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Virtual Journal for Biomedical Optics

Virtual Journal for Biomedical Optics

| EXPLORING THE INTERFACE OF LIGHT AND BIOMEDICINE

  • Editor: Gregory W. Faris
  • Vol. 5, Iss. 8 — Jun. 8, 2010
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Simultaneous high-resolution retinal imaging and high-penetration choroidal imaging by one-micrometer adaptive optics optical coherence tomography

Kazuhiro Kurokawa, Kazuhiro Sasaki, Shuichi Makita, Masahiro Yamanari, Barry Cense, and Yoshiaki Yasuno  »View Author Affiliations


Optics Express, Vol. 18, Issue 8, pp. 8515-8527 (2010)
http://dx.doi.org/10.1364/OE.18.008515


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Abstract

Adaptive optics optical coherence tomography (AO-OCT) provides three-dimensional high-isotropic-resolution retinal images in vivo. We developed AO-OCT with a 1.03-μm probing beam and demonstrated high-penetration, high-resolution retinal imaging. Axial scans are acquired with a speed of 47,000 lines/s. AO closed loop is configured with a single deformable mirror. Seven eyes of 7 normal subjects were examined. Signal enhancement was found for all subjects. A rippled interface between nerve fiber layer and ganglion cell layer, boundary between ganglion cell layer and inner plexiform layer, and chorioscleral interface were identified. Simultaneous high-resolution and high-penetration choroidal imaging may be useful for microstructural investigation of photoreceptors and glaucomatous nerve-fiber abnormalities.

© 2010 Optical Society of America

1. Introduction

Alternative probing beam wavelengths have been studied for applications in ophthalmic OCT. Although most ophthalmic OCTs use a center wavelength of an 840-nm probe for posterior-eye imaging, one of the attractive and promising probing beam wavelengths is 1.0-μm. The local minima of the water absorption and dispersion were located at around 1.0-μm [11–14

11. G. M. Hale and M. R. Querry, “Optical constants of water in the 200-nm to 200-μm wavelength region,” Appl. Opt. 12, 555–563 (1973), http://ao.osa.org/abstract.cfm?URI=ao-12-3-555. [CrossRef] [PubMed]

]. The lower absorption of melanin and lower scattering of the retinal and choroidal fundus was found at 1.0-μm compared to 840-nm [15, 16

15. M. Hammer, A. Roggan, D. Schweitzer, and G. Muller, “Optical properties of ocular fundus tissues-an in vitro study using the double-integrating-sphere technique and inverse monte carlo simulation,” Phys. Med. Biol. 40, 963–978 (1995), http://stacks.iop.org/0031-9155/40/963. [CrossRef] [PubMed]

]. These properties enable high-penetration retinal imaging at 1.0-μm, and ophthalmic OCT using a 1.0-μm probe is known to provide better image contrast of the deep region of the eye as compared to the 840-nm probe [16–22

16. Y. Yasuno, Y. Hong, S. Makita, M. Yamanari, M. Akiba, M. Miura, and T. Yatagai, “In vivo high-contrast imaging of deep posterior eye by 1-um swept source optical coherence tomography andscattering optical coherence angiography,” Opt. Express 15, 6121–6139 (2007), http://www.opticsexpress.org/abstract.cfm? URI=oe-15-10-6121. [CrossRef] [PubMed]

]. Additionally, moderate chromatic aberrations of the eye exist at the 1.0-μm wavelength band [23

23. E. J. Fernändez and P. Artal, “Ocular aberrations up to the infrared range: from 632.8 to 1070 nm,” Opt. Express 16, 21199–21208 (2008), http://www.opticsexpress.org/abstract.cfm?URI= oe-16-26-21199. [CrossRef] [PubMed]

].

Cellular-level structural information cannot be easily obtained by conventional ophthalmic OCT, because residual ocular aberrations and eye motion deteriorate the lateral resolution of retinal images. In order to overcome these problems, dynamic compensation of ocular aberrations and high-speed image acquisition are essential. Adaptive optics (AO) technology has been used to compensate the ocular aberrations. AO has been combined with several types of conventional ophthalmic imaging devices such as a flood-illumination camera [24

24. J. Liang, D. R. Williams, and D. T. Miller, “Supernormal vision and high-resolution retinal imaging through adaptive optics,” J. Opt. Soc. Am. A 14, 2884–2892 (1997), http://josaa.osa.org/abstract.cfm? URI=josaa-14-11-2884. [CrossRef]

], scanning laser ophthalmoscope (SLO) [25–29

25. A. Roorda, F. Romero-Borja, W. J. D. III, H. Queener, T. J. Hebert, and M. C. W. Campbell, “Adaptive optics scanning laser ophthalmoscopy,” Opt. Express 10, 405–412 (2002), http://www.opticsexpress.org/ abstract.cfm?URI=oe-10-9-405. [PubMed]

], and OCT [30–40

30. B. Hermann, E. J. Fernändez, A. Unterhuber, H. Sattmann, A. F. Fercher, W. Drexler, P. M. Prieto, and P. Artal, “Adaptive-optics ultrahigh-resolution optical coherencetomography,” Opt. Lett. 29, 2142–2144 (2004), http: //ol.osa.org/abstract.cfm?URI=ol-29-18-2142. [CrossRef] [PubMed]

]. These combined techniques successfully improved the lateral resolution of retinal images and provided cellular-level structural information of the in vivo retina, such as photoreceptors and microcapillaries, and blood cells. In particular, AO-OCT provides three-dimensional high isotropic resolution for in vivo retinal images. Latest studies on AO-OCT have demonstrated high-speed retinal imaging with speed of up to 120,000 A-lines/s [41

41. C. Torti, B. Považay, B. Hofer, A. Unterhuber, J. Carroll, P. K. Ahnelt, and W. Drexler, “Adaptive optics optical coherence tomography at 120,000 depth scans/s for non-invasive cellular phenotyping of the living human retina,” Opt. Express 17, 19382–19400 (2009), http://www.opticsexpress.org/abstract.cfm? URI=oe-17-22-19382. [CrossRef] [PubMed]

]. Further, polarization sensitive-OCT (PS-OCT) has been combined with AO to realize the high-resolution measurement of birefringence [42

42. B. Cense, W. Gao, J. M. Brown, S. M. Jones, R. S. Jonnal, M. Mujat, B. H. Park, J. F. de Boer, and D. T. Miller, “Retinal imaging with polarization-sensitive optical coherence tomography and adaptive optics,” Opt. Express 17, 21634–21651 (2009), http://www.opticsexpress.org/abstract.cfm? URI=oe-17-24-21634. [CrossRef] [PubMed]

].

In this study, we demonstrate high-penetration and high-resolution AO-OCT using a center wavelength of 1.03-μm by combining a SD-OCT and an AO retinal scanner [43–45

43. K. Kurokawa, D. Tamada, S. Makita, and Yoshiaki Yasuno, “Adaptive optics retinal scanner for one-micrometer light source,” Opt. Express 18, 1406–1418 (2010), http://www.opticsexpress.org/abstract. cfm?URI=oe-18-2-1406. [CrossRef] [PubMed]

].

2. Methods

2.1. AO retinal scanner

In order to minimize the aberrations of the AO retinal scanner, cumulative astigmatism was canceled by the pair of vertically off-axis SMs [27

27. S. A. Burns, R. Tumbar, A. E. Elsner, D. Ferguson, and D. X. Hammer, “Large-field-of-view, modular, stabilized, adaptive-optics-based scanning laser ophthalmoscope,” J. Opt. Soc. Am. A 24, 1313–1326 (2007), http:// josaa.osa.org/abstract.cfm?URI=josaa-24-5-1313. [CrossRef]

, 43

43. K. Kurokawa, D. Tamada, S. Makita, and Yoshiaki Yasuno, “Adaptive optics retinal scanner for one-micrometer light source,” Opt. Express 18, 1406–1418 (2010), http://www.opticsexpress.org/abstract. cfm?URI=oe-18-2-1406. [CrossRef] [PubMed]

], as shown in Fig. 1(a). Total residual RMS wavefront error of the system was measured to be less than 0.1 μm.

A Badal optometer is placed in front of the the eye pupil, L6 and L7, for the correction of large amount of defocus of the eye [27

27. S. A. Burns, R. Tumbar, A. E. Elsner, D. Ferguson, and D. X. Hammer, “Large-field-of-view, modular, stabilized, adaptive-optics-based scanning laser ophthalmoscope,” J. Opt. Soc. Am. A 24, 1313–1326 (2007), http:// josaa.osa.org/abstract.cfm?URI=josaa-24-5-1313. [CrossRef]

]. The correctable range was from -9 D to 4 D.

Chromatic aberrations of the eye between 840-nm and 1.03-μm were less than around 0.5 D [23

23. E. J. Fernändez and P. Artal, “Ocular aberrations up to the infrared range: from 632.8 to 1070 nm,” Opt. Express 16, 21199–21208 (2008), http://www.opticsexpress.org/abstract.cfm?URI= oe-16-26-21199. [CrossRef] [PubMed]

]. The difference results in the displacement of the depth and lateral position of the focus, respectively called longitudinal chromatic aberration (LCA) and transversal chromatic aberration (TCA). Although the TCA between 840-nm and 1.03-μm results in different fields of view between 840-nm and 1.03-μm, the field of view of the system is calibrated by 1.03-μm and this TCA does not introduce a significant error. The LCA between 840-nm and 1.03-μm was canceled by adjusting the lens distance between L4 and L5.

In order to eliminate the abrupt fluctuation of aberrations caused by the pull-back action of the galvanometric scanner, the frequency of the AO closed loop was configured to be identical to a submultiple of the frequency of the galvanometric scanner. Further, HASO32 was driven by a synchronization trigger that was generated from the position monitor of the fast galvanometric scanner.

Fig. 1. Schematic of the AO-OCT. WDM: Wavelength division multiplexing coupler, C: Circulator, PC: Polarization controller, FC: Fiber coupler. (a) The side and top views of the optical setup of the AO retinal scanner. L#: Lenses, LP#: Linear polarizers, BS: Beam splitter, Hs: Harmonic separator, ST: Stop, SM#: Spherical mirrors, FM#: Flat mirrors, WS: Wavefront sensor, DM: Deformable mirror, VS: Vertical galvanometric scanner, HS: Horizontal resonant scanner mounted galvanometric scanner, APD: Avalanche photodiode. (b) Reference arm. ND: ND filter. (c) Spectrometer.

A fixation target was used to reduce the eye motion and to allow imaging of given eccentricity. As the fixation target, grid lines were printed on a paper and placed at a retinal conjugated plane. The measured eccentricity was roughly calibrated using the fixation target up to an eccentricity of 9 degrees.

2.2. SD-OCT

The OCT system is based on an unbalanced Michelson interferometer, the schematic of which is shown in Fig. 1. The light source for the probing beam is 1-μm SLD (Superlum) with a center wavelength of 1.03-μm and a spectral bandwidth (FWHM) of 106 nm. The theoretical axial resolution was 3.4 μm in tissue, and the measured axial resolution was 5.6 μm in tissue. The 1.03-μm light is coupled with 840-nm light using a wavelength division multiplexing (WDM) coupler that was designed for 1.06-μm and 840-nm. Then, the light passes through a circulator and a fiber coupler, both of which are designed for 1.06-μm light. A polarization controller is placed before the circulator and aligned such that polarization mode dispersion is avoided; this mainly occurs at the circulator. After passing through the fiber coupler, 20% of the light is delivered to the reference arm and 80 % of the light is delivered to the AO retinal scanner.

The backscattered light from the retina is recoupled and sent back to the fiber coupler, where it is recombined with the reference signal. 80% of the recombined light is delivered to the spectrometer and collimated by an achromatic lens, diffracted by a reflective grating, and detected by an InGaAs line-scan camera with 1024-pixel resolution (SUI1024LDH, Sensors Unlimited, Inc.) at a speed of 47,000 lines/s. A polarization controller is aligned to optimize the polarization state; this improves the efficiency of the reflection grating. The details of the SD-OCT system have been described in [44

44. S. Makita, T. Fabritius, and Y. Yasuno, “Full-range, high-speed, high-resolution 1-μm spectral-domain optical coherence tomography using BM-scan for volumetric imaging of the human posterior eye,” Opt. Express 16, 8406–8420 (2008), http://www.opticsexpress.org/abstract.cfm?URI=oe-16-12-8406. [CrossRef] [PubMed]

, 45

45. F. Jaillon, S. Makita, M. Yabusaki, and Y. Yasuno, “Parabolic BM-scan technique for full range Doppler spectral domain optical coherence tomography,” Opt. Express 18, 1358–1372 (2010), http://www. opticsexpress.org/abstract.cfm?URI=oe-18-2-1358. [CrossRef] [PubMed]

].

The optical powers on the cornea was 0.9 mW for the probing beam and 0.07 mW for the AO beacon. These were below the American National Standard Institute (ANSI) safety limit [46

46. Z136 Committee, American National Standard for Safe Use of Lasers: ANSI Z136.1-2000 (Laser Institute of America, 2003).

, 47

47. F. C. Delori, R. H. Webb, and D. H. Sliney, “Maximum permissible exposures for ocular safety (ansi 2000), with emphasis on ophthalmic devices,” J. Opt. Soc. Am. A 24, 1250–1265 (2007), http://josaa.osa.org/ abstract.cfm?URI=josaa-24-5-1250. [CrossRef]

].

2.3. System sensitivity

In order to evaluate the system performance of AO-OCT, the system sensitivity was evaluated.

The sensitivity measurement was performed using a model eye with a mirror reflector at a retinal conjugated plane. During the measurement, the deformable mirror was flattened, and hence the system aberrations were not corrected. The measured sensitivity of AO-OCT was 83 dB.

In the case of non-AO SD-OCT which was described in Ref. [44

44. S. Makita, T. Fabritius, and Y. Yasuno, “Full-range, high-speed, high-resolution 1-μm spectral-domain optical coherence tomography using BM-scan for volumetric imaging of the human posterior eye,” Opt. Express 16, 8406–8420 (2008), http://www.opticsexpress.org/abstract.cfm?URI=oe-16-12-8406. [CrossRef] [PubMed]

, 45

45. F. Jaillon, S. Makita, M. Yabusaki, and Y. Yasuno, “Parabolic BM-scan technique for full range Doppler spectral domain optical coherence tomography,” Opt. Express 18, 1358–1372 (2010), http://www. opticsexpress.org/abstract.cfm?URI=oe-18-2-1358. [CrossRef] [PubMed]

], but operated at a half-range mode, the sensitivity was measured to be 93.0 dB, while the theoretical shot-noise-limited sensitivity is 100.8 dB. The degradation of the system sensitivity of 7.8 dB may be because of relatively high camera noise. Since the non-AO SD-OCT and AO-OCT shares a spectrometer, this sensitivity degradation may exist also in AO-OCT.

The theoretical shot-noise-limited sensitivity of AO-OCT is 95.3 dB. Accounting for the sensitivity loss based on the camera noise, a predicted sensitivity, which is limited by the camera noise and relative intensity noise, becomes 87.5 dB, and 4.5 dB loss remains unexplained. This unexplained loss could be explained by the misalignment of the mirror reflector and the system aberration. In in vivo AO-OCT measurements, the signal intensity will increase more by the aberration correction (See Section 3.2).

2.4. Measurement protocol

We carried out tests for the 7 eyes of 7 normal subjects. A patch of the retina at an eccentricity of approximately 6 degree on the nasal retina was scanned iteratively. The imaging retinal eccentricity was set by location of external fixation point. For the stabilization of fixation, the dominant eyes were measured. For all subjects, the measured field of view is 4.5 degree on the retina with calibration for emmetropia.

The subjects followed proper procedures before the measurement. First, two drops of 0.5% tropicamide and 0.5% phenylephrine hydrochloride were applied for pupil dilation; the subject were then made to sit in a chair and place their head on a chin rest. Second, the residual defocus for a subject was roughly compensated using the Badal optometer by monitoring the RMS wavefront error, and the subjects manually controlled the axial position of the fixation target. Third, we started to run the AO closed loop. The eye was centralized by the subject based on the direction of an operator who monitored the transversal location of the pupil by a Shack-Hartmann image. The depth position was controlled by monitoring the OCT images.

Note that the defocus of the eye was not constant in time even we applying the eye drops because of the temporal dynamics of accommodation [48

48. H. Hofer, P. Artal, B. Singer, J. L. Aragón, and D. R. Williams, “Dynamics of the eye’s wave aberration,” J. Opt. Soc. Am. A 18, 497–506 (2001), http://josaa.osa.org/abstract.cfm?URI=josaa-18-3-497. [CrossRef]

]. A single OCT volume requires a measurement time of a few second. For each measurement session, the aberrations of the eye was successfully compensated with sufficiently high iteration frequency of the AO closed loop. However, between each session, the amount of defocus will become larger than the stroke of the deformable mirror. This requires the defocus correction using the Badal optometer.

3. Results

3.1. Retinal images

The effect of AO correction is shown in Fig. 2. Images without AO were obtained when the deformable mirror was flattened and the defocus of the eye was corrected using the Badal optometer. With AO correction, the residual RMS wavefront error was typically 0.1 μm. The signal strength was improved and the image contrast was enhanced. However, the improvement of the lateral resolution was not huge as compared to that observed by AO-SLO, because the lateral resolution of AO-OCT is high even without AO. This difference might be attributable to the difference in the confocality (see Section 4.1).

In general, the depth dependent intensity change is not significant using 1-μm AO-OCT in comparison to 800-nm AO-OCT. There are two possible reasons: First, the 1-μm light results in wider depth of focus. Second, our system has relatively higher LCA in comparison to 800-nm AO-OCT using an achromatizer [37

37. R. J. Zawadzki, B. Cense, Y. Zhang, S. S. Choi, D. T. Miller, and J. S. Werner, “Ultrahigh-resolution optical coherencetomography with monochromatic and chromaticaberration correction,” Opt. Express 16, 8126–8143 (2008), http://www.opticsexpress.org/abstract.cfm?URI=oe-16-11-8126. [CrossRef] [PubMed]

, 51

51. E. J. Fernändez, A. Unterhuber, B. Považay, B. Hermann, P. Artal, and W. Drexler, “Chromatic aberration correction of the human eye for retinal imaging in the near infrared,” Opt. Express 14, 6213–6225 (2006), http://www.opticsexpress.org/abstract.cfm?URI=oe-14-13-6213. [CrossRef] [PubMed]

].

3.2. Analysis of image contrast

In order to confirm the feasibility of the system, several eyes were examined, as shown in Fig.4. The subject’s demographics are summarized in Table 3.2. An improvement of the image contrast was found in all subjects with AO.

In order to analyze the image contrast quantitatively, signal-to-noise ratio (SNR) and signal gains were calculated, as shown in Table 3.2. The SNR is defined as 10logmax(I)σ2. Here, max(I) is the maximum intensity of the selected layers: NFL-IPL including the regions from the NFL to the IPL, INL-ONL including regions from the INL to the ONL, PRL-Choroid including regions from the PRL to the Choroid. σ2 is a signal variance at a vitreous region which is considered as noise energy. Here, we assumed that the noise energy is constant between AO-on and AO-off. On the other hand, the signal gains were calculated as the ratio of the SNR with AO-on and the SNR with AO-off. Hence, the gain can be regarded as the gain of the signal energy. Note that the signal roll-off of the system was taking into account for the calculation. The signals were normalized by a fitting function of the measured roll-off sensitivity. As a result, positive signal gains were observed with all the subjects except NFL-IPL of subjects D and F. These negative gains are because of decreasing depth of focus by AO. Paired two-sided t-tests were performed to evaluate the gain of SNR for several depth regions. For the t-test, t.test function of a statistic language R was employed. For all 3 depth positions, statistically significant improvement of SNR was observed (P = 0.006, P = 0.004, and P = 0.014 for the regions of NFL-IPL, ONL-ONL and PRL-Choroid). The averaged signal gains were 7.0 dB, 8.2 dB, and 4.9 dB for the three depth regions, respectively.

Fig. 2. (a) Color fundus photographs with the field of view of 20 degree × 20 degree. White arrow indicates the measured location of the retina. Single B-scan images (b) without AO correction and (c) with AO correction. Black bar indicates 100 μm on the retina.
Fig. 3. Averaged B-scan image. (b) shows the region enclosed by the orange box in (a). Black bar indicates 100 μm on the retina. Red arrows indicate the chorioscleral interface. NFL: nerve fiber layer, GCL: ganglion cell layer, IPL: inner plexiform layer, INL: inner nuclear layer, OPL: outer plexiform layer, ONL: outer nuclear layer, PRL: photoreceptor layer, RPE: retinal pigment epithelium, CC: choriocapillaris, ELM: external limiting membrane, IS/OS: inner/outer segment junction.

In order to quantify the sensitivity of clinical features objectively, two OCT engineers who have more than 4-year experience in ophthalmic OCT (MY and SM) were involved as graders. The sensitivities are summarized in Table 3.2. The sensitivity is categorized in binary grades: 0 for invisible or +1 for visible. The targeted features are a rippled interface between the NFL, and the GCL (NFL/GCL), an interface between the GCL and the IPL (GCL/IPL), and a chorioscleral interface (CSI). As a result, the sensitivity of NFL/GCL was 57.1% (MY and SM), that of GCL/IPL was 85.7% (MY) and 57.1% (SM), and that of CSI was 71.4% (MY and SM). The inter-grader agreements were 71.4%, 71.4%, and 100.0% for the sensitivity of the NFL/GCL, GCL/IPL, and CSI, respectively. Note that the feature of the CSI was observed with more than 71.4% and it is the advantage of high-penetration AO-OCT.

Fig. 4. Averaged B-scan images. The left part of the image was taken without AO and the right part of the image was taken with AO. (a)–(g) are corresponding with the subject ID of A–G.

4. Discussion

4.1. Lateral resolution

In this section, we discuss the lateral resolution of AO-OCT using numerical simulations based on experimentally measured wavefront aberrations of the eye. The lateral resolution was estimated from the full width at half maximum of the point spread function (PSF). In confocal microscopy, PSF is represented by the following equation [49

49. T. Wilson and A. R. Carlini, “Size of the detector in confocal imaging systems,” Opt. Lett. 12, 227–229 (1987), http://ol.osa.org/abstract.cfm?URI=ol-12-4-227. [CrossRef] [PubMed]

, 50

50. E. Fernändez and W. Drexler, “Influence of ocular chromatic aberration and pupil size on transverse resolution in ophthalmic adaptive optics optical coherence tomography,” Opt. Express 13, 8184–8197 (2005), http:// www.opticsexpress.org/abstract.cfm?URI=oe-13-20-8184. [CrossRef] [PubMed]

]:

PSF=PSFill×(PSFobs*D)
(1)

Here, PSFill and PSFobs are point spread functions of illumination and observation, respectively. * denotes a convolution operator. D is the finite-sized pinhole; in the case of AO-OCT, a single-mode fiber acts as a detection pinhole. We applied a Gaussian function, D=exp(X2+Y22c2), to model the single-mode fiber detection. Here, c is the multiplication of the magnification and the mode field diameter of the single mode fiber, c=(M×Dmode4). In our system, M = 0.91 and Dmode = 6 μm. Note that the M = 0.91 is only valid for emmetropia because M depends on the configuration of the Badal optometer. Although a similar analysis was performed by Fernández et al. [50

50. E. Fernändez and W. Drexler, “Influence of ocular chromatic aberration and pupil size on transverse resolution in ophthalmic adaptive optics optical coherence tomography,” Opt. Express 13, 8184–8197 (2005), http:// www.opticsexpress.org/abstract.cfm?URI=oe-13-20-8184. [CrossRef] [PubMed]

], they did not take into account the detection pinhole and the magnification from the single-mode fiber to the retina, M, which is a function of the axial eye length if the eye is not emmetropic, i.e., M is a function of the stroke of the Badal optometer. In other words, M is used as a normalization factor between the retinal plane and the detector plane.

Table 1. Subject’s characteristics. Sph and Cyl: spherical and cylindrical powers in diopter, respectively, of the eye.

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Table 2. Signal gain [dB]. NFL-IPL includes the regions from the NFL to the IPL, INL-ONL includes the regions from the INL to the ONL, PRL-Choroid includes the regions from the PRL to the Choroid.

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We used Eq. (1) to estimate the lateral resolution with several magnifications, as shown in Fig. 5(a). The curves of AO-on and AO-off were calculated on the basis of typical aberrations of a normal subject (Subject-A) measured using our AO retinal scanner while AO was turned on and turned off, respectively. The residual amount of the aberration coefficients is shown in Figs. 5(b). Figure 5(c) shows the evolution of the RMS wavefront error along the time course of the AO closed loop.

The lateral resolution was estimated from Eq. (1). However, this analysis did not take into account the other factors that contributed to the lateral resolution, such as chromatic aberrations [37

37. R. J. Zawadzki, B. Cense, Y. Zhang, S. S. Choi, D. T. Miller, and J. S. Werner, “Ultrahigh-resolution optical coherencetomography with monochromatic and chromaticaberration correction,” Opt. Express 16, 8126–8143 (2008), http://www.opticsexpress.org/abstract.cfm?URI=oe-16-11-8126. [CrossRef] [PubMed]

, 50

50. E. Fernändez and W. Drexler, “Influence of ocular chromatic aberration and pupil size on transverse resolution in ophthalmic adaptive optics optical coherence tomography,” Opt. Express 13, 8184–8197 (2005), http:// www.opticsexpress.org/abstract.cfm?URI=oe-13-20-8184. [CrossRef] [PubMed]

, 51

51. E. J. Fernändez, A. Unterhuber, B. Považay, B. Hermann, P. Artal, and W. Drexler, “Chromatic aberration correction of the human eye for retinal imaging in the near infrared,” Opt. Express 14, 6213–6225 (2006), http://www.opticsexpress.org/abstract.cfm?URI=oe-14-13-6213. [CrossRef] [PubMed]

], underestimation of odd aberrations [52

52. P. Artal, S. Marcos, R. Navarro, and D. R. Williams, “Odd aberrations and double-pass measurements of retinal image quality,” J. Opt. Soc. Am. A12, 195–201 (1995),http://josaa.osa.org/abstract.cfm?URI= josaa-12-2-195. [CrossRef]

], and speckle size [42

42. B. Cense, W. Gao, J. M. Brown, S. M. Jones, R. S. Jonnal, M. Mujat, B. H. Park, J. F. de Boer, and D. T. Miller, “Retinal imaging with polarization-sensitive optical coherence tomography and adaptive optics,” Opt. Express 17, 21634–21651 (2009), http://www.opticsexpress.org/abstract.cfm? URI=oe-17-24-21634. [CrossRef] [PubMed]

].

Although our system was operated in the 1-μm wavelength band, which has moderate chromatic aberration, the effect of chromatic aberration remained, degraded the image qualities, and results in the relatively poorer depth dependent intensity change.

Double-pass aberration measurement causes underestimation of odd aberrations of the system and the eye [52

52. P. Artal, S. Marcos, R. Navarro, and D. R. Williams, “Odd aberrations and double-pass measurements of retinal image quality,” J. Opt. Soc. Am. A12, 195–201 (1995),http://josaa.osa.org/abstract.cfm?URI= josaa-12-2-195. [CrossRef]

]. Namely, if the retina was regarded as an ideal scatterer (no specular reflection), the HASO32 could measure a pure single pass aberration. In practice, the light back-scattered from the retina contains the double pass information because of the specular reflection components. This becomes the error source of the aberration measurement.

The speckle size depends on the illumination diameter and the spectral bandwidth of the light source, and it degrades the image quality of the OCT images. Therefore, the actual lateral resolution is broadened to a greater extent than the estimation. For a more accurate analysis, a reasonable and reliable estimation of the lateral resolution is required.

Table 3. Sensitivity of several features. The targeted features are a rippled interface between the NFL and the GCL (NFL/GCL), an interface between the GCL and the IPL (GCL/IPL), and a chorioscleral interface (CSI). A.S.: Average Sensitivity.

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4.2. non-AO SD-OCT vs. AO-OCT vs. AO-SLO

System performances of non-AO SD-OCT [44

44. S. Makita, T. Fabritius, and Y. Yasuno, “Full-range, high-speed, high-resolution 1-μm spectral-domain optical coherence tomography using BM-scan for volumetric imaging of the human posterior eye,” Opt. Express 16, 8406–8420 (2008), http://www.opticsexpress.org/abstract.cfm?URI=oe-16-12-8406. [CrossRef] [PubMed]

], AO-OCT, and AO-SLO [43

43. K. Kurokawa, D. Tamada, S. Makita, and Yoshiaki Yasuno, “Adaptive optics retinal scanner for one-micrometer light source,” Opt. Express 18, 1406–1418 (2010), http://www.opticsexpress.org/abstract. cfm?URI=oe-18-2-1406. [CrossRef] [PubMed]

] were compared, as shown in Table 4.2. AO-OCT provided retinal images with high isotropic resolution; however, it had a limited field of view and lower sensitivity than the non-AO SD-OCT. The field of view is limited to an isoplanatic patch [53

53. P. Bedggood, M. Daaboul, R. Ashman, G. Smith, and A. Metha, “Characteristics of the human isoplanatic patch and implications for adaptive optics retinal imaging,” J. Biomed. Opt. 13, 024008 (2008), http://link. aip.org/link/?JBO/13/024008/1. [CrossRef] [PubMed]

]. Although the problem of lower sensitivity could be overcome by the signal gain afforded by AO and the contribution of the Stiles-Crowford effect [54–59

54. A. Roorda and D. R. Williams, “Optical fiber properties of individual human cones,” J. Vis. 2, 404–412 (2002), http://journalofvision.org/2/5/4/. [CrossRef]

], the sensitivity of AO-OCT cannot easily exceed that of non-AO SD-OCT.

AO-OCT is expected to provide higher-resolution retinal images than AO-SLO with the advantage of higher confocality. On the other hand, the acquisition speed of AO-SLO reduces the transversal motion artifacts.

Table 4. System characteristics of non-AO SD-OCT, AO-OCT, and AO-SLO.

table-icon
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5. Conclusion

We have demonstrated simultaneous high-penetration choroidal imaging and high-resolution retinal imaging using AO-OCT. With AO correction, this technique exhibited a statistically significant signal gain. Although the signal gain of the transversal resolution was lower than that of AO-SLO, it was sufficiently high to observe the retinal microstructures. Although non-AO-OCT with a high numerical aperture could provide moderately high resolution, the image contrast was dim, as shown in the measurement of AO-off. Therefore, the main advantage of this AO-OCT is the simultaneous realization of high-contrast and high-resolution. The retinal images provide information about not only the thickness of the retinal nerve fiber layer but also the structure of the nerve fiber. This information might be useful for carrying out a more detailed investigation of the neural abnormalities caused by glaucoma. Further, the identification of the chorioscleral interface indicates the high penetration capability of the system. This enables the investigation of microscopic structures of the deep region of the eye.

Fig. 5. (a)Lateral resolution. (b) residual aberration coefficients and (c) residual RMS wave-front error, which are measured using the HASO32. The error bars (green and red dashed lines) indicate the standard deviation of the residual aberration coefficients of AO-off and AO-on.

Acknowledgment

This research was supported in part by Japan Science and Technology Agency through a program of Development of Systems and Technology for Advanced Measurement and Analysis.

References and links

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54.

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55.

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57.

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58.

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59.

W. Gao, B. Cense, Y. Zhang, R. S. Jonnal, and D. T. Miller, “Measuring retinal contributions to the optical stiles-crawford effect with optical coherence tomography,” Opt. Express 16, 6486–6501 (2008), http://www. opticsexpress.org/abstract.cfm?URI=oe-16-9-6486. [CrossRef] [PubMed]

OCIS Codes
(110.4500) Imaging systems : Optical coherence tomography
(170.4460) Medical optics and biotechnology : Ophthalmic optics and devices
(170.4500) Medical optics and biotechnology : Optical coherence tomography
(110.1080) Imaging systems : Active or adaptive optics

ToC Category:
Medical Optics and Biotechnology

History
Original Manuscript: January 13, 2010
Revised Manuscript: April 1, 2010
Manuscript Accepted: April 7, 2010
Published: April 8, 2010

Virtual Issues
Vol. 5, Iss. 8 Virtual Journal for Biomedical Optics

Citation
Kazuhiro Kurokawa, Kazuhiro Sasaki, Shuichi Makita, Masahiro Yamanari, Barry Cense, and Yoshiaki Yasuno, "Simultaneous high-resolution retinal imaging and high-penetration choroidal imaging by one-micrometer adaptive optics optical coherence tomography," Opt. Express 18, 8515-8527 (2010)
http://www.opticsinfobase.org/vjbo/abstract.cfm?URI=oe-18-8-8515


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References

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  4. R. Leitgeb, C. Hitzenberger, and A. Fercher, "Performance of Fourier domain vs. time domain optical coherence tomography," Opt. Express 11, 889-894 (2003), http://www.opticsexpress.org/abstract.cfm?URI=oe-11-8-889. [CrossRef] [PubMed]
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