1. Introduction

Optical coherence tomography (OCT) is a non-invasive imaging technique that is being widely applied in clinical applications to provide high-resolution, cross-sectional imaging of internal microstructure of materials and biological tissue morphology by computing the echo time delay and backscattered light intensity. OCT is a non-ionized, cost-effective, and real-time imaging modality whose resolution is 1-15 µm and despite optical attenuation from tissue scattering and absorption, it provides depth images up to a 2–3 mm range in most tissues [1]. Because of these qualities, OCT is widely used for numerous applications, including ophthalmology [2,3], dentistry [4,5], dermatology [6,7], tissue imaging [8,9], agriculture [10,11], and industrial application [12,13]. OCT faces a tradeoff between lateral resolution and depth of imaging and these two factors decide what kind of features can be imaged [14]. In a conventional OCT system, low numerical aperture (NA) objectives are used to get a large depth of imaging, which limits the lateral resolution of the system [15].

Optical coherence microscopy (OCM) is the integration of OCT and confocal detection, which can achieve high lateral resolution [16]. OCM has a persuasive field-of-view (FOV) of several millimeters, 3D visualization capability of deep tissue in a few micrometer resolutions, and a feasible temporal resolution that allows for real-time measurement [17]. To provide cellular imaging at the cost of a shallower depth of focus (100–200 µm), OCM employs a higher NA objective (i.e., ∼0.2) than conventional OCT (i.e., ∼0.04) [13]. In OCT and OCM, the light source has a direct impact on transverse and axial resolutions, as well as imaging depth, because the transverse resolution is linearly proportional to the central wavelength of the source, whereas the axial resolution is inversely proportional to the bandwidth, and longer wavelengths are more effective at penetrating deeper into tissue [14]. Light sources with a 1300 nm spectral band are commonly used to achieve greater imaging depth in scattering tissue because the light scattering is lower around 1300 nm than in the conventional near-infrared (NIR) band of 650 to 900 nm [18]. OCM with a 1700nm spectral band for high-resolution label-free deep tissue imaging has also been reported [19]. In full-field OCM (FF-OCM), a parallel acquisition scheme has been employed to acquire only the high-resolution in-focus en-face images [20–23]. Despite having a faster acquisition capability for parallel scanning beams, the FF-OCM is more vulnerable to cross-talk and multiple scattering than the point-scanning technique [24–26].

As an aspect of improving the depth of imaging, numerous approaches have been demonstrated besides using light sources with a high spectral band. In the literature, several dynamic focusing imaging methods with and without moving components have been proposed. One solution uses the mechanical motion of the sample arm to scan the sample in the depth direction with a defined depth, and only the defined portion of data volumes are captured at each focus depth, then combined to form a complete volume [27], but this approach is limited by the large data size and long acquisition time. Moreover, the path length mismatch caused by moving the focusing lens to obtain precisely focused images must be corrected by an extra hardware adjustment. Dynamic focusing involving non-moving parts such as liquid-lens-based approaches has been demonstrated to get invariant lateral resolution throughout a large cubic millimeter imaging sample [28–31]. However, liquid lens technology has some limitations, including a complex operation for tuning the focus, difficulties in producing big aperture lenses, and sensitivity to vibration [32]. The deformable microelectromechanical system (MEMS) mirror technique has also been used as a non-moving component approach for dynamic focus control [33,34]. Furthermore, to get extended imaging depth, the Bessel beam has been employed in several studies for its extended DOF [35–37], and cellular resolutions have been achieved [38–40]. However, unlike Gaussian beams, the use of the Bessel beam approach is constrained by nonideal side lobes in the point spread function and lower contrast in the focal plane [41]. Therefore, a non-invasive optical imaging system is required with variable DOF to obtain precisely focused high-resolution images.

The fundamental scope of this study was to develop a variable DOF-OCM system to vary and improve the DOF for obtaining high-resolution and precise focused images of the sample. The beam diameter of the sample arm was increased and decreased to vary the DOF of the system using a deviated focusing in the optical pathway before the galvo scanners. The variable DOF imaging capability and the accurate focusing over the varying DOF were confirmed by analyzing the OCM cross-sectional images of lemon and scotch tape samples, respectively. Also, the high-resolution volumetric imaging capability of this system was verified through en face images of apple and grape specimens in healthy and infected states. Thus, the proposed optically deviated focusing method applied in the sample arm setup can be used to vary DOF for obtaining high-resolution and precisely focused images of the sample.

2. Materials and methods

2.1 Variable DOF-OCM System configuration

A schematic diagram of the developed variable DOF-OCM system is shown in Fig. 1. A broadband light source (BroadLighters T-850-HP, Superlum) with a central wavelength of 860 nm and a full width at half maximum bandwidth of 165 nm was used in the setup. To divide the optical power equally between the sample and reference arms, a 50:50 fiber coupler (TW850R5A2, Thorlabs) was used. Polarization controllers (FPC023, Thorlabs, Newton, NJ, USA) were used in both reference and sample arms to control the polarized state of the transmitted light. A fiber adapter was used at the fiber coupler end, which was connected to the sample arm. A deviating lens (C330TME-B, f = 3.1 mm, Thorlabs, USA) was placed at its back focal distance after the fiber adapter in the optical pathway of the sample arm. The deviating lens can be positioned toward and away from the fiber adapter using a motorized based autofocus module (PT-AF700; Piezoelectric Technology Co. Ltd., Korea) to vary the beam diameter. The samples were scanned using 2D galvo scanners (GVS002, Thorlabs) in the sample arm. A two-lens relay configuration was used after the galvo mirrors, where light passes through a lens of 30 mm focal length followed by a lens with 50 mm focal length. An objective lens with a high numerical aperture (OLYMPUS, Plan 20x/ NA 0.4) was used to obtain OCM images with high lateral resolution. The same relay configuration and objective lens were used in the reference arm to avoid dispersion. The interference signal obtained from the interferometer was transferred to a customized spectrometer containing a collimator, diffraction grating, achromatic doublet lens, and a 12-bit complementary metal-oxide-semiconductor (CMOS) line-scanning camera (Aviiva EM4, E2V, USA). When the camera was set at an exposure time of 14.1 µs and the power at the sample path was 1 mW, the sensitivity of the developed system was approximately 110 dB near zero optical delay. The lateral resolution of this system was measured at 2.19 µm using a resolution target (USAF 1951, Edmund Optics, Barrington, NJ, USA), and the axial resolution of this system was 1.97 µm, determined by the central wavelength and spectral bandwidth of the light source.

 

Fig. 1. Schematic of variable depth-of-focus OCM system. BLS: broadband light source; C: collimator; DG: diffraction grating; FC: fiber coupler; FA: fiber adapter; GM: galvo mirror; L: lens; LSC: line scan camera; M: mirror; MS: micro stage; OL: objective lens; PC: polarization controller; S: sample; DL: deviating lens.

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2.2 Principle of varying DOF and increasing lateral resolution

Beam propagation through the fiber adapter and the deviating lens is shown in Fig. 1 with a blue dotted box, where it is visualized that the beam diameter can be increased and decreased by moving the deviating lens toward and away from the fiber adapter, respectively. In the first case, the initial positions (1) of the fiber adapter and the deviating lens are shown, where the fiber adapter is located at the back focal point of the deviating lens. In the second case (2), it is visualized that when the deviating lens moved towards the fiber adapter ((-) direction), the beam diameter increased. In the third case (3), when the deviating lens moved away from the fiber adapter ((+) direction), the beam diameter decreased. In the system schematic, orange, green, and blue dotted lines (for the first, second, and third cases, respectively) represent the variable DOF-acquiring process through the use of a deviating lens. The use of a fiber adapter-deviating lens combination and a relay configuration helps to very the DOF in high-resolution imaging.

3. Result

3.1 Experimentally confirmed lateral resolution

The experimentally confirmed lateral resolution of the variable DOF-OCM system is shown in Fig. 2. The OCM en face image of USAF-1951 1x resolution target, shown in Fig. 2(a), was obtained using OLYMPUS, Plan 20x/ NA 0.4 objective lens in the sample arm. Figure 2(b) shows the A-scan depth profile result of resolution target images of 2(a). The region of interest (ROI) of Fig. 2(b) was group 7 of the resolution target image, which is marked by a red dotted line in Fig. 2(a). In the A-scan profile, it is visualized that the intensity peaks from element 6 of group 7 are very much distinguishable. According to the USAF-1951 1x resolution target chart, shown in Fig. 2(c), the calculated lateral resolution of the variable DOF-OCM system was 2.19 µm using OLYMPUS, Plan 20x/ NA 0.4 objective lens. Figure 2(d) shows an OCM en face image of the USAF-1951 1x resolution target, which was obtained using the OLYMPUS Plan 40x/NA 0.65 objective lens in the sample arm. Figure 2(e) is the extended view of groups 8 and 9 of the resolution target image of 2(d), where the red dotted line is the scan position for obtaining an A-scan intensity profile. The A-scan profile of group 8 of the resolution target image, shown in Fig. 2(f), visualizes the distinguishable intensity peaks of element 6 of group 8, and the computed lateral resolution of the variable DOF-OCM system employing an OLYMPUS Plan 40x/ NA 0.65 objective lens was 1.10 µm, according to the resolution chart shown in Fig. 2(c). The experimentally confirmed lateral resolution of 2.19 µm using OLYMPUS, Plan 20x/ NA 0.4 objective lens was sufficient for further cross-sectional imaging to assess the variable DOF imaging approach.

 

Fig. 2. Resolution target images and their A-scan profiles for measuring lateral resolution of variable DOF-OCM system. (a) variable DOF-OCM resolution target image taken using OLYMPUS, Plan 20X/ NA 0.4 objective lens. (b) A-scan profile of OCM image shown in image (a). (c) lateral resolution chart of groups 6, 7, 8, and 9 of USAF 1951 1x resolution target. (d) variable DOF-OCM resolution target image taken using OLYMPUS, Plan 40X/ NA 0.65 objective lens. (e) extended view of groups 8 and 9 from image (d). (f) A-scan profile of OCM image shown in image (e).

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3.2 Accurate focusing capacity evaluation

Figure 3 represents the accurate focusing capacity of the variable DOF-OCM system based on the movement of the deviating lens away from the fiber adapter. Figure 3(a-e) shows the 2D OCM images of a scotch tape sample, which were taken by moving the deviating lens at 0, 0.5, 1, 1.5, and 2 mm distances from its original position, respectively. Figure 3(a) was taken at the initial position of the deviating lens and then it was gradually moved away from the fiber adapter and the remaining images (shown in Fig. 3(b-e)) were taken. The depth intensity profiles of scotch tape images, which were taken at different positions of the deviating lens are shown in Fig. 3(f), and the ROI is marked with a blue dotted box in Fig. 3(a). The depth intensity profile was acquired by combining and averaging 200 A-lines from the given ROI using a MATLAB (Mathworks, USA)-coded program. Only one OCT image from each position of the deviating lens was used to make that intensity curve for each position. The black, blue, green, magenta, and red depth intensity profiles represent the depth profiles of OCM images, which were taken at 0, 0.5, 1, 1.5, and 2 mm positions, respectively. In the depth intensity profile, the scotch tape layers are distinguishable till the last layer. More crucially, the backscattered intensity of OCM cross-sectional images gradually increases with the movement of the deviating lens from 0 to 2 mm distance from its original position away from the fiber adapter, as seen in the depth intensity profile. Figure 3(g) shows the curve-fitted depth intensity profile of OCM images to present a significant difference in the back-scattered intensity of obtained OCM images. Curve fitting is a technique for determining a mathematical model that fits the experimental data and smoothing the outcome. In the curve fitted result, it is visualized that the backscattered intensity increases from the focused region of OCM cross-sectional images when the deviating lens is gradually moved from 0 to 2 mm distance from its original position away from the fiber adapter. The rise of backscattered intensity from cross-sectional images due to the shifting of the deviating lens away from the fiber adapter approves the proper focusing through the variable DOF imaging approach.

 

Fig. 3. Intensity measurement of variable-DOF-OCM system acquired 2D cross-sectional image. (a-e) scotch tape 2D cross-sectional images. (f) A-scan profiles of 2D cross-section images, shown in (a-e). (g) curve fitted A-scan profiles of 2D cross-section images, shown in (a-e).

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3.3 Verification of variable DOF imaging approach

A sliced lemon specimen was used for cross-sectional imaging to evaluate the variable DOF imaging approach. Figure 4(a-e) shows 2D cross-sectional images of sliced lemon obtained by moving the deviating lens at 0, 0.5, 1, 1.5, and 2 mm distances from its original position away from the fiber adapter, respectively. In OCM cross-section images, it is visualized that the DOF increases with the gradual movement of the deviating lens from 0 to 2 mm from its initial position, marked by the red dotted box and double arrow in Fig. 4(a-e). A-scan analysis was performed on cross-sectional images to get quantitative data to verify the variable DOF imaging approach of the developed system. The ROI on which A-scan analysis was performed is marked by the green box in Fig. 4(a). 450 consecutive A-lines were summed up and averaged to get the DOF region data. Figure 4(f) depicts the graphical representation of the change in DOF in the 2D cross-sectional images as a function of the varied lens position, where it is seen that the DOF increases with the gradual movement of the deviating lens from 0 to 2 mm from its initial position. The mean intensity of 2D cross-sectional images was measured from the same ROI to further confirm the variable DOF technique in cellular-level imaging because image intensity rises proportionally with DOF. Figure 4(g) shows a graphical representation of the change in mean intensity of 2D cross-sectional images as a function of deviating lens position, showing that the mean intensity rises as the deviating lens moves from 0 to 2 mm from its original location. En face images of a sliced lemon are shown in Fig. 4(h), which were acquired at depth direction intervals of 70 µm from a 3D volumetric record when the deviating lens was 1.5 mm distant from the fiber adapter point. In en face images, the inner morphology of a lemon can be seen with the intercellular water, intercellular space, intracellular water, and cell wall. The variable DOF imaging method of this system was approved by the rise in mean intensity of cross-sectional OCM images obtained by gradually shifting the deviating lens away from the fiber adapter.

 

Fig. 4. Evaluation of DOF variation in 2D cross-sectional imaging. (a-e) lemon 2D cross-sectional images. (f) graphical representation of the change of DOF of lemon 2D cross-sectional images according to the position of the deviating lens. (g) graphical representation of the change of mean intensity of lemon 2D cross-sectional images according to the position of the deviating lens. (h) en face images of a sliced lemon at 70, 140, 210, 280, 350, and 420 µm depths.

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3.4 Application of high-resolution imaging of variable DOF-OCM system

Figure 5(a) shows eight OCM en face images of a healthy apple specimen, which were taken at 70 µm intervals at depth direction when the deviating lens was 1.5 mm distant from the fiber adapter point. In Fig. 5(a), intercellular water, intercellular space, intracellular water, and cell wall are marked with arrows. In addition, the intensity of en face images appears to be reduced at the depth direction (70 µm to 560 µm). The high backscattered intensity was detected in the fleshy space and nearly zero backscattered intensity was detected in the intracellular water region in en face images of healthy apple specimens. Similarly, the en face images of abnormal/infected apple specimens are shown in Fig. 5(b), which were acquired at depth direction intervals of 70 µm from a 3D volumetric record when the deviating lens was 1.5 mm distant from the fiber adapter point. In the depth direction (70 µm to 560 µm), the intensity of en face images of diseased apple specimens appears to be decreased at depth direction. More crucially, the en face images of infected apple specimens are visualized with the shrink cell membrane, cell wall, and intracellular space due to the loss of intracellular and intercellular water, and the high intensity was detected in the fleshy shrink cell region compared to the water region.

 

Fig. 5. Variable DOF-OCM system acquired en face images of healthy and infected apple specimens. (a) en face images of healthy apple specimens at 70, 140, 210, 280, 350, 420, 490, and 560 µm depths. (b) en face images of abnormal apple specimens at 70, 140, 210, 280, 350, 420, 490, and 560 µm depths.

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Figure 6(a) shows six OCM en face images of a healthy grape specimen, which were taken at 70 µm intervals at depth directions, where intercellular water, intercellular space, intracellular water, and cell wall are marked with arrows. The strong backscattered intensity was observed in the fleshy intercellular space, and cell wall region, while nearly zero backscattered intensity was seen in the intracellular water area. Figure 6(b) shows en face images of an infected grape specimen taken at 70 µm intervals at depth directions, demonstrating how the cell wall and intracellular space shrank as intracellular and intercellular water was lost. The intensity of en face images of diseased grape specimens appears to be decreased in the depth direction (70 m to 560 m). Because of the strong backscattered recorded from only the fleshy parts, the fleshy intercellular space and cell wall areas seem brighter than other regions. The en face images in the depth direction reveal the interior morphological structure of apple and grape specimens, also confirming the volumetric imaging capacity of the variable DOF-OCM with a DOF of a few hundred micrometers. The en face images of both healthy and infected grape specimens were taken from a 3D volumetric record when the deviating lens was 1.5 mm distant from the fiber adapter point.

 

Fig. 6. Variable DOF-OCM acquired en face images of healthy and infected grape specimens. (a) en face images of healthy grape specimens at 70, 140, 210, 280, 350, and 420 µm depths. (b) en face images of abnormal grape specimens at 70, 140, 210, 280, 350, and 420 µm depths.

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4. Discussion

OCM is based on the same principle as OCT but uses an objective lens with high NA to increase the lateral resolution, which limited the DOF in optical imaging. Different approaches have been demonstrated to overcome the trade-off between lateral resolution and DOF of OCM, but those approaches are limited either by cross-talk and multiple scattering [20] or long data size and acquisition time or nonideal side lobes in point spread function and lower contrast in the focal plane [41] or low SNR [42]. In this study, firstly, OLYMPUS Plan 40X/ NA 0.65 objective lens was employed for high-resolution imaging, but the limited DOF of this lens disrupts the main purpose of varying DOF of this study. Later, OLYMPUS Plan 20X/ NA 0.4 was used to get comparatively better DOF and the high-resolution imaging capability of this study was verified by assessing the resolution target images. For the variable DOF imaging approach, a combination of a motorization-based deviating lens and a static fiber adapter was used in the sample arm to vary the beam diameter for obtaining OCM images with precisely focused and variable DOF. Although the lateral resolution of this system was affected by changes in beam diameter due to the inverse proportional relationship between beam diameter and lateral resolution [43], the resolution was sufficient for cellular-level imaging.

Usually, the relay configuration is used to expand the beam diameter and to get a fixed focus on the sample. However, the deviating lens has been employed in this study to achieve a varying depth of focus and precise focusing capability of the OCM system by changing the beam diameter. Also, the deviating lens can be easily adjustable through the micro-stage. In contrast, the focal lengths of the relay-configured lenses are usually matched with the galvo scanner and objective lens, thus, the image quality can be degraded by changing the position of the lens. However, to compensate for the dispersion mismatch between the sample and reference arms, a software-based dispersion compensation method was used in this study. Some blurring effects occurred due to the objective lens focusing deviation, although it was negligible.

5. Conclusion

In this study, a variable DOF-OCM imaging approach was successfully demonstrated for precisely focused high-resolution images utilizing the optically deviated focusing method. A fiber adapter and a deviating lens were used in the sample arm to vary the beam diameter in the optical pathway before the galvo scanner for proper focusing and to extend the DOF. The capacity of precise focusing and varying DOF imaging using the variable DOF-OCM imaging approach was verified by assessing the cross-sectional images of scotch tape and lemon specimens, shown in Figs. 3 and 4, respectively. Also, the high-resolution volumetric imaging capability of this variable DOF-OCM system at the cellular level was demonstrated through the assessment of en face images of apple and grape specimens in both healthy and infected states. Thus, the system can be a proper alternative to the conventional OCM systems to ensure precisely focused high-resolution images with varying DOF of a few hundred micrometers.