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Abstract
A Powell lens is used in a line-field spectral domain OCT (PL-LF-SD-OCT) system to generate a line-shaped imaging beam with almost uniform distribution of the optical power in the line direction. This design overcomes the severe sensitivity loss (∼10 dB) observed along the line length direction (B-scan) in LF-OCT systems based on cylindrical lens line generators. The PL-LF-SD-OCT system offers almost isotropic spatial resolution (Δx and Δy ∼2 µm, Δz ∼1.8 µm) in free space and sensitivity of ∼87 dB for 2.5 mW imaging power at 2,000 fps imaging rate with only ∼1.6 dB sensitivity loss along the line length. Images acquired with the PL-LF-SD-OCT system allow for visualization of the cellular and sub-cellular structure of biological tissues.
© 2023 Optica Publishing Group under the terms of the Optica Open Access Publishing Agreement
1. Introduction
OCT technology can be classified in three categories, point-scanning, full field and line field, based on the scanning approach used for generation of volumetric images [1]. In point-scanning OCT (PS-OCT) a focused Gaussian beam is raster scanned in the transverse (XY) plane while simultaneously recording depth profiles (Z) for generation of a 3D image. Jitter from the mechanical scanning causes phase instability in the X and Y directions in the acquired imaging data. This phase instability can hinder the ability of OCT technology to measure accurately blood flow or functional responses of neural tissue to external stimulation and can cause motion artefacts in OCT images acquired with spatial resolution of ~ 1 µm. In Full Field OCT (FF-OCT), a wide optical beam is incident on the surface of the imaged object and scattered light from is projected onto the sensor of a 2D camera. Volumetric FF-OCT images are acquired by translating the focal plane in Z direction with simultaneous change of the optical pathlength in the reference arm of the system. Since there is no mechanical scanning in the XY plane, FF-OCT offers high phase stability in the transverse imaging plane [1,2]. In Line-Scan (LS) or Line-Field (LF) OCT, a line shaped beam is projected onto the surface of the imaged object and scanned in Y direction while data in Z direction is acquired simultaneously to form a volumetric image [3]. Therefore, LF-OCT offers high phase stability in the XZ plane with some scanning mirror jitter related phase instability in the Y scanning direction [4].
While the concept of LF-OCT was first proposed two decades ago [5], it took more than a decade for suitable fast, 2D camera technology to be developed. Over the past ∼10 years multiple research groups have developed LF-OCT for structural [6], vascular [7,8] and functional [9–11] imaging of the human retina, cellular resolution imaging of the human cornea and limbus [12], as well as cellular resolution imaging of human skin [13–16].
Almost all LF-OCT systems that have been reported so far use a cylindrical lens as the line generator, which results in an elliptically shaped transverse profile of the beam with Gaussian intensity distribution along the major and minor axis of the ellipse. This leads to progressive loss of image contrast from the center of the line (B-scan width) to its edges. The typical quick solution to this problem adopted by many research groups is to simply cut the low contrast areas of the acquired LF-OCT image. However, this approach is wasteful as it significantly reduces the image FOV. Expanding the line length by a factor of 4x or more and using only the central part of the Gaussian profile for imaging can reduce the loss of image contrast in cylindrical lens based LF-OCT systems. However, this approach has multiple drawbacks: a) waste > 50% of the optical power, which will require much more powerful light source to compensate for such a loss; b) require lenses with much larger focal lengths, that will increase the footprint of the LF-OCT system; c) parts of the laser line that are not used for imaging present a safety hazard for both the operator of the system and the imaged subject, therefore they need to be properly blocked, which may not be a trivial task. Digital compensation of the sensitivity loss is another option, though it will result in progressively higher noise level from the center to the edges of the FOV in the line direction.
An alternative approach is to replace the cylindrical lens with a Powell lens, which generates a top-hat light intensity profile in the line-direction. The Powell lens has been used as line a generator for numerous industrial applications for decades since its invention [17]. However, the first Powell lens-based LF-SD-OCT (PL-LF-SD-OCT) system was only reported recently [18]. While the design of this system is compact, simple and likely offers easy optical alignment, it has significant limitations in terms of spatial resolution and sensitivity. The authors used 15 mW optical power incident on the imaged biological tissue sample to achieve ∼87 dB maximum sensitivity at an imaging rate of 3,500. This power exceeds the maximum permissible exposure (MPE) for ocular tissues as defined by the ANSI standard [19] by more than 3 times. Furthermore, the system has ∼16 dB sensitivity roll-off and nearly 2× degradation of the FWHM of the axial PSF over a scanning range of ∼ 850 µm (from ∼2 µm near the zero-delay line to ∼ 4 µm at 850 µm depth). Therefore, this design of the PL-LF-SD-OCT system may not be suitable for ophthalmic applications.
Here, a novel design of a PL-LF-SD-OCT system is presented, which offers ∼2 µm × 2 µm × 1.8 µm (x × y × z) resolution in free space, ∼87 dB maximum for 2.5 mW imaging power at 2,000 fps image acquisition rate. More importantly, the sensitivity loss in a single B-scan along the line direction is only ∼1.6 dB. The system was validated by imaging plant tissue (cucumber) and animal cornea (rats).
2. Methods
2.1 Powell lens
Line generators such as cylindrical lenses and Powell lenses are designed to convert circularly shaped optical beams to line-shaped beams by restricting the propagation of light only in one of the transverse directions of the beam. Cylindrical lenses utilize a two-dimensional spherically shaped front surface, to bend the incident beam in one direction while leaving the orthogonal direction unchanged such that a line-shaped beam is formed. The line-shaped beam maintains Gaussian intensity distribution along the line direction (Fig. 1(A)–1(C)). In contrast, Powell lenses utilize a two-dimensional aspherical, conically shaped front surface, described with Eq. (1):
to generate a line-shaped beam with almost uniform intensity distribution along the line direction (Fig. 1(D)–1(F)) by introducing spherical aberrations in the original light profile [17,20]. Here r is the radial distance from the light propagating axis, k is the conic parameter, and c is the radius of curvature of a sphere. Since the Powell lens does not have a focus, the concept of fan angle, which is the maximum expansion angle of the beam in the line direction, is used to characterize the lens. The back surface of the Powell lens can be either planar or curved, and its profile is used to control the fan angle size.
Fig. 1. Zemax simulations of beam propagation though a cylindrical lens (A) and a Powell lens (D). 1D and 2D cross-sectional light intensity distributions for the cylindrical (B and C) and Powell (E and F) lenses.
2.2 Layout of PL-LF-SD-OCT system
A schematic diagram of the PL-LF-SD-OCT system is shown in Fig. 2. The system is powered by a supercontinuum laser (SuperK Extreme, NKT Phonics). A custom filter unit (combination of long pass (600 nm cut-off), short pass (875 nm cut-off) and neutral density (ND) filters) is used to select a portion of the emission spectrum suitable for this study. A reflective collimator (RC04APC-P01, Thorlabs) is used to generate a collimated beam with 3.9 mm 1/e2 diameter and the measured output power directly after the collimator is ∼12 mW. A Powell lens with 5° fan angle (Laserline Optics) is used to generate a line-shaped beam in the vertical (Y) direction. A telecentric pair of achromat doublets L1 (f = 75 mm) and L2 (f = 100 mm) is used to relay and magnify the beam. A non-polarizing beamsplitter (70:30 (R:T), BS023, Thorlabs) is used to split the incident beam between the sample and the reference arms of the Michelson interferometer. Multiple slits are used throughout the system to eliminate unwanted reflections from the optical components. In the sample arm of the system, the beam is focused (in Y direction) onto a 1D galvanometric scanner (GVS011, Thorlabs) to allow for acquisition of volumetric images. A telecentric pair of achromat doublets L3 (f = 80 mm) and L4 (f = 100 mm) is used to relay and magnify the beam in order to partially fill (∼60%) the entrance aperture of a microscope objective (M Plan APO NIR 10×/0.26 NA, Mitutoyo). The optical power measured at the image plane is ∼2.5 mW, which is below the maximum permissible exposure for human corneal and retinal tissue imaging as defined by the ANSI standard [19].
Fig. 2. A schematic diagram of the PL-LS-SD-OCT system. RC, reflective collimator; PL, Powell lens; L1-L6, achromatic doublets; BS, non-polarized beam splitter; CL, cylindrical lens; NDF, neutral density filter; TS, translation stage; DC, dispersion compensation unit; M, mirror; GS, galvanometric scanner; MO, microscopic objective; S, adjustable slit; G, transmissive grating; L8, camera lens; C, camera CMOS sensor. PL = 5° fan angle, L1 = 75 mm, L2 = L4 = 100 mm, L3 = 80 mm, L5 = 30 mm, L6 = 200 mm, L8 = 85 mm, CL1 = 75 mm, CL2 = 150 mm, and CL3 = 250 mm.
In the reference arm of the system, a cylindrical lens, CL1 (f = 75 mm) forms a telecentric pair with L2 in the vertical direction to convert the diverging line-shaped beam into an elliptically shaped parallel beam. Neutral density filters (NDF) are mounted along the optical path after CL1 to prevent saturation of the camera. A custom-built dispersion compensation unit (a pair of BK7 prisms) is used to compensate low orders of dispersion mismatch introduced by the optical components of the sample and reference arms (what we refer to as “hardware dispersion compensation”, or HDC). An iteration-based custom Python algorithm is used to numerically compensate dispersion mismatch with an approach similar to the one described in the Ref. [21]. An achromat doublet L5 (f = 30 mm) is used to focus the reference beam onto a mirror, mounted on a small translation stage. The mirror, L5 and the DC unit are mounted onto a large linear translation stage to control the optical path difference between two arms.
In the detection arm of the system, a combination of lenses, L6 and L7 is used to relay the beam to the spectrometer. L6 is an achromat doublet (f = 200 mm), while L7 represents a pair of cylindrical lenses CL2 (f = 150 mm) and CL3 (f = 250 mm) with mutually orthogonal orientation, that are used to control the magnification of the beam separately in the X and Y direction [9,22]. The spectrometer is comprised of a volume phase holographic (VPH) transmission grating (990 l/mm @ 805 nm, Wasatch Photonics), and a camera lens (Planar T* 1.4/85, Zeiss). The transmitted optical beam is projected onto the sensor of a 2D CMOS camera (J-PRI, AOS technologies). The camera sensor has an area of 2,560 × 1,920 pixels, with a pixel size of 7.8 µm × 7.8 µm. For this design of the PL-LF-SD-OCT system, all 2,560 pixels were utilized in the spectral direction in order to achieve largest possible OCT scanning range. In the spatial (B-scan) direction, 600 pixels were used for volumetric image acquisition. The system’s acquisition rate was set to 2,000 fps to achieve system sensitivity sufficient for imaging semi-transparent tissues such as cornea and retina. The maximum acquisition rate for the chosen region of interest (ROI) on the sensor (2,560 × 600 pixels) is 6,000 fps.
2.3. Data acquisition and processing
A custom LabVIEW-based algorithm was developed for data acquisition with the PL-LS-SD-OCT system. A set of custom Python-based algorithms were developed for processing of the raw data and generating dispersion compensated images. Amira (ThermoFisher Scientific) was used to render volumetric images and display enface projections from selected ROI.
3. Results
3.1. System performance
The performance of the PL-LF-SD-OCT system in terms of resolution and sensitivity was evaluated using either a protected silver mirror or a United States Air Force (USAF 1951) resolution target as the imaged object, and results from the tests are summarized in Fig. 3. The spectra measured separately at the detection arm of the system from mirror reflections in the reference and sample arms of the system are shown in Fig. 3(A). Note that the spectra were generated by averaging 50 consecutive frames to suppress the effect from the relative intensity noise (RIN) of the light source. The detected spectrum is centered at 730 nm with a FWHM spectral bandwidth of 135 nm. Figure 3(B) shows the system’s axial point-spread function (PSF) measured at a depth of 100 µm relative to the zero-delay line after hardware dispersion compensation (HDC) only (black color) and after additional software dispersion compensation (SDC, red color). The PSF’s FWHM is ∼1.8 µm in free space, corresponding to ∼1.3 µm in biological tissue assuming an averaged reflective index of 1.38. The system’s axial resolution degrades slowly with depth, only by ∼6% over 1 mm scanning range (Fig. 3(C)). The system’s sensitivity was measured for ∼2.5 mW incident optical power and 2,000 fps camera rate. As shown in Fig. 3(D), the median sensitivity measured at a depth of ∼100 µm is ∼87 dB with ∼ 6 dB sensitivity roll-off over 700 µm and ∼13 dB roll-off over the 1 mm scanning range. Figure 3(E) shows the sensitivity results for all 600 A-scans within a single B-scan, measured at a depth location of 100 µm away from the zero-delay line. These results show only ∼1.6 dB loss of sensitivity from the highest peak of the B-scan to its edges.
Fig. 3. (A) Spectra from the reference and sample arms measured at the detection end of the system. (B) Axial PSF measured at depth of 100 µm after HDC and HDC + SDC. (C) Depth-dependent degradation of the axial resolution. (D) Depth-dependent sensitivity. (E) Sensitivity distribution along the B-scan, measured at 100 µm depth. (F) Image of the USAF 1951 resolution target. (G) Normalized intensity profiles acquired from Group 7, Element 6 at the locations marked with the green and red lines in Fig. 3(F).
The system’s lateral resolution and Field of View (FOV) were evaluated by imaging a USAF 1951 resolution target. With the current design of the system, the FOV is 263 µm (X) × 658 µm (Y), corresponding to 300 B-scans (X) and 600 A-scans in each B-scan (Y). Figure 3(F) shows an image of groups 6 and 7 of the resolution target. The intensity plots shown in Fig. 3(G) correspond to the locations in Fig. 3(F) marked with the green and red lines. Since both the horizontal and vertical bars of group 7, element 6 can be clearly resolved and the width of 1 line of this element is equivalent to ∼2.2 µm, therefore both the horizontal and vertical transverse resolution of the PL-LF-SD-OCT system are better than 2.2 µm.
3.2. Images of biological tissue
While the PL-LF-SD-OCT system is designed for imaging the human anterior eye segment (cornea and limbus), due to COVID-19 related restrictions on conducting clinical imaging studies, ethics clearance for use of the PL-LF-SD-OCT system for in-vivo imaging of the human anterior segment (cornea and limbus) has been delayed. Therefore, the performance of the system was evaluated by imaging plant tissue such as cucumber that has optical properties and cellular structure with size of the smallest cells similar to that of the human cornea, as well as imaging animal corneas (rats). For all imaging sessions, the optical power incident on the surface of the imaged object was 2.5 mW and the camera acquisition rate was set to 2,000 fps.
3.2.1. Cucumber
Figure 4 (A) shows a digital photograph of a transverse slice from cucumber. A magnified view (6×) of the ROI marked with the red square in Fig. 4(A) is shown in Fig. 4(B). The red arrow marks a cucumber seed with semi-transparent surrounding tissues. Figures 4 (C-H) were generated using Amira software. XZ, YZ and enface (XY) images of the cucumber seed and the surrounding tissue are shown in Fig. 4(C), 4(D) and 4E respectively, while Fig. 4(F) shows a volumetric image of the same region. Figures 3 (H) and 3 (G) show two enface images from the same 3D stack that correspond to different depths. Small cells of ∼10 µm in size (Fig. 3 G, red arrow) located along the boundary of the cucumber seed (white arrow), as well as cellular nuclei in the larger cells are clearly resolved. Small reflective features were observed in the cytoplasm of larger cells (Fig. 3 H, green arrow), as well as double nuclei in one of the larger cells (Fig. 3 H, blue arrow).
Fig. 4. Images of cucumber tissue. Digital photograph of transverse slide from cucumber (A). magnified view of the region of interest marked with the red square (B). PL-LS-SD-OCT images of the cucumber: XZ projection (C), YZ projection (D), enface projection (E), volumetric image (F), two enface projections corresponding to different depths (H and G). Cellular features are marked with colored arrows in Figs. 4 H and 4G: seed (white), small cells (red), cellular nuclei (orange), reflective features in the cytoplasm (green), cell with 2 nuclei (blue).
3.2.2. Rat cornea
The corneas of male Sprague-Dawley rats (∼ 1 year old) were imaged with the PL-LF-SD-OCT system. All imaging sessions were conducted in compliance with the ethics regulations of the Office of Research Ethics, University of Waterloo. Immediately after euthanasia, rats were placed on a holder mounted on a XYZ translations stage as shown in the digital photograph in Fig. 5(A). A representative B-scan (XZ direction) of the rat cornea is shown in Fig. 5(B).
Fig. 5. (A) Digital photograph of the imaged animal (B). Representative B-scan of the rat cornea (B)EPI -Epithelium; BM – Bowman’s membrane; STR - Stroma; DM - Descemet’s membrane; END - Endothelium; Red arrow: Bowman’s layer, green arrow – basal cell layer of the epithelium. Enface projections acquired from the anterior (C) and posterior (D) stroma. Arrows mark: keratocytes (yellow); thin nerves in the anterior stroma (blue); thick nerve in the posterior stroma (white).
While all 5 of the major corneal layers (EPI – epithelium; BM – Bowman’s membrane; STR – stroma; DM - Descemet’s membrane; and END – endothelium) [23–25] are clearly resolved in the image, many of these layers appear blurred. The blur is caused by the very short (∼18 µm) depth-of-focus of the current design of the PL-LF-SD-OCT system. For the imaging data presented in Fig. 5 (B-D), the focal plane of the microscope objective was positioned at the anterior stroma to allow for imaging of keratocyte cells. The red arrow in Fig. 5(B) marks the Bowman’s membrane, while the green arrow marks the basal cell layer in the corneal epithelium. Figures 5(C) and 5(D) show enface projections acquired from different locations the anterior and posterior stroma respectively. The yellow arrows mark stromal keratocytes, the blue arrow – thin corneal nerves in the anterior stroma, white arrow – larger stomal nerve located in the posterior stroma.
Cross-sectional and enface images of the posterior rat cornea that were acquired with the focal plane positioned at the endothelial layer are shown in Fig. 6. The cross-sectional images (Fig. 6(A) and 6(B)) were flattened and enface images of the endothelial layer (Fig. 6(C) and 6(D)) were generated using maximum intensity projection (MIP). The enface images (Fig. 6(C) and 6(D)) clearly show the honeycomb-like pattern of the endothelial cells, as well as dark, round spots inside the cells that could correspond to cellular nuclei (red arrows) [12,25]. Line artefacts in the enface images (Fig. 6(D), blue arrow) are caused by the integer step-based flattening algorithm.
Fig. 6. Images of the corneal endothelium. XZ (A) and YZ (B) cross-sectional images of the posterior cornea showing the Descemet’s membrane and the endothelium. Enface images of the endothelium (C and D) showing the cellular structure. Red arrows mark cellular nuclei, blue arrows mark artefacts generated by the integer-based flattening algorithm.
4. Discussion
The novel design of the PL-LF-SD-OCT system resulted in very high spatial resolution in biological tissue: ∼2 µm isotropic lateral and ∼ 1.3 µm axial (Fig. 3(B), 3 G and 3F), which was sufficient to visualize the cellular and sub-cellular structure of plant tissues (Fig. 4) and animal cornea (Fig. 5 and Fig. 6). Furthermore, the broadening of the axial PSF function over the entire scanning range (1 mm) was limited to only ∼ 6% (Fig. 3(C)) compared to nearly 100% change reported by the Singaporean research group [18] over a scanning range of ∼ 850 µm. Maximum sensitivity of 87 dB was achieved (Fig, 3D) for 2,000 fps image acquisition rate and 2.5 mW imaging power, which is well below the MPE recommended by the ANSI standard for in-vivo imaging of human ocular tissue. Given equal conditions (same frame rate and imaging power), the novel design proposed here offers an improvement of ∼ 5 dB in the maximum sensitivity measured close to the zero-delay line, compared to the design proposed by the Singaporean group [18]. A very important feature of the proposed novel design is the small (only ∼ 1.6 dB) loss of sensitivity along the width of a B-scan (Fig. 3(E)), which is a very significant improvement compared to the sensitivity loss associated with LF-OCT systems based on cylindrical line generators (∼10 dB or higher) [5–10,12–15]. Since the Singaporean research group [18] has not reported the B-scan sensitivity loss for their design, unfortunately, we cannot provide direct comparison between the 2 designs at this time.
In June 2022, our research group published a cylindrical lens-based design for a LF-SD-OCT system [12]. That design had two major limitations: a) Gaussian distribution of the optical power along the line, which resulted in ∼ 10 dB loss of sensitivity from the center to the edges of the FOV along the line direction; and b) simplified design of the detection arm that did not allow for separate optimization of the transverse OCT resolution in the X and Y direction. The LF-SD-OCT system utilized a femtolaser (INTEGRAL, Femtolasers GmbH) with 3 dB spectral bandwidth of 130 nm, centered at 785 nm to achieve 1.7 µm axial and ∼2 × 3 µm (X × Y) lateral resolution in biological tissue. The lower RIN noise of the femtolaser combined with the camera pixel size (11 µm × 11 µm) contributed to OCT system’s sensitivity of ∼92 dB measured near the zero-delay line for 2.6 mW optical power incident on the images object.
The PL-LF-SD-OCT system described here was designed to overcome some of the limitations of the LF-SD-OCT system. The substitution of the cylindrical lens with a Powell lens reduced the loss of sensitivity across the FOV from ∼10 dB to ∼1.5 dB. The use of supercontinuum laser with spectrum centered at ∼730 nm and 3 dB spectral bandwidth of ∼135 nm resulted in improvement of the axial OCT resolution in biological tissue from 1.7 µm to 1.3 µm. The addition of 2 cylindrical lenses (CL2 and CL3) in the detection arm of the PL-LF-SD-OCT system allowed for independent adjustment of the OCT lateral resolution in the X and Y direction.
While the current design of the novel PL-LF-SD-OCT system offers sufficiently high spatial resolution and sensitivity for imaging the cellular structure of semi-transparent biological tissues such as cucumber (Fig. 4) and rat cornea (Fig. 5 and Fig. 6), the design leaves plenty of room for improvement:
- a) Depth of focus (DOF): The current design resulted in ∼18 µm DOF, which is too short for imaging the cellular structure of the human cornea in one volumetric data set, as seen in Fig. 5(B). One approach to resolving this issue would be to trade lateral resolution for extended DOF by replacing the microscope objective in the current design with a lower magnification one. However, this approach is not desirable, as it will compromise the ability to visualize the cellular structure of corneal tissue, which was the main goal for designing the new PL-LF-SD-OCT system. An alternative approach would be to apply digital adaptive optics (DAO) [7,26–29] to correct for defocus and higher order aberrations in the PL-LF-SD-OCT images. We plan to utilize this approach in the near future.
- b) Sensitivity and sensitivity roll-off: The current design offers maximum sensitivity of 87 dB near the zero-delay line with ∼13 dB sensitivity roll-off (Fig. 3(D)), which is sufficient for imaging the cellular structure of semi-transparent tissues such as the human and animal cornea, though will be problematic for imaging biological tissues that are more scattering such as skin. The sensitivity of the novel PL-LF-SD-OCT system is dependent on several factors: imaging power, image acquisition rate, efficiency of the optical design and RIN of the light source. Increasing safely the imaging power so that it is below the ANSI recommended MPE for ocular tissues is a very limited option, and in our case can result in only ∼1.5 dB sensitivity gain. Decreasing the camera acquisition rate to 1,000 fps will result in 3 dB sensitivity gain, however, this approach will introduce unwanted unvoluntary eye motion artefacts in the in-vivo human corneal images [8,30]. While the use of a broadband supercontinuum laser resulted in improvement of the PL-LF-SD-OCT axial resolution, the higher RIN noise of the laser in combination with the smaller pixel size (7.8 µm × 7.8 µm) of the AOS camera contributed to a slightly lower SNR (87 dB) measured near the zero-delay line with 2.5 mW imaging power. The use of low RIN noise light sources such as superluminescent diodes can further improve the SNR of the PL-LF-SD-OCT. The current optical design of the sample and detection arms of the PL-LF-SD-OCT system includes a large number of optical components that are lossy. Optimizing the efficiency of the collection of light scattered from the imaged object should improve the system’s sensitivity. Furthermore, the smaller camera pixel size poses significant constraints on the optical design of the spectrometer and the PL-LF-SD-OCT system as a whole. The use of a camera with larger pixel size would improve the system’s sensitivity, as well as the sensitivity roll-off, however, this approach will result in shorter scanning range for the same spectral range. Camera efficiency is another factor that contributes to the system’s sensitivity. The use of cameras with better quantum efficiency would improve the system’s sensitivity and possibly the sensitivity roll-off.
- c) Powell lens: While the use of a Powell lens greatly reduced the loss of sensitivity along the width of a B-scan compared to cylindrical lens-based LF-OCT systems, integration of the Powell lens in an OCT system is challenging. The Powell lens is a type of aspherical lens that does not have focus. The center of the optical beam incident on the lens needs to be precisely aligned with the conical tip in the transverse direction and must be coaxial with the central axis of the lens. Even small later shifts or tipping the beam relative to the axial direction results in non-uniform distribution of the optical power along the line length. As the Powell lens does not have a very well-defined focus, it is difficult to generate a collimated beam of the OCT system to allow for changes in the reference pathlength without significant loss of system’s sensitivity due to misalignment between the refence and sample arm beams. In our case, this issue was resolved by use of a cylindrical lens in the reference arm to convert the line-shaped beam to an approximately circular and collimated beam in the reference arm of the system. Also, it should be noted that the quality of the Powell lens varies between manufacturers and imprecision in the Powell lens design can affect negatively both the resolution and the sensitivity of the PL-LF-SD-OCT system.
5. Conclusion
A novel PL-LF-SD-OCT system that utilizes a Powell lens instead of cylindrical lens as the line generator was developed. This design resulted in significantly improved uniformity of the illumination along the line direction and only ∼ 1.6 dB sensitivity loss between the B-scan’s center and edges. The system’s high spatial resolution allowed for imaging the cellular structure of plant tissues and the animal cornea, as well as resolving small morphological features such as cellular nuclei in the endothelial cells. Future clinical applications of the PL-LS-SD-OCT system include in-vivo imaging of the healthy and pathological human cornea and limbus.
Funding
Canadian Institutes of Health Research (PJT-178018); Natural Sciences and Engineering Research Council of Canada (RGPIN-2020-06308, RTI-2021-00780).
Acknowledgments
The authors would like to thank Jean Flanagan for assistance with the animal imaging.
Disclosures
The authors declare no conflicts of interest related to this article.
Data availability
Data underlying the results presented this paper are not publicly available at this time but may be obtained from the authors upon reasonable request.
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