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Abstract
Corneal imaging is important for the diagnostic and therapeutic evaluation of many eye diseases. Optical coherence tomography (OCT) is extensively used in ocular imaging due to its non-invasive and high-resolution volumetric imaging characteristics. Optical coherence microscopy (OCM) is a technical variation of OCT that can image the cornea with cellular resolution. Here, we demonstrate a blue-light OCM as a low-cost and easily reproducible system to visualize corneal cellular structures such as epithelial cells, endothelial cells, keratocytes, and collagen bundles within stromal lamellae. Our blue-light OCM system achieved an axial resolution of 12 µm in tissue over a 1.2 mm imaging depth, and a lateral resolution of 1.6 µm over a field of view of 750 µm × 750 µm.
© 2022 Optica Publishing Group under the terms of the Optica Open Access Publishing Agreement
1. Introduction
Corneal disease is the one of the leading causes of visual loss globally making it crucial for studies related to corneal health [1,2]. The cornea, the optically transparent and outermost part of the eye, is vital for light transmission. It consists of five different layers including the epithelium, Bowman’s layer, stroma, Descemet’s membrane, and the endothelium. Some of the corneal structures that exist in these layers include the anterior and posterior stromal keratocytes, stromal nerves, and visible cell nuclei. Currently, specular microscopy (SM) and in vivo confocal microscopy (IVCM) technologies are used in clinical settings for cell evaluation, offering very high lateral resolution to observe individual cells and even cell nuclei in corneal tissue [3]. However, SM and IVCM have some disadvantages, such as limited field of view (FOV) resulting in prolonged image acquisition and processing time, and lack of volumetric imaging capability [4]. SM is only able to visualize the endothelial cell layer, while IVCM can demonstrate all cellular layers of the cornea and even some infectious pathogens, but requires contact with the eye and is highly operator-dependent [5].
Optical coherence tomography (OCT) is a non-invasive imaging modality used in many fields, including ophthalmology [6]. A key feature of OCT is the independent control of axial and lateral resolutions: axial resolution is related to the central wavelength and spectral bandwidth of the light source, whereas the lateral resolution depends on the numerical aperture (NA) of the focusing lens in the sample arm [7]. Currently, most OCT systems use near-infrared light to image the retinal layers due to its ability to penetrate the deeper layers of scattering tissue and its commercial availability [8]. Additionally, in recent years, visible light OCT has been expanding in ocular imaging to increase contrast in ex vivo and in vivo tissue structures with sub-micrometer axial resolution [9,10]. Using shorter wavelengths has been reported to help increase the contrast of corneal lamellar structures utilizing lasers with a second-harmonic generation [11–14]. OCT systems mostly use low NA objectives combined with broadband sources such as supercontinuum, femtosecond lasers, and a less expensive alternative, super-luminescent diode, to achieve high axial resolution, but the low NA objectives ultimately limit the lateral resolution to tens of microns [3,7,15–18]. Optical coherence microscopy (OCM) is an imaging modality based on OCT and confocal microscopy that incorporates large NA objectives to achieve high lateral resolution for the visualization of cellular structures [19]. Some OCM technologies that also demonstrate micrometer resolution for corneal imaging are ultra-high resolution OCT, micro-OCT with a large NA and spectral domain system, time-domain full-field OCT, Fourier domain OCT, and Gabor-domain OCT [4,20–28,3].
Using short-wavelength super-luminescent diodes for coherent signal detection can be challenging but would reduce system costs while utilizing the advantages of improved contrast for corneal structures. In this study, we present a blue-light OCM system with a center wavelength of 450 nm and a 6 nm bandwidth. Applying a super-luminescent diode and a suitable microscope lens with 10× magnification simplifies the system setup and significantly reduces costs. We imaged ex vivo corneal cellular structures, and were able to visualize epithelial cells, endothelial cells, and anterior and stromal collagen lamellae. Additionally, we evaluated the system's optical performance to characterize its axial and lateral resolution, effective FOV, and sensitivity falloff over the entire imaging depth in air.
2. Methods
2.1 Design of the optical coherence microscopy system
The blue-light OCM system was based on a spectral domain OCT engine. The schematic diagram of the blue-light OCM is shown in Fig. 1. The light source was a super-luminescent light emitting diode (EXS210099-03 Butterfly, EXALOS, Switzerland) centered at 450 nm with a 6 nm full width at half maximum (FWHM) bandwidth, and was guided into a 50:50 fiber coupler (TW470R5A2, Thorlabs Inc., USA). The sample and reference arms were connected to 1.1 m and 1 m fiber patch cords (460HP, Thorlabs, USA) respectively to compensate for the dispersion mismatch. The output power from the single mode fiber pigtail (S405-XP) of the light source was 5 mW. Due to the mismatch of the fiber core diameter between the S405-XP (core diameter = 3 µm) and the single mode fiber used in the fiber coupler (460-HP, core diameter = 3.5 µm), there was a significant insertion loss when connecting the two types of fiber. The output powers on the sample arm, after the microscope objective was 0.345 mW, and the output power on the reference arm was 0.304 mW.
Fig. 1. Schematic diagram of the blue-light OCM system. SLED – super-luminescent light emitting diode; FC – 50:50 fiber coupler; PC – polarization controller; C1 - C3 – collimator; L1 - L4 – achromatic lenses; VFL – variable focus liquid lens; X, Y – galvanometer scanning mirrors; M – mirror; DM – dielectric mirror; VPHG – volume phase holographic grating; OL – objective lens; GPU – graphics processing unit.
In the sample arm, a reflective collimator (RC04APC-P01, Thorlabs Inc., USA) collimated the beam to a variable focus liquid lens (clear aperture = 3.9 mm, Corning, USA) for autofocusing into different layers of the sample [29]. The 4 mm beam was then deflected by a pair of galvanometer scanners (Saturn 5B, Pangolin Laser Systems Inc., USA) to a 2× beam expander comprised of two pairs of visible achromatic doublet lenses, L1 and L2, that had effective focal lengths of 50 mm and 100 mm respectively (Fig. 1). A 1” broadband dielectric mirror (BB1-E02, Thorlabs Inc., USA) was placed in between the aforementioned relay telescope. Finally, a long working distance (34 mm) 10× magnification objective lens (378-803-3, Mitutoyo, Japan) was used to focus the beam onto the imaging plane.
OpticStudio (Zemax, LLC, USA) was used to model the blue-light OCM sample arm as shown in Fig. 2(a). It should be noted that a paraxial lens was used in lieu of a Zemax model of the objective lens, due to its unavailability, to evaluate the performance of the relay telescope. Spot diagrams at various scanning angles (1.1° × 1.1°, which translates to 750 µm × 750 µm) are shown in Fig. 2(b). All focus spots with 750 µm FOV were within the airy disk (radius of 1.361 µm), thus suggesting diffraction limited performance. The Zemax simulation was exported into SolidWorks to design mechanical mounts for the components at each specific location (Fig. 3(a)). These components were fitted on a 12” × 12” × 3/8” aluminum breadboard (MSB12, Thorlabs Inc., USA) as shown in Fig. 3(b). The sample was mounted on a vertical stage for precise alignment with the incident beam reaching the apex of the rabbit cornea shown in Fig. 3(c).
Fig. 2. (a) Ray-trace diagram from the blue-light OCM sample arm’s optical simulation in OpticStudio. (b) Spot diagrams with 750 µm × 750 µm FOV centered on the apex of the cornea model. Each spot diagram is encircled by an airy disk radius of 1.361 µm shown in black.
Fig. 3. (a) SolidWorks mechanical layout of the sample arm. (b) Photograph of 3D printed mounts with optical components. (c) Sample (rabbit cornea) mounted on an artificial anterior chamber on a vertical stage shown in a red dashed box. C2 – collimator; L2 - L3 – achromatic lenses; VFL – variable focus liquid lens; DB – driver board for VFL; X, Y – galvanometer scanning mirrors; M – mirror; DM – dielectric mirror; OL – objective lens.
The beams from the sample and reference arms travelled back to the spectrometer, which is depicted in Fig. 4. The spectrometer was constructed using a reflective collimator (RC12APC-P01, Thorlabs Inc., USA) which collimated the beam onto a volume phase holographic transmission grating (WP-1500/477-30, 1500 lp/mm, Wasatch Photonics, USA) that was angled incidentally at 20.9°. The grating dispersed the collimated beam into its wavelength components, which were then focused with a focusing lens (AC254-200-A, Thorlabs, USA) to a 2048-pixel line-scan camera (Octoplus, Teledyne e2v Ltd., UK). Although the camera in the blue-light OCM system had 2048 pixels, only 512 pixels were used due to the length constraints of the spectrometer arm, the limited bandwidth of the light source, and the spectral resolution provided by the grating. The camera operated at a line rate of 80 kHz with a 12-bit depth mode. A PCIe frame grabber (PCIe-1437, National Instrumental Corp., USA) was used to capture the recorded interferometric fringes. All OCT volumes presented in this manuscript have a dimension of 1000 A-scans per B-scan, and 1000 B-scans per volume. The imaging system’s triggering and scanning waveform generation were provided by a multifunctional data acquisition card (PCIe-6353, National Instrumental Corp., USA). Blue-light OCM images were acquired and processed by our custom written graphics processing unit (GPU) accelerated OCTViewer software package [29,30].
Fig. 4. (a) Zemax simulation of the spectrometer with a zoomed inset of the rays reaching the camera emphasized with the purple box. (b) SolidWorks mechanical layout of the spectrometer. (b) Photograph of 3D printed mounts with optical components shown in the yellow dashed box. C3 – collimator; VPHG – volume phase holographic grating; L4 – focusing lens, f = 200 mm.
3. Results
3.1 Signal processing
Standard spectral-domain OCT imaging processing procedures were applied to this system, including DC subtraction, wavenumber resampling, dispersion compensation, and Fourier transform. The interference signal detected on the line-scan camera was periodic with a phase-characteristic determined by the spectrometer design and dispersion mismatch between the reference and sample arms. We first performed the spectrometer calibration step to identify the relative relationship between the pixels in each A-scan in wavenumber space. We built a free-space interferometer that is free of dispersion mismatch for the purpose of calibrating spectrometer. Interference signals at various pathlength mismatch were recorded. The phase of the periodic signal was numerically retrieved using the Hilbert transform [31].
The axial mirror positions in Fig. 5 were measured with the micrometer gauge of a translational stage and therefore represent the point spread function (PSF) in air. It should be noted that the falloff does not have a maximum toward the zero position (z = 0 mm) due to the DC removal that rejects some of the lower frequencies close to zero hertz and detrending the noise floor. The falloff over an axial distance of 770 µm was 7 dB, which is larger than the theoretical value of around 4 dB over the full imaging depth [32]. However, a stronger falloff was expected due to imperfect collimation in the reference arm. The Gaussian fitted PSF was plotted as shown in Fig. 6 and the FWHM was evaluated to determine the axial resolution in air, which was measured to be 16 µm. This is reasonable compared to the theoretical axial resolution, which is 15 µm in air.
Fig. 5. (a) Intensity falloff of up to an axial depth of 1.2 mm (b) PSF measurement of FWHM of ∼16 µm in air (∼12 µm in tissue). (c) En face view of a standard USAF 1951 resolution test target imaged with blue-light OCM. (d) A zoomed image of the yellow dashed box in (c) depicting the estimation of lateral resolution for group 9 and element 3 with 1.6 µm line spacing.
Fig. 6. (a) Cross-sectional image of scotch tape to indicate image depth with the blue-light OCM system. Seven layers can be visualized with the aid of the variable focus liquid lens. (b) B-scan of a finger with visible sweat ducts. (c) Fungal filaments demonstrated on moldy bread. (d) E. coli bacteria in 1:10 glycerol solution (inset about 10× digitally zoomed in the orange dashed box).
The lateral resolution and FOV were evaluated using a positive high resolution United States Air Force 1951 resolution test target (64-862, Edmund Optics, USA). The line pairs of group 9 element 3 in Fig. 6 are discernible suggesting a lateral resolution measurement of 1.6 µm in free space. The theoretical lateral resolution was calculated to be 1.4 µm. The FOV was measured to be 750 µm × 750 µm.
3.2 Exploratory imaging of various structures
To further evaluate the system and inspect image depth, single non-averaged cross-sectional images of scotch tape and finger were collected [Fig. 6(a) and (b)]. With autofocusing adjustment of the variable focus liquid lens, at least seven layers of the tape can be visualized. Similarly, sweat ducts underneath the epidermis of the skin can be visualized with the blue-light OCM system. To investigate the possibility of imaging infectious pathogens known to cause microbial keratitis in humans, we also imaged filamentous fungi (demonstrated on moldy bread) as well as cultured Escherichia coli [Fig. 6(c) and (d)]. Common fungi that cause keratitis and moldy bread are Aspergillus and Fusarium [33]. E. coli (DH5a strain) were prepared in 1:10 glycerol solution and a single drop was placed on a microscope slide before imaging.
3.3 Ex vivo rabbit cornea
Five enucleated adult white rabbit eyes (Sierra for Medical Science, USA) were imaged with the blue-light OCM system. All eyes were shipped with intact eyelids that covered the cornea to prevent epithelial damage and imaged within 30 hours of harvesting. Corneal scleral buttons from the enucleated eyes were each placed on an artificial anterior chamber inflated by normal saline (0.9%w/v NaCl) within the normal corneal physiological intracameral pressure range. Corneas were examined under a slit-lamp microscope to verify the intact status of epithelium before imaging on the blue-light OCM system.
Representative en face blue-light OCM images of endothelial cells, collagen lamellae, and keratocytes of rabbit cornea are shown in Fig. 7 with a field of view of 750 µm × 750 µm. Clear contrast and lateral resolution can be observed with the blue-light OCM.
Fig. 7. En face images of rabbit corneal (a) endothelial cells, (b) collagen lamellae, and (c) keratocytes with 750 µm × 750 µm FOV. (Insets about 5× digitally zoomed).
We compared our blue-light OCM system with the Heidelberg HRTII-RCM, which is commonly used for clinical corneal imaging, presented in Fig. 8. We imaged corneal structures at a FOV of 450 µm × 450 µm at cellular resolution. Of note, the collagen lamellae demonstrated in Fig. 7(b) are not visible on the in vivo confocal microscopy (IVCM). Subjectively, the contrast and cell boundaries of the epithelium and endothelium are more readily apparent on blue-light OCM, whereas the keratocytes are better delineated with IVCM.
Fig. 8. Epithelium (a, d), endothelium (b, e) and keratocytes (c, f) in ex vivo rabbit cornea imaged with the blue-light OCM system (top row) and Heidelberg HRTII-RCM (bottom row).
Collagen lamellar density was also visualized with the blue-light OCM system in the anterior and posterior stroma of the rabbit cornea. A clear, striated pattern of collagen lamellae in the rabbit cornea was observed in different layers indicated by the B-scan images in Fig. 9. These collagen lamellae decrease in density as the variable focus liquid lens moves closer to the posterior cornea. We are not aware of any other clinical imaging device that is able to demonstrate the collagen lamellar architecture of the cornea, but other experimental studies demonstrate collagen lamellae in in vivo rat cornea [34].
Fig. 9. (a-c) En face images of collagen lamellae in parallel stripes in ex vivo rabbit cornea. (d-f) Representative B-scans of each en face to emphasize the anterior and posterior sections of the fibers shown in green dashed boxes. The density of the fibers decreases as imaging depth increases.
4. Discussion
In this study, a novel spectral-domain OCM system was built with a 450 nm super-luminescent light emitting diode (SLED) light source. To the best of our knowledge, this is the first time that a spectral domain OCM system in the short wavelength range was demonstrated. With our custom designed blue-light OCM system, we demonstrated en face OCT images with cellular lateral resolution to visualize ex vivo rabbit corneal epithelial and endothelial cells, keratocytes, and collagen lamellae in the anterior and posterior stroma with up to 750 µm × 750 µm FOV. These volumes were recorded in 12.5 s with a spatial resolution of 16 µm axially and 1.6 µm laterally in air. We also demonstrated the ability of this modality to clearly delineate filamentous fungi and bacteria, both of which are known to cause infectious keratitis in humans. Although the current standard of care in vivo cellular imaging microscopy (IVCM) has a higher resolution than the described system, the blue-light OCM can acquire actual 3D image volumes at a higher speed and detect structure sizes close to IVCM resolution. The blue-light OCM might provide a pathway for image-based diagnosis of two of the most common causes of corneal blindness globally: keratoconus (characterized by disorganization of the collagen lamellae) and infectious keratitis (characterized by stromal invasion of microbial pathogens, most commonly bacteria) [35,2]. However, further investigations are needed. The current optical properties show blue-light OCM has more speckle noise compared to the confocal microscopy images. However, speckle noise can be reduced with physical and computational methods [36–45].
Although it is possible to construct a blue-light OCM system with a supercontinuum light source, it can cost up to ten times more than the blue SLED used in this project. Moreover, a supercontinuum laser is inherently unstable in terms of the output power and spectrum shape. Hence, a visible light OCT with supercontinuum usually suffers from high relative intensity noise. Despite the relatively narrow bandwidth (6 nm), the center wavelength of 450 nm provides an axial resolution of 15 µm in air which is comparable to a Fourier Domain OCT system with a 1310 nm center wavelength and 50 nm bandwidth, a common wavelength of choice in anterior chamber imaging. However, compared to an experimental OCM system using 800 nm wavelength achieving 1.5 µm isotropic resolution [4], our blue light OCM is indeed at a disadvantage. Although sacrificing some penetration depth by using blue light for OCM for weakly scattering samples, we were able to achieve cellular level transverse resolution with only a 10× objective lens. For the described system, a lens from Mitutoyo was used which considerably reduces costs compared to lenses with higher magnifications. The moderate power of the objective lens also enables long working distance and non-contact imaging. Another advantage of using shorter wavelengths is that with the same spot size, shorter wavelength beam has a longer depth of focus. For example, the depth of focus in our blue-light OCM system was 7.2 µm, whereas a 1310 nm OCM system with similar spot size would have a depth of focus of 2.5 µm.
One limitation of our current blue-light OCM system is the relatively low sensitivity compared to the OCM systems with NIR light source and the potential phototoxicity in in vivo imaging. The theoretical sensitivity of our imaging system is calculated to be 91 dB. Because it is difficult to account for optical losses reliably, we expected a lower measured sensitivity that was obtained with a value of 80 dB. Candidates reducing the sensitivity are losses on lenses and the line-scan camera’s low quantum efficiency (∼15%). Future systems can use better anti-reflection coatings, and a more efficient camera can be selected. Moreover, a custom-made fiber coupler with S405-XP fiber could significantly reduce the insertion loss and allow us to optimize the splitting ratio between the reference and sample arms while maintaining laser safety. Improving the sensitivity would enable faster imaging speed to the point that it is feasible for in vivo human corneal imaging. Future studies could use a longer wavelength SLED (510 nm) with a broader bandwidth (10 nm), which is considered safer to use by the American National Standard Institute (ANSI) Z136.1-2014 standards in ocular imaging than blue light for in vivo imaging of human cornea [46].
Future studies using a short-wavelength range for OCM will focus on imaging such as 3D mapping of corneal collagen structures or assessing human corneas as well as other species to avioid sample bias. Literature shows the significance of analyzing aspects, such as the lamellar orientation in the corneal collagen, measuring the biomechanical properties, or assessing the microstructure during and before transplantation on human eyes. Many reported methods for imaging corneal structures include polarization-sensitive setups, second harmonic generation, or two-photon autofluorescence, which are expensive and inherently complex to build [14,47–51]. Consequently, this study provides an initial step describing the advantages of using OCM imaging systems with short-wavelengths that can reduce costs of systems, reduce complexity, and make its applications more accessible to clinical environments thus helping more people in need.
Funding
National Institutes of Health (P30 EY010572, R01 EY028755, R01HD107494); Oregon Health and Science University (unrestricted departmental grant); Research to Prevent Blindness (Career Advancement Award).
Acknowledgments
Along with all the contributors who helped with this work, we would like to thank Dr. Kate Keller for her time and effort to provide us with many samples of E. coli.
Disclosures
David Huang: Optovue Inc. (F, I, P, R). These potential conflicts of interest have been reviewed and managed by OHSU. Other authors declare no relevant conflicts of interest related to this article.
Data availability
Data underlying the results presented in this paper are not publicly available at this time but may be obtained from the authors upon reasonable request.
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