In this paper, we have demonstrated a SECM probe that comprehensively images luminal specimens while adaptively adjusting the focus so that it always resides within the tissue. The images of a tissue phantom and an excised swine tissue demonstrated the effectiveness of the adaptive focusing mechanism and the feasibility of acquiring three-dimensional volumetric images using a single helical scan of the probe. Several technical challenges need to be addressed before the SECM probe is applicable for human imaging in vivo, however.
The image acquisition speed in this study, 25-kHz line rate, was slower than the maximum speed of the line scan camera in the spectrometer, 70 kHz, mainly due to the method used for generating the feedback control signal. The control signal was created within the image acquisition software coded in the LabVIEW platform (National Instruments, Austin, TX), and the computational burden of generating the control signal in this programming environment limited the image acquisition speed. Developing more efficient code would allow the image acquisition to fully utilize the maximum speed of the line scan camera. The image acquisition speed can be also increased by separating the feedback control signal generation from the image acquisition software. A simple optoelectronic apparatus comprising a grating, a position-sensitive detector (PSD; i.e. quadrant photodetector), and an electric control circuit could be used to analyze a portion of the spectrum of the returning light from the SECM probe and generate a control signal independently of the image acquisition software.
The imaging speed needed to be further reduced when imaging biological tissues due to the low level of detected intensity. The light throughput of the SECM probe was only 2%, partly caused by the use of off-the-shelf optical components that were not optimized for the wavelengths of light used in our setup. We can improve the light throughput by customizing the optical components, and can utilize a supercontinuum source with a 3 times higher spectral density (SC450-6, Fianium, UK) than the present source, resulting in an order-of-magnitude improved sensitivity, thereby allowing the line scan camera to operate at its maximum speed.
The adaptive focusing mechanism was demonstrated to provide in-focus images over most of the imaged areas ( and ). However, due to the slow update rate of the control signal, there were regions where the adaptive focusing did not track the tissue surface reliably (dotted arrow in ). In our experiment, the control signal was updated discretely at a rate of 125 Hz; rapid changes of the focal distance during this 8-msec interval caused the tissue surface to be outside of the focus of the SECM probe. More frequent updates of the control signal will enable improved continuous tracking of the tissue surface and provide in-focus images over even larger areas. Higher control signal throughput can be achieved by developing more efficient code and/or employing an independent apparatus of generating the control signal as mentioned previously.
Increasing the control signal update rate will also enhance the dynamic performance of the adaptive focusing. The current adaptive focusing mechanism can track up to a 1-Hz dynamic focal deviation with a displacement amplitude of 250 µm. Based on previous experience in esophageal imaging of living animals and human patients with a balloon-centering catheter [2
], a focal deviation of ±250 µm at the rate of 2 Hz is anticipated in human esophageal imaging. With the improvement on the update rate described in the previous paragraph, we anticipate that the adaptive focusing mechanism will be able to reliably track the dynamic focal deviations encountered when imaging human patients.
We used a PZT-driven linear actuator for adaptive focusing in this paper due to its small size, large travel range and fast moving speed. Several other variable focusing methods have been previously studied for endoscopic microscopy applications, including moving the distal tip of the fiber pneumatically and mechanically [24
] and using a pressure-controlled variable focus liquid lens [25
]. However these methods need significant modifications to transfer the force or the pressure rapidly from the proximal end to a rotating SECM probe. MEMS deformable mirrors have been used to change the imaging depth of optical coherence tomography (OCT) [26
] and are expected to provide a focusing range as large as 1mm for an objective lens with NA of 0.4 [27
]. Electrically-tunable varifocal liquid lenses [28
] and voice-coil motors [29
] have been used to conduct auto-focusing in cell-phone cameras. In future probe development, we will investigate these electrically-driven variable focusing mechanisms and will select a method that meets the requirements of SECM imaging and that can be integrated within a rotating imaging probe.
While the SECM images of a resolution target and a tissue phantom demonstrated that the SECM probe has good transverse resolution, cellular features were not clearly visualized in the swine tissue images. The off-the-shelf aspheric singlet that was used as the objective lens was designed for use in air, which increased the specular reflection from the tissue surface and caused spherical aberration when imaging below the tissue surface. In the future, we will fabricate a custom objective lens that allows for water immersion and will fill the centering balloon with the immersion medium, decreasing the specular reflection and the spherical aberration and subsequently enabling high-resolution imaging of sub-surface regions of the tissue. Speckle noise presented in the SECM images also made it hard to appreciate cellular features. The speckle noise will be reduced utilizing a single-mode illumination and multi-mode detection method though a double-clad fiber [30
]. Although the rotational scan of the probe provided good spatial registration between circumferential scans and generated large-area images without noticeable stitching artifacts (, and ), there were locations where non-uniformity of the motor rotation speed made it difficult to mosaic circumferential scans together with microscopic precision. The rotational non-uniformity was primarily caused by the low torque capacity of the motor. We will utilize a motor and a driveshaft with a higher torque capacity to provide more uniform rotation of the probe. A rotary encoder can be integrated into the probe to measure the actual rotation speed, which can be used to correct the image distortions generated by residual rotational non-uniformity.
The spectral bandwidth of 30 nm was dispersed over 1024 pixels of the line scan camera, resulting in an effective coherence length of 11 mm. While interference fringes were not likely generated from the probe optics due to the anti-reflection coating of each component, relatively large distances between the components, and confocal gating, interference between back-reflections from the inner and outer surfaces of the plastic tube generated visible fringes (). We expect that the water-immersion approach described above will significantly reduce the back-reflection from the tube’s inner surface and will subsequently decrease the visibility of the interference fringes. When imaging lens paper, fringe patterns were imposed on the SECM images obtained from superficial regions (D = 14 µm in ), caused by the interference between reflections from the tube outer surface and the specimen. However, fringe patterns were not observed when imaging swine tissue because the constant contact between the plastic tube and the tissue reduced the back-reflection from the outer surface of the tube.
The sub-optimal performance of the aspheric objective lens also limited the field depth to 56 µm. In the next version of our SECM probe for human imaging, we will increase the field depth to 100 µm by designing a custom objective lens that has a larger diffraction-limited FOV for the spectral band of interest. Since most epithelial disorders manifest near the surface, an imaging depth range of 100 µm is expected to provide sufficient histomorphologic information to render accurate diagnosis.
The next step in our research will be to construct a clinically-viable SECM probe that can be used in human patients. The technology development will be focused on addressing the challenges discussed above and reducing the probe size further. We anticipate that the next-generation clinical SECM probe will acquire volumetric confocal images of the entire distal esophagus with a volume of 39 cm2
(surface area) × 100 µm (ranging depth) in less than 10 minutes. Following the image acquisition, the comprehensive volumetric data will be analyzed to locate regions with high probabilities of harboring severe dysplasia or early cancer. The SECM probe can then be repositioned to the identified high-risk regions, and high-power laser light will be delivered through the SECM probe to generate laser-burn marks around the regions [4
]. The marks will be visible under video endoscopy, and clinicians will be able to take biopsies from the high-risk regions rather than random locations, which will increase the likelihood that patients will receive a much more accurate diagnosis than the current standard of care.