FULL-FIELD OCT METHOD AND SYSTEM FOR GENERATING AN IMAGING OF AN OCULAR FUNDUS

Abstract

The invention relates to a full-field OCT method for generating an imaging of an ocular fundus (31), in which short-coherent light (22) is emitted and split into an object beam path (25) and a reference beam path (24). The object beam path (25) is directed onto the ocular fundus (33). The reference beam path (24) and a portion of the object beam path (25) reflected by the ocular fundus (31) are directed onto an image sensor (32), such that an interference between the reference beam path (24) and the object beam path (25) occurs on the image sensor (32), wherein the reference beam path (24) impinges on the image sensor (32) at an angle deviating from the object beam path (25). Before impinging on the image sensor (32), the reference beam path (24) impinges on an optical correction element (27) in order to reduce a chromatic aberration within the reference beam path (24). Intensity information and phase information is determined from a capturing of the image sensor. A focus-adjusted image of the ocular fundus is calculated. The invention also relates to a system that is suitable for carrying out said method. Images of the ocular fundus can be captured without the beam path being previously adapted to the refractive power of the eye lens.

Claims

1-13. (canceled)

14. A full-field OCT method for generating an image representation of an ocular fundus (31), including the following steps: a. emitting short-coherent light (22); b. splitting the short-coherent light (22) into an object beam path (25) and a reference beam path (24), wherein the object beam path (25) is guided onto the ocular fundus (33); c. guiding the reference beam path (24) and part of the object beam path (25) reflected by the ocular fundus (31) onto an image sensor (32) such that interference between the reference beam path (24) and the object beam path (25) arises on the image sensor (32), wherein the reference beam path (24) strikes the image sensor (32) at an angle that deviates from the object beam path (25) and wherein, prior to the incidence on the image sensor (32), the reference beam path (24) strikes an optical correction element (27) in order to reduce a chromatic deviation within the reference beam path (24); d. ascertaining intensity information and phase information from a recording of the image sensor (32); e. calculating a focus-corrected image of the ocular fundus (31).

15. The full-field OCT method of claim 14, wherein the direction of the reference beam path (24) is deflected using a reflection element (26).

16. The full-field OCT method of claim 15, wherein the reference beam path (24) upstream of the reflection element (26) is separated from the reference beam path (24) downstream of the reflection element (26).

17. The full-field OCT method of claim 14, wherein a length of an optical path of the reference beam path (24) is alterable.

18. The full-field OCT method of claim 15, wherein the optical correction element (27) is disposed between a beam splitter (23) and the reflection element (26).

19. The full-field OCT method of claim 14, wherein the optical correction element (27) is embodied as a transmission grating or as a reflection grating.

20. The full-field OCT method of claim 14, wherein the optical correction element (27) is set such that a normal of a pulse front of the reference beam path (24) emerging from the optical correction element (27) includes an angle with a propagation direction of the object beam path (25) that is smaller than an angle between a propagation direction of the reference beam path (24) and the propagation direction of the object beam path (25).

21. The full-field OCT method of claim 14, wherein a phase factor used when calculating the focus-corrected image is derived from a known refractive error of the patient.

22. The full-field OCT method of claim 14, wherein the image sensor (32) is designed to record an en-face section within a period of less than 800 s, preferably less than 500 s, further preferably less than 300 s.

23. The full-field OCT method of claim 14, wherein the optical elements (23, 28) disposed in the object beam path are rigid.

24. The full-field OCT method of claim 14, comprising a fixation light (39), which the patient sees during the recording.

25. The full-field OCT method of claim 24, wherein the fixation light (39) is a focus-independent fixation light.

26. A full-field OCT system comprising a recording device (15) and a computing unit (19), the recording device (15) comprising a light source (20) for emitting short-coherent light (22), a beam splitter (23) for splitting the short-coherent light (22) into an object beam path (25) and a reference beam path (24), wherein the object beam path (25) is guided to an exit opening of the recording device (15), and comprising an image sensor (32), on which the reference beam path (24) and part of the object beam path (25) reflected by an object are made to interfere, wherein the reference beam path (24) strikes the image sensor (32) at an angle deviating from the object beam path (25) and wherein the reference beam path (24) strikes an optical correction element (27) prior to the incidence on the image sensor (32) in order to reduce a chromatic deviation within the reference beam path (24), and wherein the computing unit (19) is designed to calculate a focus-corrected image of the object from image data recorded with the image sensor (32).

Description

BRIEF DESCRIPTION OF THE DRAWINGS

[0039] In the following, the invention is described in exemplary fashion on the basis of advantageous embodiments, with reference being made to the attached drawings. In detail:

[0040] FIG. 1: shows a schematic illustration of a full-field OCT system according to the invention;

[0041] FIG. 2: shows an illustration of the beam paths of the hand-held device of FIG. 1;

[0042] FIG. 3: shows a schematic illustration of the calculation of a focus-corrected image;

[0043] FIG. 4: shows another aspect of the hand-held device of FIG. 2; and

[0044] FIG. 5: shows an embodiment of a focus-independent fixation light.

DETAILED DESCRIPTION

[0045] In the full-field OCT system according to the invention shown in FIG. 1, a patient 14, whose ocular fundus is the subject of imaging, holds a hand-held device 15 in the hand. The hand-held device 15 is disposed in front of the eye 16 of the patient 14 such that the patient 14 can see a fixation light in the interior of the hand-held device 15. Once the patient 14 has set their viewing direction on the basis of the fixation light, they actuate a switch 17 that is used to trigger the recording of an image representation of the ocular fundus.

[0046] Once the recording has been completed, the image data are transferred via a data network 18 to a central computer 19 that is distant from the patient 14. By way of example, the central computer 19 can be disposed at the headquarters of a service provider who operates the full-field OCT system. The system can be configured such that the central computer 19 receives image data from a multiplicity of hand-held devices 15, which are operated at different locations or in different installations. In particular, the computing steps required for a possibly required focus correction can be carried out on the central computer 19. In alternative embodiments of the invention, the hand-held device 15 itself is designed to carry out the computing steps for the focus correction.

[0047] According to FIG. 2, the hand-held device 15 comprises a superluminescent diode 20, which emits short-coherent light with a coherence length of 10 m, for example. A first lens 21 is used to generate a collimated beam path 22 of the short-coherent light, which is incident on a beam splitter 23. The beam splitter 23 is used to split the short-coherent light 22 into a reference beam path 24 and an object beam path 25.

[0048] The reference beam path 24 is incident on a roof prism 26, enters the prism 26 through a roof face, is reflected at a hypotenuse face of the prism 26 and emerges from the roof prism 26 again in the opposite direction, through the other roof face but with a parallel offset. The reference beam path 24 is incident on a transmission grating 27 and guided back in the direction of the beam splitter 23 by the transmission grating 27.

[0049] Proceeding from the superluminescent diode 20, the object beam path 25 passes through the beam splitter 23 and is guided to the eye 16 of the patient via a second lens 28 and a stop 29. The object beam path 25 passes through the lens of the eye 30 into the interior of the eye 16 and illuminates the ocular fundus 31 over an area.

[0050] Components of the short-coherent light 22 cast back by the ocular fundus 31 return to the beam splitter 23 via the lens of the eye 30 and the second lens 28, the object beam path being deflected in the direction of an image sensor 32 at said beam splitter. The optical elements in the object beam path 25 are disposed in such a way that the ocular fundus 31 is imaged onto the image sensor 32. This is illustrated in FIG. 2 using the example of an object point 33, which corresponds to an image point 34 on the image sensor 32.

[0051] The reference beam path 24 is overlaid on the object beam path 25 between the beam splitter 23 and the image sensor 32. There is interference between the object beam path 25 and the reference beam path 24 in the plane of the image sensor 32. The interference pattern is recorded using the image sensor 32.

[0052] There is an angle between the object beam path 25 and the reference beam path 24 upon incidence on the image sensor 32. The object beam path 25 strikes the image sensor 32 at right angles. When striking the image sensor 32, the reference beam path 24 includes an angle of slightly less than 90 with the image sensor 32, for example an angle of 87.

[0053] The direction of the reference beam path 24 is set by the transmission grating 27. Before passing through the transmission grating 27, the reference beam path 24 propagates in a direction that is at right angles to the image sensor 32. The reference beam path 24 changes its propagation direction when passing through the transmission grating 27. Here, the transmission grating 27 is designed in such a way that a chromatic deviation is avoided. Expressed differently, the transmission grating 27 is designed in such a way that the pulse front of the reference beam path 24 emerging from the transmission grating 27 is parallel to the image sensor 32. Accordingly, the pulse front is not at right angles to the propagation direction of the reference beam path but includes a different angle with the direction of the reference beam path 24. The transmission grating 27 forms an optical correction element within the meaning of the invention, by means of which a chromatic deviation within the reference beam path 24 is reduced.

[0054] This avoids smearing in Fourier space of the multispectral wave field emanating from an object point 33. This also applies if, unlike what is shown in FIG. 2, the object beam path 25 is not focused on the image sensor 32 but has a focusing error. Accordingly, the focus-corrected image can be calculated without the result being impaired by chromatic smearing.

[0055] FIG. 2 shows an object beam path 25 in which the object point 33 is imaged to infinity by the lens of the eye 30. Thus, the eye 16 does not have a refractive error. The hand-held device 15 is adjusted in such a way that there is focused imaging onto the image sensor 32 when the eye 16 does not have a refractive error.

[0056] A focusing error arises in the case of a refractive error of the eye 16, within the scope of which the image plane does not correspond to the image sensor 32 but is located in a plane in front of or behind the image sensor 32. An image recorded by the image sensor 32 appears blurred.

[0057] The steps for correcting the focusing error are explained on the basis of FIG. 3. Each object point 31 of the ocular fundus 33 can be considered to be the starting point of a divergent spherical wave 36. The divergent spherical wave 36 is converted into a convergent spherical wave 37 by the lens of the eye 30 and the second lens 28, which are illustrated together as a lens component 35 in FIG. 3. In the case of a lens of the eye without a refractive error, the convergent spherical wave 37 opens into an image point 34 on the image sensor 32.

[0058] If the eye 16 has a refractive error, the convergent spherical wave does not have the correct angle of curvature following the lens component 35 and the image point lies either in front of the image sensor 32 or behind the image sensor 32. In FIG. 3, this is illustrated on the basis of a convergent spherical wave 38, which is curved too strongly and the image point of which is located in front of the image sensor 32.

[0059] In order to generate a focus-corrected image of the ocular fundus 33 in the case of such a focusing error, the washed-out image data recorded by the image sensor 32 are initially transmitted to the central computer 19. The central computer 19 subjects the image data to a Fourier transform so that a suitable phase factor can be added in Fourier space. The spectrum altered thus is then transformed back into real space, where a refocused (sharp) image of the object is obtained. Since the reference beam path 24 was chromatically corrected by the transmission grating 27, the correction calculation can be performed together for all wavelengths contained in the short-coherent light 22.

[0060] FIG. 4 illustrates the fixation light 39 of the hand-held device 15. The lens of the eye 30 images the object point 33 from the ocular fundus 31 to infinity, and so the eye does not have a refractive error. The fixation light 39 is disposed in such a way that it is perceived in focus by an eye 16 without a refractive error.

[0061] FIG. 5 shows an embodiment of a fixation light 39 which can be perceived in focus at different axial positions. The light emitted by a light source 40 is guided into a hollow cylinder 41, which is dimensioned in such a way that each light ray is reflected precisely once in the interior of the hollow cylinder 41. An axial section of the length 42, in which the light can be perceived as a sharply delimited spot, arises on the other side of the hollow cylinder 41. This is an example for a focus-independent fixation light 39, which can also be perceived in focus by a patient with a refractive error without needing an adjustment of optical elements within the hand-held device 15.