imaging modalities requires a sensor with a bandwidth of 1 MHz or greater. Variable gain is useful to account for differing signal strengths. Photomultiplier tubes (PMT) and associated amplifiers used for detecting MPM signals must be highly sensitive in the red to NIR wavelengths and have gain-bandwidth products sufficient for handling weak high-speed signals. Preferably, both MPM signals should be detected and digitized synchronously to ensure co-registered imaging. Since we implemented spectral domain configuration for OCM, the associated line scan camera must be capable of readout speeds in the kHz to obtain video-rate cross-sectional OCM images with 100 line scans or more (i.e., columns) per image.

An equipment rack should enclose all sources, detectors, and computational equipment. It should also be portable and comply with electrical and laser safety requirements. The software interface should control imaging parameters and display images to the user. Table 3 summarizes the acquisition requirements.

Table 3

Summary of acquisition requirements.

4.1.

Distal Tip Configuration

A diagram and photograph of the components housed within the distal ferrule (no protective hypotube), and a distal tip end view of the completed endoscope, are shown in Figs. 3 and 4, respectively.

Fig. 3

(a) Rendering and (b) photograph of the endoscope distal assembly under construction. The protective hypotubing is not yet installed.

Fig. 4

(a) Distal tip end-view rendering and (b) photograph of the completed endoscope.

The custom ferrule (Majer Precision Engineering, Inc., Arizona, USA) is made of medical-grade stainless steel and was designed to hold all the optical components, the scanning system, and the irrigation and accessory channel tubing in a rigid configuration. It has a small outer diameter of 3.3 mm, which fits inside the hypotubing of the insertable portion. The ferrule’s complex shape was achieved using fine wire electrical discharge machining (EDM). Within the ferrule, the largest cutout (1.52 mm diameter) holds the quartered piezo scanning system and a lens system, used for illumination in all three imaging modalities (navigation, OCM, and MPM) and collection of remitted light for OCM and MPM. Twelve small cutouts hold collection fibers for the remitted navigation light. A 1.35 mm diameter cutout holds the polyimide accessory channel tubing (480-II, Microlumen, Florida, USA), which has a 1.34 mm outer diameter and 0.06 mm wall thickness. The resulting 1.22 mm clear channel is sufficiently large to pass accessories such as biopsy forceps (SpyBite, M00546270, Boston Scientific Corporation, Massachusetts, USA). The final 0.97 mm diameter cutout holds smaller diameter tubing (335-II, Microlumen, Florida, USA) with an outer diameter of 0.96 and a 0.05 mm wall thickness. This tubing is used for irrigation with saline and the delivery of contrast agents. The latter two functionalities are shown in Fig. 5.

Fig. 5

(a) Biopsy forceps (SpyBite) inserted through the accessory channel and (b) water being flushed through the irrigation channel.

4.2.

Optical and Scanning System Configuration

All three optical modalities utilize a common double-clad fiber (DCF) (DCF13, Thorlabs, New Jersey, USA), scanning system, and lens system. All modalities utilize the single-mode core of the DCF for illumination. Based on the requirements in Sec. 2, the specified illumination wavelength, remitted signal wavelength, illumination spot size (resolution), illumination FOV, and collection mechanism for the three modalities are specified in Table 4.

Table 4

Illumination and collection mechanisms for the different optical modalities.

An illumination wavelength of 561 nm was chosen for navigation based on laser availability and optical tissue properties. This wavelength provides high contrast when imaging blood vessels due to the oxy- and deoxy-hemoglobin absorption peaks near this wavelength.²³ A laser diode with 35 mW power output was used (R560503FX, LaserGlow Technologies, Toronto, Canada). For OCM, a dual-superluminescent diode with a center wavelength of 1300 nm, 70 nm full-width-half-maximum bandwidth, and 8 mW power was chosen (Broadlighter 1300, SuperLum, Co., Cork, Ireland). For MPM, a custom $\sim 100\mathrm{fs}$, 40 MHz pulsed laser centered at 1400 nm with 160 mW output power was utilized.²⁴ Laser pulse dispersion in the optical fiber was mitigated by prechirping with a grating pair. The power output measurements at the distal tip of the endoscope were lower than the source outputs, and they were confirmed to meet ANSI Z136.1 standard for permissible laser exposure on skin. The beam rapidly diverges from the endoscope and operational controls (e.g., lasers only energized with endoscope inside the patient) limiting eye safety concerns.

The differing illumination spot sizes and FOVs of the reflectance system compared to those of the OCM/MPM system are made possible by the novel lens system previously described.²⁵ In brief, the lens system contains a triplet with two dichroic coatings at the distal and proximal inner surfaces in the shape of a dot and a ring, respectively. The coatings are arranged similar to the primary and secondary mirrors of a telescope. The coatings transmit light at visible wavelengths so that the final objective lens weakly focuses the light into a low 0.007 NA beam with a depth of focus from 5 to 22 mm and a central spot size of $10.98\mu \mathrm{m}$ at the best focus. The reflectance signal from the illuminated spot is collected with twelve 0.5 NA multimode fibers held by the ferrule, enabling good light collection efficiency. At NIR wavelengths, the coatings reflect light in a telescope-like configuration to expand the beam, and the final objective lens strongly focuses them to a high 0.44 NA beam with a $50\mu \mathrm{m}$ front focal distance and an $\sim 2\mu \mathrm{m}$ spot size. The OCM light is collected back into the single-mode core of the DCF, as required for this interferometric technique. The MPM signal is collected in the 0.2 NA first cladding of the DCF, which enables increased light collection over utilizing the core only. Note that MPM only captures the central portion of the collected light since the wavelengths generated by the non-linear effects are not affected by the dichroic surfaces. The optical paths used are illustrated in Fig. 6.

Fig. 6

Tunnel diagrams, where light propagates form left to right including reflected light, for the illumination, optical paths, and collection mechanisms for the different optical modalities. The sample is represented with an orange line.

The quartered-piezo (piezo) scanning assembly with a cantilevered DCF and fiber mass is shown in Fig. 7. Scanning endoscopes have been demonstrated before using Lissajous and spiral scanning patterns.²⁶ Our endoscope works with a raster pattern since relatively slow (tens of Hz) scanning is required for OCM, and relatively fast (thousands of Hz) scanning is required for video-rate en face navigation and MPM imaging. Raster scanning can also ease image reconstruction because it is acquired in a more intuitive sequence than spiral or Lissajous scanning. However, image calibration adjustment is still needed due to the sinusoidal motion of the piezo. Raster scanning at slow and fast frequencies can be achieved by utilizing multiple vibrational resonance frequencies (e.g., first and third) that can be adjusted by tuning the fiber length and utilizing fiber-mounted masses.²⁴

Fig. 7

Quartered piezo scanning assembly including the quartered piezo tube, fiber, wires, collar, and brass mass. The brass mass is shown sectioned through the middle to show the internal “bow-tie” shape, the fiber channel, and the hole used to glue the mass to the fiber.

The scanning system shown in Fig. 7 utilizes a custom piezo with 1.0 mm outer diameter, 0.2 mm wall thickness, 11 mm length, 4 longitudinal outer electrodes (electrically isolated quadrants at 90 deg), and PIC255 material (PI, Massachusetts, USA). A custom polyether ether ketone piezo collar (Optics Technology, New York, USA) holds the proximal piezo stationary and centered inside the ferrule cutout while allowing the distal piezo end and the cantilevered fiber to oscillate with relative freedom. Additionally, the collar protects the exposed leads that apply voltage to the piezo electrodes. Desired resonant frequencies of $\sim 38\mathrm{Hz}$ first resonance and $\sim 3800\mathrm{Hz}$ third resonance were achieved using 16 mm long cantilevered fiber with a custom brass mass (Optics Technology). The mass shape is cylindrical to maximize its weight in the 1.52 mm ferrule cutout while allowing a tip deflection amplitude of up to $400\mu \mathrm{m}$ peak-to-peak. The cylinder’s inner diameter was shaped in a “bow-tie” fashion and glued to the fiber at one point only, through a hole on the side of the mass. Thus the fiber is free to bend and oscillate as required at third resonance. The mass is placed such that the center of mass sits at a location between 75% and 80% of the fiber’s length, aiming for the third resonance distal nodal point. The mass, placed as described, lowers the first harmonic frequency while keeping the third harmonic frequency relatively unchanged,²⁴ allowing for a 100× difference in the two frequencies for raster scanning.

The ratio of scanning frequencies limits the number of pixels in one dimension of the image (the slow axis), in this case, to 100 pixels, and the DAQ board sampling rate of 2 MHz limits the number of pixels in the other dimension (the fast axis) to about 130 pixels. For the high NA imaging modalities, 100 pixels enable Nyquist sampling given the $100\times 100\mu \mathrm{m}$ FOV and $2\mu \mathrm{m}$ resolution. For low NA reflectance imaging, the image may be undersampled depending upon the FOV chosen. However, it is worth noting that the number of image pixels obtained is similar to a 10,000 elements fiber bundle, which has been widely used for miniature endoscopes.²⁷

4.3.

Endoscope Handle

The fibers, piezo control wires, and irrigation and accessory tubing are carried through the hypotube to the 3D-printed handle, where they are separated into three different channels, as shown in Fig. 8. The center channel contains the 12 multimode collection fibers, the dual-clad fiber, and the four piezo control wires. A curved side-channel houses the accessory channel tubing with a 1.22 mm inner diameter. The tapered entrance allows the tool to be inserted smoothly. A second side-channel holds the Luer lock tip of a 21G hypodermic needle with a shortened cannula (Exelint International, California, USA), which is mated with and glued into the irrigation channel tubing. This Luer lock connector provides an attachment point that allows for the introduction of saline or contrast agents with a standard Luer lock syringe.

Fig. 8

(a) Open handle showing the internal channels guiding the different components and (b) handle with components installed.

Proximal to the handle, the fibers and piezo control wires are protected with a 2.0 m long, 3/16 in. outer diameter vinyl tube (4027470, ACE, USA). The tubing is glued to the handle and protected with a separate heat shrink tubing at the handle interface (MT5000-1/2-0-SP, TE Connectivity, Schaffhausen, Switzerland). Components exit the vinyl tube’s proximal end and lead into three connectors that attach the handheld endoscope to the instrument rack. A standard FC/APC connector is used for the DCF, a standard SMA connector is used for all 12 multimode fibers combined into a single $900\mu \mathrm{m}$ lumen, and a 4-pin connector is used for the piezo control wires (SP13-4pin, Szjelen, China). A complete prototype of the handheld endoscope is shown in Fig. 9.

Fig. 9

Endoscope prototype in sterilization box.

All exposed materials were carefully selected to ensure compatibility and sterilizability of the endoscope. The hypotube is made of 304 stainless steel, and a custom cover glass is made of fused silica (W-SIO2UVE-RS0003, Bern Optics, Massachusetts, USA). The irrigation and accessory tubing is made of polyimide. The gaps between the cover glass and the hypotube and the gaps between the inner irrigation and accessory tubing and the cover glass are sealed with a medical-grade epoxy (EP42HT-2MED Clear, Masterbond, New Jersey, USA). Gaps between components such as the handle-hypotube interface are further sealed with medical-grade adhesive (Loctite 4013, Henkel Adhesives, Düsseldorf, Germany). The 3D-printed handle material (VeroBlack) and the vinyl tubing are not biocompatible, but they can be covered with a sterile ultrasound sheath during clinical use. Sterilizability of the entire handheld endoscope was confirmed with deliberate contamination, then swabs for common pathogens after hydrogen peroxide sterilization. All tests revealed that no pathogens were present after sterilization.

5.1.

Navigation Components

Referring to the system diagram in Fig. 1, the 561 nm laser is connected to a custom wavelength division multiplexer (WDM) (OZ Optics, Ontario, Canada) optimized for combining this wavelength with the 1300 nm OCM source. The WDM output then enters a custom WDM/dual-clad fiber coupler (DCFC) (DBM-1300-200901, Castor, Quebec, Canada), which adds in the 1400 nm MPM source and directs all wavelengths into the core of the DCF that runs through the endoscope. The single $900\mu \mathrm{m}$ lumen SMA connector containing light reflected from the sample is mated to an SMA connector containing a single $1500\mu \mathrm{m}$ core fiber at the instrument rack interface (M93L01, Thorlabs). This fiber is coupled to an SMA-adapted silicon detector (New Focus 2107-FC-M, MKS, California, USA) with a 6-MHz bandwidth at the $3\times 102$ gain used during testing.

5.2.

OCM Components

The 1300 nm OCM source is coupled into a fiber interferometer containing a circulator (37474 FOC-12N-11, OZ Optics) and a 90:10 fiber coupler (FUSC-280125223, Precision Micro-optics, Massachusetts, USA). Ten percent of the power is directed into a reference arm with dispersion matching, polarization control, and an optical path length adjustable to match the endoscope. The remaining 90% of the power is directed into the WDMs for passage through the DCFC and output through the DCF core. Light reflected by the tissue into the DCF core follows the same return path, where it is recombined with reference arm light and directed by the circulator to the spectrometer. The spectrometer is equipped with a reflective collimator (MPD124-M01, Thorlabs), 300 mm focal length custom-made lens (AC254-300-C-ML, Thorlabs), reflective grating (GR25-0613, Thorlabs), and a 2048 pixel linear array (GL2048L, Sensors Unlimited, New Jersey, USA). This camera was chosen due to its high quantum efficiency (60% to 70%), user-controlled exposure, high dynamic range ($>1200:1$), and appropriate read rate of up to 76,000 lines per second.

5.3.

MPM Components

MPM source light from the custom 1400 nm femtosecond laser is combined with the other wavelengths into the DCF core via the second WDM and DCFC. Remitted light in the DCF inner cladding is delivered via the multimode arm of the DCF to a custom-made PMT enclosure. A reflective collimator (RC02FC-P01, Thorlabs) delivers the light to a series of dichroic mirrors that separate SHG and 2PEF (FF01-890/SP, FF725-SDi01-$25\times 36\times 3.5$, FF01-850/310-25, FF01-770/SP-25, Semrock, New York, USA) and direct the signals to PMTs (H10721-20, Hamamatsu, Iwata, Japan). Transimpedance amplifiers (PMT-4V3, Advanced Research Instruments Corp., Oregon, USA) provide variable gain with a 2 MHz bandwidth at the $1\times {10}^6$ gain used in testing. A programmable filter (3945, Krohn-Hite, Massachusetts, USA) can be used to remove high-frequency noise.

5.4.

Data Acquisition, Software, and Display

To generate piezo control signals, acquire detector/camera data, and reconstruct and display images, a custom LabView program was written in collaboration with INTAQ (Arizona, USA). Based on user input image dimensions and image rate, piezo control signals for each axis are generated and output through a multipurpose DAQ board (PCIe-6321, National Instruments, Texas, USA). The control signals are further amplified at the piezo driver (TD250, Piezo Drive, Shortland, Australia). One-dimensional slow-axis scanning is used for OCM imaging, and two-dimensional fast- and slow-axis raster scanning is used for reflectance and MPM modes. Piezo control signals are synchronized with the data acquisition. Reflectance and MPM signals are acquired with the same multipurpose DAQ board, capable of up to four synchronous analog input channels at up to 3.5 MHz each (PCIe-6374, National Instruments). Due to limitations in the current software only acquiring on one side of the sinusoidal piezo movement, the number of pixels in the slow scanning imaging dimension is limited to 50. The number of pixels in the fast scanning dimension is only limited by the sampling rate of the DAQ board and the piezo fast scan frequency and could be up to 460. OCM line scan camera signals are acquired with a frame grabber (PCIe-1433, National Instruments). The number of A-scans (line scan camera reads) per image is limited by the camera line rate and the piezo slow scan frequency and could be up to 950.

The software enables the acquisition and the display of navigation, OCM, or MPM images. Navigation and MPM images are displayed in real-time in their as-acquired form, which is distorted due to the scanning system’s sinusoidal movement, piezo scanning axes that may not be perfectly perpendicular, and minor stable whirling of the fiber. In practice, typically 10% of data points at each image edge are excluded, as the fiber movement at these locations is so slow as to cause significant distortion and provide limited additional FOV. Therefore, the number of pixels in the slow scanning imaging dimension is normally 40. A calibration obtained from imaging a grid pattern can be applied to process images into their true dimensions.

OCM camera reads are processed by applying calibration files created using published methods²⁹ and taking the Fourier transform of each A-scan. B-scan images are constructed by combining processed A-scans. Acquired raw data, processed data, and configuration files are saved upon request.

6.1.

Reflectance Test

The endoscopic optical system was tested in a tabletop configuration without the hypotube to allow for alignment of the components. Reflectance modality images of a black on white paper USAF resolution target were analyzed (#53-715, Edmund Optics, New Jersey, USA) to evaluate FOV and lateral resolution. At a working distance of 5.0 mm, up to a $2.5\mathrm{mm}\times 2.5\mathrm{mm}$ FOV could be imaged, which equates to a 28 deg FOV. To evaluate lateral resolution, the FOV was reduced to $\sim 0.25\times 0.25\mathrm{mm}$ by controlling the piezo oscillation amplitude voltage. The reduction of the FOV resulted in $\sim 6\mu \mathrm{m}$ distance between pixels when a square image of $40\times 40\text{pixels}$ was obtained. The contrast was calculated utilizing the grayscale values of the images, and the resolving criterion was set to a contrast of 0.01 or greater using:

Grayscale values were read from a line traced through the horizontal or vertical bars of the respective element. An $I_{\max }$ or $I_{\min }$ was taken at each line’s inflection points, and four contrast values were calculated, one for each adjacent $I_{\max }$ and $I_{\min }$ values, using the contrast equation. The average of these contrast values is the reported value. Figure 10 shows reflectance images of the USAF resolution target. The smallest element with contrast above the threshold is group 4 element 3, equivalent to a lateral resolution of $24.80\mu \mathrm{m}$ or 20.16 lines per millimeter. Vertical lines displayed a contrast of $\sim 0.024$, and horizontal lines displayed a contrast of $\sim 0.011$. The differences in contrast between the horizontal and vertical fringes seen in Fig. 10 may be attributable to the suboptimal alignment of the lenses causing astigmatism.

Fig. 10

Reflectance images of a USAF target showing the (a) vertical and (b) horizontal lines for group 4 element 3. The contrast and brightness of these images were digitally adjusted for clarity. (c) Grayscale values extracted from the red lines traced over the horizontal and vertical patterns shown on the left side.

To further test the capability of reflectance imaging and to measure the working distance range, images of a reflectance target were obtained at 5, 10, 15, and 30 mm from the endoscope distal cover glass. The target consists of a surface coated with Spectralon (Labsphere Inc., New Hampshire, USA) and geometrical shapes made of Fineshut KIWAMI (Koyo Orient Japan Co. Ltd., Saitama, Japan). These materials provide some of the highest and lowest, respectively, reflectance values obtainable, thus providing higher contrast than the USAF target. The reflectance images and the target white light image can be seen in Fig. 11. The finite thickness of the Fineshut material causes some artifacts, especially at close range. The reflectance images show that the shapes are visualized over the entire range with a contrast of up to 0.58. At 30 mm, the low collection efficiency of the reflected signal results in a noisier image.

Fig. 11

(a)–(d) Images of a reflectance target in white light and reflectance at 5, 10, 15, and 30 mm from the endoscope distal cover glass. (e) The white light image has been labeled for the approximate location of the scans.

6.2.

OCM Test

To test OCM, the lateral FOV was reduced to $\sim 50\mu \mathrm{m}$ by controlling the piezo oscillation amplitude voltage. The reduction of the FOV resulted in $\sim 0.3\mu \mathrm{m}$ distance between A-scans with a total of 166 lines per frame. The best focus was found to be at $\sim 60\mu \mathrm{m}$ from the distal cover glass. A variable line grating test target was imaged to test OCM lateral resolution (R1L3S6P, Thorlabs). The OCM image is shown in Fig. 12. This test was performed as a simple pass/fail test for OCM since the system’s lateral resolution is theoretically $2\mu \mathrm{m}$, but the smallest pattern in the target is $4\mu \mathrm{m}$. This image shows excellent contrast between the lines in the lateral direction with a signal-to-background ratio (SBR) of about 68.

Fig. 12

OCM image of a variable line grating test target (R1L3S6P, Thorlabs, New Jersey, USA).

The $14.2\mu \mathrm{m}$ FWHM of the line in the axial direction (vertical on the image) is significantly degraded from the theoretical axial resolution of $\sim 9\mu \mathrm{m}$. We attribute the discrepancy to suboptimal calibration of the spectrometer.

6.3.

MPM Test

The broken edge of a gallium arsenide (GaAs) sample was imaged to test the MPM modality due to the strong emission of 2PEF at the edge. The FOV was reduced to $\sim 50\times 50\mu \mathrm{m}$ by controlling the piezo oscillation amplitude voltage. The reduction of the FOV resulted in $\sim 1.25\mu \mathrm{m}$ distance between pixels when a square image of $40\times 40\text{pixels}$ was obtained. The best focus was found to be at $\sim 60\mu \mathrm{m}$ from the distal cover glass. The multiphoton image obtained by averaging 32 frames is shown in Fig. 13. The edge image allows us to use a non-linear technique to characterize lateral resolution.³⁰ This technique consists of fitting the edge profile using:

Fig. 13

(a) 2PEF image of a GaAs edge and (b) the profile taken from the white line in the image.

The resolution computed by this technique is $3.97\mu \mathrm{m}$. Additionally, an image of a computer chip was acquired to further test the MPM capability and determine SBR. The image can be seen in Fig. 14, showing two wires that are used to connect the chip’s internal components. Chip materials such as silicon-doped hafnium oxide and metals are likely the SHG and 2PEF sources. SBRs of 8.1 and 10.4 for SHG and 2PEF, respectively, were measured from the image.

Fig. 14

(a) SHG, (b) 2PEF, and (c) white light images of computer chip. The long vertical strands in the center are wires that connect the internal components of the chip.