We use our previously developed adaptive optics (AO) scanning laser ophthalmoscope (SLO)/ optical coherence tomography (OCT) instrument to investigate its capability for imaging retinal vasculature. The system records SLO and OCT images simultaneously with a pixel to pixel correspondence which allows a direct comparison between those imaging modalities. Different field of views ranging from 0.8°x0.8° up to 4°x4° are supported by the instrument. In addition a dynamic focus scheme was developed for the AO-SLO/OCT system in order to maintain the high transverse resolution throughout imaging depth. The active axial eye tracking that is implemented in the OCT channel allows time resolved measurements of the retinal vasculature in the en-face imaging plane. Vessel walls and structures that we believe correspond to individual erythrocytes could be visualized with the system.
© 2015 Optical Society of America
Retinal vasculature plays an important role for the metabolism of the retina. Diseases such as diabetes affect the normal vasculature which may lead to vision loss and finally blindness. Currently fluorescein angiography is regarded as clinical gold standard for assessing the retinal vasculature in vivo. However, this requires the injection of a dye in order to increase vessel contrast which is invasive and may lead to unwanted complications in some patients. Optical coherence tomography (OCT) provides depth resolved information and using intrinsic motion within a vessel as contrast mechanism retinal vasculature can be visualized [1–4]. Even though small capillaries can be imaged with this technique the transverse resolution remains limited which prevents the visualization of smaller structures such as vessel walls or individual erythrocytes. In order to increase the transverse resolution adaptive optics (AO), a technique known from astronomy has to be used. Details of the technology and its importance for ophthalmic imaging can be found in previous work [5–13]. AO has been combined with a variety of different retinal imaging techniques such as fundus photography (FP) [14–17], scanning laser ophthalmoscopy (SLO) [18–25], or optical coherence tomography (OCT) [26–33]. However, using AO fundus photography or AO-SLO the contrast obtained from vessels within these images is quite limited. This is mainly caused by a strong contribution from the highly reflective nerve fiber layer or the top vessel wall to the detected signal intensity. In order to overcome this limitation, the use of forward scattered light for imaging  or the combination with fluorescence imaging has been proposed . Although these methods greatly improve vessel contrast, the depth resolution of AO-FP and AO-SLO is not comparable to what can be achieved with OCT. Therefore details that may be contained only within a plane with limited depth extension might be obscured by highly scattering structures outside the plane of interest. The combination of AO with OCT might overcome this limitation. However standard AO-OCT systems that are based on recording A-scans have the drawback that the en-face imaging plane which shows the vessel structure has to be extracted from a 3D data set [33, 36–38]. Currently this data is recorded rather slowly compared to AO-SLO or AO-FP which prevents the observation of fast dynamic processes such as blood flow in this imaging plane. En-face OCT [29, 39–43] uses a scanning protocol that is identical to AO-SLO and might therefore be an alternative approach for investigating the vessel structure with AO-OCT. Recently, we introduced an en-face OCT/SLO instrument that is equipped with AO and an axial eye tracker which yielded high resolution images of the photoreceptor mosaic .
In this paper we present an adapted version of our instrument and demonstrate its capabilities for investigating retinal vasculature. A new dynamic focus scheme was implemented that allows for recording of 3D OCT data with high transverse resolution throughout imaging depth. Since the instrument records both, SLO and OCT images simultaneously, a direct comparison between the two imaging modalities becomes possible. Using the capability of the system to record en-face OCT images with a high frame rate at a certain depth within the retina, dynamic processes such as blood flow are investigated. The high axial and transverse resolution allows for visualization of small structures such as vessel walls and individual blood cells.
The used instrument is a lens based AO-SLO  in combination with en-face OCT. Details of the system have been presented previously . In brief SLO and OCT images are recorded simultaneously at a frame rate of 20 or 40 frames per second. The light source is a superluminescent diode with a center wavelength of 840nm and a bandwidth (full width at half maximum) of 50nm which yields an axial resolution of ~4.5µm in retinal tissue. The light that is backscattered from the retina is split into several parts and is used for wavefront sensing, the SLO channel and the OCT channel, respectively. In order to control the depth location of the imaging plane of the en-face OCT a high speed axial eye tracker is used which keeps the residual tracking error below the depth resolution of the system. AO correction is performed in closed loop using a 52 element deformable mirror (DM), a Shack Hartmann wavefront sensor and commercially available AO software (Casao, Imagine Eyes, France). The theoretical transverse resolution of the system within the eye is ~2.2µm (assuming a pupil diameter of 8mm). However, the actual resolution will be smaller and depends on the pupil diameter of the subject. In order to maintain the high transverse resolution throughout imaging depth, a dynamic focusing scheme was implemented into the instrument. For this purpose the DM was used to add additional defocus to the system, thus shifting the focal plane through the retinal tissue. The DM provides sufficient stroke in order to cover the entire retinal depth. One key issue for the dynamic focus is the synchronization between the shift of the coherence plane and the focal plane. To achieve this goal corresponding software had to be developed. Data acquisition and AO control is performed on two separate computers which required corresponding communications between the two. In order to determine the correct speed and starting point for the defocus change of the DM during data acquisition which should be matched to the speed of the coherence gate that is moved in depth, measurements in a model eye (consisting of a 30mm lens and a resolution test target (RTT)) were performed. Thereby the functionality of the dynamic focus was tested by recording OCT volumes with and without dynamic focus of the RTT that was intentionally placed in-focus and out of focus (in both directions). For this test we used the same procedure as we have published previously with a different dynamic focus scheme .
In vivo measurements were performed on healthy volunteers after informed consent from the subject was obtained. All measurements were approved by the local ethics committee and adhered to the tenets of the Declaration of Helsinki. Imaging was performed without an artificially dilated pupil in subjects that had a sufficiently large pupil (>6.5mm) during the experiment in the dark measurement environment. No drugs were administered to prevent accommodation. Different scanning angles between 0.8x0.8 degree and 4x4 degree at a frame rate of 20 or 40 frames per second were used. Each frame consists of 1152 (x) times 790 (y) pixels. Both scanning directions (forward and backward scan) of the 8 kHz resonant scanner were used. The scanning depth (z) for the 3D OCT volumes was set to 600μm (optical) while for measurements of dynamic changes the scanning depth was kept at the same axial position. Data acquisition took 6 sec resulting in a total of 120 or 240 recorded frames. The signal in both channels (SLO and OCT) was sampled with the same data acquisition board operating at 20M samples per second. Before starting the measurements, aberrations introduced by imperfections of the eye were corrected by closing the AO-loop. Data acquisition was started after the residual wavefront error was below 0.1µm. During a volume scan with dynamic focus the AO loop was opened and the shape of the DM was changed to adapt the defocus according to the depth scanning speed of the OCT channel. The opening of the loop is essential in order to avoid an influence of the AO control to the dynamic focus scheme.
Image post processing included the correction for sinusoidal motion of the resonant scanner and correction of transverse motion artifacts. For the in-plane data a reference frame (that showed minimal or no motion artifacts) in the SLO data set was manually chosen and all other SLO frames were registered to this reference frame using a cross correlation algorithm. The same correction was then applied to the OCT frames. In the case of 3D volumes this procedure cannot be used because the individual frames will show different structures. Therefore each frame (with exception of the first frame) is registered to the previous frame of the data set.
An exemplary volume scan of the retina recorded in vivo with dynamic focus is shown in Fig. 1. The fly through movie (Media 1) starts at a depth location corresponding to the choroid. The video shows then the recorded en-face images while the coherence gate and focus plane are moved simultaneously towards anterior layers. Note that the data was recorded in the same order as is shown in the video. On the left hand side of the movie en-face OCT image frames can be seen while on the right hand side the corresponding SLO images are displayed. The OCT frames allow for a clear separation between individual retinal layers. It is worth to note that when the coherence plane is located within the large vessel, individual erythrocytes can be observed. Since the scanning speed of our system is limited, the visibility of the erythrocytes depends on the flow velocity within the vessel which is varying from frame to frame because of the cardiac cycle. The best visibility of these cells in this rather large vessel is obtained when the flow speed is low. This is the case for example in frame No. 85 of the data set. In the SLO images a clear separation between photoreceptor layer and retinal nerve fiber layer (including the top vessel wall) can be achieved. However, the strong signal arising from these layers outshines contributions from layers in between. Media 2 shows the corresponding B-scan images (y-z- imaging plane, the B-sans correspond to vertical cross sections in Fig. 1) extracted from the 3D data set. Since OCT and SLO images are recorded simultaneously, the depth resolution or depth of focus of the SLO can be determined experimentally. Thereby, SLO-B-scans are extracted from the entire data set and compared with the OCT B-scans. If we take the wave guiding properties of cones into account we can assume that the signal from the photoreceptor layer in the SLO emerges from a single plane (i.e. the external limiting membrane). This plane is broadened in the SLO B-scans by the corresponding depth of focus of the system. We can then use the full width at half maximum intensity distribution in depth of the SLO signal from the photoreceptor layer in order to determine the confocal gating. For the data set shown in Fig. 1 we estimated a depth of focus of ~50µm in tissue.
Figure 2 shows averaged OCT and SLO images of the same location on the retina while the focus and coherence plane are kept at a constant depth within the retina corresponding to the center of the large vessel. In the OCT image the full extension of the vessel can be observed. Due to motion and the high backscattering potential of the particles within the vessel (mainly erythrocytes) the interior of the vessel appears in the averaged images as bright and rather homogenous structure. Clearly visible are structures along the vessels (indicated by white parenthesis) that we believe correspond to the walls of the large vessel. From the image we estimated a 7µm thickness of the vessel wall which is in good agreement with results obtained previously . In the lower right corner of the image, shortly after the bifurcation, the vessel wall appears to be thickened. The confocal SLO image shows poor contrast of the internal vessel structure because the signal is mainly dominated by highly backscattering structures such as the retinal nerve fiber layer and the top vessel wall (the high backscattering of this location can also be observed in the OCT data set shown in Fig. 1). In order to increase the vessel contrast in the SLO images an offset pinhole configuration or a split detector scheme has to be used [34, 47].
The entire recorded data set can be viewed in Media 3. In the movie small moving particles which we think correspond to individual erythrocytes can be observed in the OCT images. Although some motion within the vessel can be seen in the SLO images, the strong backscattering signal arising from the top vessel wall and the retinal nerve fiber layer makes such observation much more difficult.
When viewing the movie it can be recognized that in some image frames the particles appear elongated (elliptical) parallel to the vessel while in other image frames they have an almost round shape. One possible explanation could be the varying flow velocity of the particles due to the cardiac cycle. However, further investigations are needed in order to confirm this interpretation. Figure 3 shows an example frame where this elongation can be seen (indicated by the white circle). This elongation allows a determination of the flow direction within the vessel because such elongations will only appear if the flow vector of the particles has a component that is parallel to the slow scanning direction. In our case the scanning of the slow axis starts at the bottom of the image which means that the flow direction is roughly from right to left (indicated by the red arrow in Fig. 3). This also indicates that the observed vessel is an artery because the particles move from the main vessel to the small branches. It is interesting to note that the flow within the lower branch is approximately anti parallel to the scanning direction. Therefore the particles will appear compressed in the flow direction (indicated by the red circle in Fig. 3). In principle it should be possible to determine via the elongation the flow velocity. However, individual erythrocytes are of plate-like shape with a height that is smaller than the diameter of the cell. Therefore the shape of individual cells within the images may vary depending on the cell orientation in respect to the imaging beam. This effect cannot be separated from shape distortions introduced by flow. In order to overcome this limitation many particles have to be investigated as can be done by switching into a different scanning mode that allows rapidly imaging of the same AO-SLO line over time and measuring the slope of the corresponding time resolved image [48, 49]
In a next step we recorded images with a larger field of view. Figure 4 shows an overview fundus image as well as en-face OCT and SLO images with a field of view of ~4°x4° and the focus and coherence plane kept constant. Even without additional contrast mechanism, the capillary network can be observed with high contrast in the OCT image (cf. Figure 4b). However, the interpretation of the en-face image deserves some care because the coherence plane is not parallel to the retina (retinal pigment epithelium). Therefore different layers corresponding to different depths within the retina are visible in the image. In the upper right corner individual nerve fiber bundles can be observed while in the lower left corner the capillary network of the inner plexiform layer can be observed. In the SLO image residual backscattering signal from the photoreceptor layer can be observed which leads to the typical appearance (bright central core that is surrounded by a dark area on both sides) of the large vessel in confocal SLO images . The central core originates from the strong backscattering of the erythrocytes and the top vessel wall while the dark bands originate from the shadow caused by the erythrocytes that is visible in layers below the vessel. In contrast to this observation the same vessel appears much thicker in the OCT image because only light that originates from the coherence plane contributes to the signal. Figure 5 shows averaged OCT and SLO images of the upper left vessel bifurcation region recorded with a smaller field of view.
Figure 6 shows small field of view images recorded from the central left part of the vessel visible in Fig. 4. The confocal SLO image (cf. Figure 6a) does not show details on the internal structures of the vessel because the top vessel wall is highly reflective. In order to investigate the contributions from different layers to the SLO signal we recorded en-face OCT images at different depths. When the coherence plane is placed at the top vessel wall (cf. Figure 6b) a similar image as the SLO image can be observed indicating that indeed the major contribution in the SLO originates from this layer. When the coherence plane is moved further in depth, internal structures (backscattering at erythrocytes) of the vessel can be observed (cf. Figure 6c). Placing the coherence gate at the center of the vessel reveals interesting features as can be observed in Fig. 6d). First of all, the vessel walls can be observed similar to Fig. 2. In addition three separated blood streams (marked with the numbers 1-3 in the image) can be observed within the vessel. By observing elongations of the erythrocytes in the data set we concluded that this vessel is a vein. The separated blood streams 1 and 2 originate from the large bifurcation (visible in Fig. 5a) while the stream 3 originates from the small branch on the left hand side of the image. Interestingly the blood streams do not completely mix together after the bifurcation. Noteworthy in this image are the observation of larger structures (15-20µm, indicated with an arrow in Fig. 6d) in the surrounding tissue which we could not identify with known structures of the retina. When the coherence gate is moved further into depth (cf. Figure 6e) the shadow originating from the vessel and a faint contribution from the bottom vessel wall can be clearly observed. In this layer the tissue surrounding the vessel shows a rather regular structure.
4. Discussion and conclusion
We demonstrated the capability of en-face OCT to investigate retinal vasculature with high transverse resolution. The instrument supports different fields of view ranging from 0.8°x0.8° up to ~4°x4° degrees. The larger field of view greatly simplifies the determination of the exact imaging location. The simultaneous recording of the SLO images enables a direct comparison between both imaging modalities. Thereby the same confocal detection optics (the light is coupled into a single mode fiber) is used for both channels. It is interesting to note that although we have a pixel to pixel correspondence between both imaging modalities different physical quantities are measured. While the SLO measures the light intensity backscattered from the retina, the light amplitude is measured in OCT. This leads to the high dynamic range within OCT images. To account for this the OCT data is usually converted to light intensity (by squaring the OCT signal amplitude) and compressed using a log scale for display. However, this results in a broadening of the point spread function and a different appearance of structures. We therefore displayed most of the OCT images in a linear amplitude scale as a compromise between broadening of structures and visibility of weak backscattering structures. Figure 7 shows for a comparison the different display options together with an SLO image. The full internal extension of the vessel is best seen in the image using a logarithmic intensity scale (Fig. 7a). In the image using a linear amplitude scale (Fig. 7b) the vessel appears slightly thinner because of the rather weak light backscattering of the particles that are close to the vessel wall. In the linear intensity scale image (Fig. 7c) a similar vessel appearance as in the SLO image (Fig. 7d) can be observed. This vessel appearance (bright central core that is surrounded by dark bands on both sides) is very typical for confocal SLO images of the vessels . One clear advantage of the OCT is the coherence gating which successfully eliminates contributions from other layers (nerve fiber layer and top vessel wall) that are clearly present in the SLO image and obscure the vessel structure.
However the comparison is performed only for the confocal SLO configuration which we realized using a single mode fiber approach. Similar results as in the conventional confocal detection scheme using a pinhole are obtained. It would be interesting to compare our results with an off-set pinhole configuration or a split detector scheme. However, these technologies are currently not implemented into our system.
In this study frame averaging was performed after registration of the images to a reference frame. However, due to motion in-frame distortions may occur which have not been corrected for. In order to improve the registration performance a strip wise registration procedure has been proposed . We omitted this procedure because we found that even without this strip wise registration the averaged frames showed a preservation of the structures seen in the single frame images. In addition our strip wise registration algorithm failed for images showing large vessels. Motion within the vessels will result in a displacement of structures from frame to frame resulting in registration errors. Perfect image registration, however, is essential in order to use additional vessel contrast methods that are based on measuring the variation between image frames [34, 52–54]. Due to our non-perfect image registration we did not include this analysis into our study.
An interesting observation with the en-face OCT system is the good visibility of the vascular network even without the use of contrast enhancements such as phase variance methods . A possible advantage of en-face OCT lies in the high modulation frequency of the OCT signal (3MHz) which results in reduced phase wash out due to motion compared to FD-OCT methods operating at 200 kHz A-scan rates. Therefore a high backscattering signal can be observed even in the case of larger vessels. This improves the visibility of these structures.
An asset of the system is the possibility to dynamically shift the focal plane simultaneously with the coherence gate through the tissue which provides 3D volumes of the retina with high resolution throughout imaging depth. Entirely sharp volumes of the retina with high transverse resolution can currently not be obtained with other OCT techniques. One critical issue with the proposed dynamic focus scheme is the alignment of the coherence plane and the focus plane. For that reason the online imaging plane of the OCT mode (before starting a measurement) is set at the junction between inner outer segments of photoreceptors (IS/OS). (For the labeling of the different retinal layers we refer to recent work by R. J. Jonnal et al. ). This can be done by changing the reference arm length. After closing the adaptive optics loop a fine adjustment of the focus can be necessary. Especially at imaging locations with a thick nerve fiber layer, the focus will be shifted to anterior layers because of the increasing contribution of this layer to the signal on the wavefront sensor. The correct focus is found when the photoreceptor mosaic appears sharpest. Another problem is accommodation of the subject. In this study we did not administer drugs in order to prevent accommodation. Therefore a residual mismatch between coherence plane and focal plane may arise. In order to minimize this effect, a lens was placed after the fixation target in order to produce an image of the target for fixation that is located in infinity. The 3D imaging mode is essential in order to determine the exact depth location of the images that are recorded with the 2D imaging mode over time (which are located at the same imaging depth). The en-face imaging mode enables the investigation of dynamic processes such as blood flow and allows for frame averaging similar to AO-SLO.
Financial assistance from the Austrian Science Fund (FWF projectP22329-N20) is gratefully acknowledged. The authors further acknowledge equipment and financial support from W. Drexler (Medical University of Vienna)
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