Measurement of refractive error in the human eye using optical coherence tomography

Abstract

Certain embodiments disclosed herein provide an apparatus for performing ophthalmoscopic measurements, comprising an optical coherence tomography (OCT) imaging system configured to output a collimated laser beam that is scanned across a patient's eye. Lines in the OCT image are identified, and the lines represent the retina of the patient's eye. The shape of the lines is then used by a refractive error calculator to measure the refractive error of the patient's eye. The height of the lines due to curvature indicates the degree of myopia or hyperopia. More complex line curvatures correspond to higher-order aberrations. OCT images at multiple cross sections can be used to calculate a wavefront aberration map.

Description

[Technical Field]

[0001] The present disclosure relates generally to techniques for measuring refractive error of the eye. [Background technology]

[0002] Optical coherence tomography (OCT) systems direct a coherent beam of light into a patient's eye, where a portion of the coherent beam is scattered and returned to the OCT system. Interference between the scattered light and a portion of the original coherent beam is used to measure the location of the scattered point. Ophthalmic OCT is primarily used to measure eye dimensions and diagnose various eye conditions. Ophthalmic OCT may also be used to image the anterior chamber of the eye. Summary of the Invention [Means for solving the problem]

[0003] The present disclosure relates generally to measuring refractive error of the eye using OCT.

[0004] In this patent application, an optical coherence tomography (OCT) imaging system is configured to scan a parallel beam across a patient's eye. A computing device is configured to receive OCT signals reflected from the retina. The apparent shape of the retina, and in particular, the deviation of the apparent shape of the retina from a straight line, is used to calculate the refractive error of the eye. In an eye without any optical error, the apparent image of the retina is a straight line.

[0005] The following description and the related drawings set forth in detail certain illustrative features of the one or more embodiments.

[0006] The accompanying drawings depict certain aspects of one or more embodiments and therefore should not be considered as limiting the scope of the disclosure. [Brief explanation of the drawings]

[0007] [Figure 1]An ideal, aberration-free crystalline lens is shown illuminated with multiple parallel rays so that all rays are focused at a focal point F. [Figure 2A] 1 shows an OCT image of an eye without aberrations, showing the actual light ray propagation to the focal point on the retina of the eye. [Figure 2B] 1 illustrates an aberration-free OCT image of an eye showing an extended focus representation, in accordance with certain embodiments. [Figure 3] Illustrates Fermat's principle as applied to light rays passing through a crystalline lens. [Figure 4] Let us denote a surface S, which is a surface for which the light propagation time from plane A to surface S is equal. [Figure 5A] 1 shows an OCT image of a plane P offset above the focal point of the crystalline lens in the OCT image, according to certain embodiments. [Figure 5B] 1 shows an OCT image of a plane P offset above the focal point of the crystalline lens in the OCT image, according to certain embodiments. [Figure 6A] 1 illustrates apparent retinal images acquired for a myopic eye and a hyperopic eye, according to certain embodiments. [Figure 6B] 1 illustrates apparent retinal images acquired for a myopic eye and a hyperopic eye, according to certain embodiments. [Figure 7] 1A-1C illustrate an approach for relating the sag of an apparent image of the retina representing a plane to the distance of the plane from the lens focal plane, according to certain embodiments. [Figure 8A] FIG. 1 illustrates a system for testing the relationship between the sag of an apparent image of the retina representing a plane and the distance of the plane from the lens focal plane, in accordance with certain embodiments. [Figure 8B] FIG. 1 illustrates a system for testing the relationship between the sag of an apparent image of the retina representing a plane and the distance of the plane from the lens focal plane, in accordance with certain embodiments. [Figure 9] This is an OCT image of the reflective surface P captured by OCT when the aspheric lens was not in the optical path. The image of the reflective surface P is a straight line, proving that the rays of the scanned OCT beam are parallel to each other and that the reflective surface P is perpendicular to the OCT rays. [Figure 10] 1 is a set of OCT images of the apparent reflecting surface P acquired at different vertical positions of the crystalline lens, according to certain embodiments. [Figure 11] 1 is an OCT image showing apparent image sag of a reflective surface P, according to certain embodiments. [Figure 12A] An apparent image of a reflective surface P facing the incident OCT beam and having a flat surface of an aspheric lens tilted to induce high-order aberrations is shown. [Figure 12B] An apparent image of a reflective surface P facing the incident OCT beam and having a flat surface of an aspheric lens tilted to induce high-order aberrations is shown. [Figure 13] FIG. 1 is a diagram of an eye showing dimensions that may be used to calculate the refractive error of the eye based on the sag measured in an OCT image, in accordance with certain embodiments. [Figure 14A] 10 illustrates the measurement of high-order aberrations using OCT images, according to certain embodiments. [Figure 14B] 10 illustrates the measurement of high-order aberrations using OCT images, according to certain embodiments. DETAILED DESCRIPTION OF THE INVENTION

[0008] For clarity, where possible, the same reference numerals have been used to denote identical elements that are common to the figures, and it is contemplated that elements and features of one embodiment may be beneficially incorporated in other embodiments without specific indication.

[0009] The refractive error of the human eye can be measured using Shack-Hartmann, Talbot-Moiré, Tracy, or Tscherning wavefront meters. Below, we describe a new class of devices that can measure the refractive error and wavefront error of the human eye. The devices are based on Fermat's principle and use an OCT scanner. Although ophthalmic OCT was invented in 1991, it has not been used or considered as a device capable of measuring the refractive error and wavefront error of the eye.

[0010] An optical coherence tomography (OCT) system directs one beam from a laser beam (called the signal beam) into the patient's eye and another beam from the laser beam (called the reference beam) into the reference arm. The reflected beams returning from the two arms are then combined. The combined beams form a wavelength-dependent interference spectrum. By analyzing the interference spectrum, two pieces of information can be obtained: (a) the intensity of the reflected light returning from the eye, and (b) the Z depth location of the origin of the returned reflected light. The Z location can be expressed either as Z distance or as the light propagation time T = Z / c, where c is the speed of light. For the purposes of the embodiments described herein, the propagation time T is the most important parameter. A measurement along a given propagation direction is called an A-scan. Thousands of A-scans can be obtained by scanning the OCT beam in the X and Y directions. From multiple A-scans, a 3D image of the eye can be reconstructed. Typically, the 3D image displays an X / Z or Y / Z side view or an X / Y front view. Ophthalmic OCT is primarily used to image the cornea, anterior chamber, lens, vitreous, and retina. The images are used to measure eye dimensions and diagnose various types of eye conditions.

[0011] Referring to Figure 1, Fermat's principle, which has a history of over 350 years, is used to measure refractive error in the eye using OCT. This principle states that the propagation time of flight from plane A to focal point F is the same for each ray. Fermat's principle can be applied to an ideal lens L illuminated with a plane wave propagating parallel to the lens's optical axis. The geometric path length from plane A to focal point F is shortest for ray 1, but ray 1 is the ray that travels the longest path through the lens material. The propagation time through the lens is slower, corresponding to c / n, where c is the speed of light in a vacuum and n is the refractive index of the lens material.

[0012] The following Figures 2A-2B are used to illustrate how Fermat's principle applies to OCT images of an eye with no refractive error. Note that the OCT device displays the propagation time of the OCT ray as a function of the input R position of the OCT ray.

[0013] Referring to Figure 2A, collimated OCT light beams are propagated into the eye. If the eye has no refractive error and all beams converge on the retina, the propagation time will be the same for all rays shown in Figure 2A.

[0014] When creating an OCT image, the propagation time is plotted as a function of the R input position of the incident beam. Refraction and beam deflection on the curved surfaces of the cornea and lens are ignored when displaying an OCT image of the eye. Because the propagation time is the same for all rays, the retinal focus appears as a straight line (4), as shown in Figure 2B. The straight line (4) in Figure 2B is not an image of the retina, but rather a horizontally expanded image of point (4) on the retina in Figure 2A. It is important to note that the horizontal line is not a retinal image; it is merely an image of the retina. In reality, the horizontal line is an image of a small horizontally expanded portion of the retina. The length of the expansion is equal to the scan length of the OCT beam above the cornea. The latter is often limited by the size of the pupil.

[0015] Certain OCT devices partially account for beam deflection caused by refraction on inclined surfaces, but for purposes of the embodiments described herein, such OCT software is used that does not correct for beam deflection on curved and inclined surfaces.

[0016] Currently, there are two existing techniques for OCT imaging: retinal OCT and anterior chamber OCT. In anterior chamber OCT, the central axial beam of the OCT above the cornea is always strictly vertical and scanned along the horizontal direction. The beam waist of the incident OCT beam is located at the anterior chamber-lens complex, which helps the OCT efficiently collect backscattered light from the anterior chamber. In retinal OCT, the beam is focused at the retina level, and the axial beam of the OCT beam is not strictly parallel. Instead, the incident beam forms the shape of a hand-held fan whose rotation axis is located at the pupil of the eye. In this way, an extremely large angular range of the retina can be scanned.

[0017] The OCT described in this patent application is similar to anterior chamber OCT in the sense that the central axial beam is always strictly vertical and scanned horizontally, but the beam waist is focused at the level of the retina. Such focusing ensures efficient collection of backscattered light from the retina. Such systems are referred to in this patent application as improved anterior chamber OCT. Some variation from strictly vertical can be corrected, as described below, in some implementations so that the axial beams are only substantially vertical and parallel to each other, e.g., within 2 degrees of vertical and parallel.

[0018] Referring to Figure 3, suppose a circular surface S is centered at the focal point F of a lens L. Also, Fermat's principle dictates that the propagation time of light from plane A to spherical surface S is equal for all rays. Therefore, S is a surface of equal propagation time.

[0019] Referring to FIG. 4, a reflecting plane P is assumed to be above the focal point F. As shown in FIG. 4, rays 1 and 2 arrive at plane P first. Ray 3 is the last ray to arrive at plane P. The arrival time difference Δ is Δ / c, where c is the speed of light. The propagation time from plane A to spherical surface S is equal for all rays. Referring to FIGS. 5A and 5B, the appearance of surface P on an OCT image measuring refractive error is a curve with a sag Δ measured along the optical axis of the lens L.

[0020] Figure 6A shows that when imaging a myopic eye, the line representing light reflected from the retina bends upward (towards the OCT scanner), and Figure 6B shows that when imaging a hyperopic eye, the line representing light reflected from the retina bends downward (away from the OCT scanner).

[0021] Figure 7 shows the calculation of sag Δ using an aspheric lens L and an OCT scanner. Δ refers to the phase plane sag at a specific X distance from the focal plane. The OCT scanner had an OCT scan length of 6 mm (the limit of the OCT used) and scanned horizontally across an AL108 Edmund Optics aspheric lens with a numerical aperture (NA) of 0.55 and an input diameter of 10 mm. For a horizontal scan length of 6 mm, the peripheral OCT ray has an NA = 0.55 * 0.6 = 0.33. The angle of incidence α of the peripheral ray is α = arc sin NA = arc sin 0.33 = 19.27°.

[0022] Sag Δ can be calculated as Δ=X*[(1 / cos19.27°)−1)=X*0.0594 Another way to express Δ is Δ=59.4 μm / mm from the focal plane FP.

[0023] 8A and 8B, anterior chamber OCT, i.e., the Optoview iVue OCT using an anterior chamber scanning scheme, was used to measure the value of sag Δ for multiple values ​​of X. In the experimental setup, X was varied by moving lens L relative to reflection plane P. The value of X corresponds to the distance of lens L from the position where the focal point of lens L is at reflection plane P. Because the distance between the OCT and plane P did not change, moving the lens up or down relative to reflection plane P does not move the center point of the bow-like image at reflection plane P up or down. Instead, only the curvature of the bow-like image changes.

[0024] Figure 9 shows an OCT image of a plane P captured without an aspheric lens in the OCT beam. The reflective surface P was a flat, polished stainless steel surface. The straight line demonstrates that the OCT scanner was optically correctly aligned, i.e., the incident OCT light beams were (a) parallel and (b) perpendicular to the stainless steel surface.

[0025] Figure 10 shows OCT images taken with an aspheric lens in the OCT beam at various distances X relative to the initial position of Figure 8A. For positive values ​​of X, the focus of lens L is below the reflection plane P. For negative values ​​of X, the focus of lens L is above the reflection plane P. X = 0 corresponds to the lens focus being on the reflection plane P, and therefore the image at X = 0 is a straight line.

[0026] Figure 11 shows an OCT image acquired at X = 5 mm. The expected sag Δ can be calculated as Δ = 5 mm × 59.4 μm / mm = 297 μm. The measured sag Δ from the OCT image was 310 μm. Therefore, the measurement accuracy is 310 / 297 = 1.044, or 4.4%. This accuracy is acceptable given the available measurement accuracy of the X-shift of the aspheric lens.

[0027] 12A and 12B are OCT images showing the apparent OCT image of the reflecting plane P for an optical configuration with high-order aberrations. The high-order aberrations were introduced by tilting and inverting lens L so that the OCT beam is incident on the flat surface of lens L (aspheric lens L is designed to illuminate an arcuate surface).

[0028] Referring to Figure 13, the concepts outlined above can be used to measure the refractive error of an eye using OCT images. For a typical eye, a 1 mm change in the eye's axial length causes a refractive error of 2.82 diopters. This 2.82 figure is used by various IOL (intraocular lens) calculators to select the appropriate IOL after cataract surgery. The following realistic example is used to estimate the resolution of an OCT-based refractive error measurement device. The angle of incidence of a peripheral ray in an 8 mm dilated pupil is α = arcsin(4 / 20) = 11.5°. A 1 mm error in axial length causes a sag Δ, as defined above: Δ = 1 mm * [1 / cos(11.5°) - 1] = 20.6 μm. In the vitreous, Δ = 20.6 μm corresponds to a distance of 1.34 * 20.6 μm = 27.6 μm, where 1.34 is the refractive index of the vitreous. In summary, a 2.82 diopter error causes an OCT image sag Δ = 27.6 μm (air units). Therefore, assuming a realistic axial resolution of 4 μm for OCT, the estimated refractive error resolution is 2.82 diopters * 4 / 27.6 = 0.41 diopters. With a technically feasible OCT resolution of 2 μm, the resolution can be approximately 0.2 diopters. By using appropriate image analysis and segmentation software, the accuracy of the Δ measurement can be an order of magnitude less than the OCT depth resolution. In this way, refractive error measurements can exceed 0.2 diopters.

[0029] Recently, so-called phase-resolved OCT has also been developed. The software for phase-resolved OCT calculates not only the Fourier amplitude component of the interference spectrum, but also the phase component. Phase-resolved OCT can have a depth resolution much better than 2 μm, and thus phase-resolved OCT-based systems can have much better diopter resolution.

[0030] The angle of incidence α and numerical aperture of the patient's eye are functions of measurements of the diameter of the patient's pupil and the position of the patient's pupil along the optical axis obtained using conventional techniques (including anterior chamber OCT). The angle of incidence α is a function of the distance between the pupil and the patient's retina, which is also easily measured using an OCT scanner. Alternatively, the diameter of the pupil can be measured using a surgical microscope.

[0031] Figure 14A shows the sag Δ of an eye that has only myopic defocus aberration, but no astigmatism or higher order errors that result in a parabolic shape resulting from the distance of the retina from the eye's focal point. Higher order aberrations of the eye manifest as deviations from the apparent parabolic shape of the retina, as shown in Figure 14B.

[0032] OCT scanners can be used not only to measure spherical refractive error, but also to measure the entire wavefront. For this purpose, multiple large-scale A-scans arranged in a raster or radial configuration can be used. A set of OCT images, each at a different cross-sectional plane of the patient's eye, is acquired, with each pixel in each image representing a measurement of the eye at X, Y, and Z coordinates within the patient's eye. Thus, the pixels representing the retina in the multiple images (e.g., the line pixels representing the retina, as discussed above) can be assembled to obtain a surface, i.e., a surface defined by an array of X, Y, and Z coordinates. This three-dimensional surface is the wavefront aberration map used in wavefront science. For example, the Z value (i.e., aberration) at each (X, Y) coordinate in the three-dimensional surface represents the phase delay value of the wavefront aberration map. If this surface curves upward, the patient's eye is myopic; if this surface curves downward, the patient's eye is hyperopic; and, as described above, the amount of sag of the surface indicates the degree of myopia or hyperopia.

[0033] The wavefront aberration map can be analyzed using the well-known Zernike, Fourier, or Seidel methods. The optical characteristics of the eye, such as modulation transfer function, point spread function, Strehl ratio, Snellen resolution, astigmatism, coma, spherical aberration, higher-order aberrations, lens prescription parameters, etc., can be derived from the wavefront aberration map, i.e., from the two-dimensional sag of the surface. The derivation of the optical parameters of the eye using the measured wavefront aberration map is a routine procedure for those skilled in the field of human ocular optics, and does not need to be described in detail here.

[0034] An OCT scanner can be used to measure refractive error using any of the techniques described herein by configuring the OCT scanner to generate a strictly parallel incident beam on the cornea. If there is some error in the parallelism of the OCT beam (e.g., caused by the OCT scanning system), a strictly parallel test surface can be measured with OCT. Non-parallelism of the incident OCT beam will appear as a deviation from a straight line on the OCT image of the test surface. This deviation is a calibration error and should be used during calculations to derive the correct wavefront error of the eye.

[0035] To measure the wavefront aberration map, the reflection plane P (see Figure 8A and Figure 8B) imaged by the OCT scanner should ideally be flat. Furthermore, a flat anatomical retinal layer must be selected to measure the ocular wavefront map. In the human eye, the ideal flat anatomical layer is the Bruch's membrane-retinal pigment epithelium (RPE) complex. Anatomical studies have shown that the presence of a pit does not affect the flatness of the RPE / Bruch's complex and can be used to measure sag Δ.

[0036] The following conclusions and applications result from the above approach: Using the techniques described herein, the refractive error of the entire eye can be obtained without needing to know the refractive index of the cornea, the refractive index of the lens, or any topographical and geometrical data of the surfaces. All that is required is to trust the validity of Fermat's principle, which has been shown to be valid for the past 350 years. The techniques described herein can be performed without performing OCT-based ray tracing. The methods are independent of ray tracing. Refractive error, measured in diopters, is approximately proportional to sag delta, which is approximately proportional to the square of the pupil diameter. Higher order aberrations appear as deviations from a parabolic sag surface of the apparent image of the retina. As will be described later, by applying certain modifications to the OCT beam scanner, OCT can be transformed into a multifunctional diagnostic tool. Anterior chamber OCT can be transformed into a multifunctional diagnostic tool by (a) strictly collimating the incident OCT beam and (b) increasing the OCT reference arm length by the distance between the pupil and the retina (approximately 20 mm), thereby shifting the imaging zone downward from the anterior chamber to the retina. An OCT scanner implementing the techniques described herein can be integrated into a surgical microscope used during ophthalmic surgery. In this way, the surgical microscope becomes an intraoperative wavefront aberrometer, such as ORA (Alcon Inc.) or Holos (Zeiss).

[0037] The techniques for measuring the refractive error of a patient's eye as described herein may be performed using a computing device. The computing device may receive one or more OCT images from an OCT scanner or may be incorporated into the OCT scanner itself. The computing device comprises one or more processing devices and one or more memory devices operably connected to the one or more processing devices. The one or more memory devices store executable code that, when executed by the one or more processing devices, causes the one or more processing devices to calculate the refractive error of the patient's eye as described above and any other calculations or measurements from the OCT images as described above.

[0038] The above description is provided to enable any person skilled in the art to practice the various embodiments described herein. Various modifications to these embodiments will be apparent to those skilled in the art, and the general principles defined herein may be applied to other embodiments. Accordingly, the claims are not intended to be limited to the embodiments shown herein, but are to be accorded the full scope consistent with the language of the claims.

Claims

1. an optical coherence tomography (OCT) imaging system configured to scan a collimated beam across the patient's eye; one or more processing devices; one or more memory devices coupled to the one or more processing devices, the one or more memory devices storing executable code that, when executed by the one or more processing devices, causes the one or more processing devices to: receiving one or more OCT images of the patient's eye from the OCT imaging system; identifying an apparent image of the retina of the patient's eye in the one or more OCT images; one or more memory devices for calculating the refractive error of the patient's eye from the shape of the apparent image; 1. An apparatus for performing ophthalmic measurements, comprising:

2. The executable code, when executed by the one or more processing devices, further causes the one or more processing devices to: determining that the patient's eye is myopic if the apparent image is curved upward; 10. The apparatus of claim 1, wherein if the apparent image is curved downward, the patient's eye is determined to be farsighted.

3. 10. The apparatus of claim 1, wherein the executable code, when executed by the one or more processing devices, further causes the one or more processing devices to calculate high-order aberrations of the patient's eye according to the shape of the apparent image.

4. 10. The apparatus of claim 1, wherein the executable code, when executed by the one or more processing devices, further causes the one or more processing devices to measure the refractive error of the patient's eye by calculating sag of the apparent image.

5. The executable code, when executed by the one or more processing devices, further causes the one or more processing devices to: obtaining an angle of incidence α corresponding to a pupil width of the patient's eye and a distance between the pupil and the retina; calculating refractive error as a function of said sag and said angle of incidence; 5. The apparatus of claim 4, wherein the refractive error of the patient's eye is calculated by:

6. The apparatus of claim 1 , further comprising a surgical microscope, wherein the OCT imaging system is integrated into the surgical microscope.

7. The apparatus of claim 1 , wherein the OCT imaging system is a phase-sensitive OCT imaging system.

8. The apparatus of claim 1 , wherein the OCT imaging system is a modified anterior chamber OCT imaging system.

9. 10. The apparatus of claim 1, wherein the OCT imaging system is a modified anterior chamber OCT having a reference arm length selected to image the retina of the patient's eye and configured to produce a parallel beam.

10. (a) the one or more OCT images comprise a single image, and the apparent image is a line in the single image representing the retina; (b) the one or more OCT images include multiple images of multiple cross sections of the patient's eye, and the apparent image is a surface defined by an array of X, Y, and Z coordinates corresponding to pixels in the multiple images representing the retina; The device of claim 1 , wherein the

11. an optical coherence tomography (OCT) imaging system configured to scan a collimated beam across the patient's eye; one or more processing devices; one or more memory devices coupled to the one or more processing devices, the one or more memory devices storing executable code that, when executed by the one or more processing devices, causes the one or more processing devices to: receiving a plurality of OCT images of the patient's eye from the OCT imaging system, each OCT image of the plurality of OCT images corresponding to one of a plurality of cross-sections of the patient's eye, such that each pixel in each OCT image of the plurality of OCT images represents an X, Y, and Z coordinate within the patient's eye; one or more memory devices for identifying pixels in the plurality of OCT images representing a retina of the patient's eye to obtain a surface corresponding to a wavefront aberration map of the patient's eye; 1. An apparatus for ocular imaging, comprising:

12. 12. The apparatus of claim 11, wherein the executable code, when executed by the one or more processing devices, further causes the one or more processing devices to calculate a refractive error of the patient's eye using the wavefront aberration map.

13. scanning a collimated beam across the patient's eye with an optical coherence tomography (OCT) imaging system to obtain a plurality of OCT images at a plurality of cross sections of the patient's eye; identifying pixels in the plurality of OCT images that represent a retina of the patient's eye to obtain a wavefront aberration map; calculating the refractive error of the patient's eye using the wavefront aberration map; 1. A method for performing optometric measurements, comprising:

14. (a) determining that the wavefront aberration map is upwardly curved; determining that the patient's eye is a myopic eye in response to (a); The method of claim 13 further comprising:

15. (a) determining that the wavefront aberration map is downwardly curved; determining that the patient's eye is a hyperopic eye in response to (a); The method of claim 13 further comprising:

16. 14. The method of claim 13, further comprising calculating high-order aberrations of the patient's eye according to the shape of the wavefront aberration map.

17. The method of claim 13 , further comprising calculating the refractive error by calculating a sag of the wavefront aberration map.

18. Obtaining an angle of incidence α according to a pupil width of the patient's eye and a distance between the pupil and a retina of the patient's eye; calculating the refractive error as a function of the sag and the angle of incidence α; 20. The method of claim 17, further comprising:

19. 14. The method of claim 13, further comprising calculating a refractive error of the patient's eye according to the shape of the wavefront aberration map when performing ophthalmic surgery using a surgical microscope, wherein the OCT imaging system is integrated into the surgical microscope.

20. The method of claim 13 , wherein the OCT imaging system is a modified anterior chamber OCT imaging system.

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