Optical coherence tomography for measurement on the retina

Abstract

Optical coherence tomograph for examining an eye (3) which features: - a lighting device (4, 5) for providing source radiation, - an illumination and measurement beam path (7) comprising a splitting element (6) for splitting the source radiation into illumination radiation (B) and reference radiation (R), with the illumination radiation (B) illuminating an illumination field in the eye (3) and collecting backscattered illumination radiation in the eye (3) as measurement radiation (M), wherein the illumination and measurement beam path (7) comprises a scanner (13) for adjusting the lateral position of the illumination field in the eye (3) and a front optic (12), - a reference beam path (8) which provides for the reference radiation (R) an optical path length (21) corresponding to an optical path length from the splitting element (6) to the illumination field and back to a superposition point (71), - a detection beam path (14, 15, 17) which receives the measuring radiation (M) from the illumination and measuring beam path (7) and the reference radiation (R) from the reference beam path (8) and superimposes them at the superposition point (71) and directs them onto a detector (19), - the illumination and measuring beam path (7) further exhibits - a beam splitter (11) for separating the measuring radiation (M) collected by the eye (3) from the illumination radiation (B) directed to the eye (3), wherein the beam splitter (11) directs the separated measuring radiation (M) to the detection beam path (14, 15, 17), and - a light splitting element (32, 34) that splits the illumination radiation (B) into spots in order to illuminate the retina (2) with a multi-spot pattern, characterized in that - the lighting device (4, 5) is tunable with respect to the wavelength of the source radiation, - the detector is an area detector (19), - the detection beam path further - an optical element (14) acting only on the measuring radiation (M), which interacts with the front optics (12) and adjusts the numerical aperture with which measuring radiation (M) is collected in the eye (3), and - an aperture (15) which is arranged in front of the area detector (19), in or near an intermediate image plane and which defines the size of an object field from which the measuring radiation (M) reaches the area detector (19), and - wherein the aperture prior to the area detector is designed as a first multi-hole aperture (15) and a first multi-lens array (36) is arranged between this multi-hole aperture and the area detector (19), which focuses the radiation emanating from each hole of the first multi-hole aperture (15, 15a, 15b) onto a pixel area of ​​the area detector (19) which has a spatial resolution of 4 to 100 pixels in one direction, preferably as a 2D pixel area with 5 to 50 pixels or 5 to 40 pixels per direction.

Description

[0001] The invention relates to an optical coherence tomograph for examining an eye, comprising an illumination device for providing source radiation whose wavelength is tunable, an illumination and measurement beam path which has a splitting element for dividing the source radiation into illumination radiation and reference radiation, which illuminates an illumination field in the eye with the illumination radiation and collects backscattered illumination radiation in the eye as measurement radiation, wherein the illumination and measurement beam path has a scanner for adjusting the lateral position of the illumination field in the eye and a front optic, a reference beam path which provides an optical path length for the reference radiation that corresponds to an optical path length from the splitting element to the illumination field and back, and a detection beam path.which receives the measurement radiation from the illumination and measurement beam path and the reference radiation from the reference beam path and superimposes them onto a surface detector.

[0002] The invention further relates to a method for optical coherence tomography for examining an eye, wherein source radiation is provided and tuned with respect to its wavelength and split into illumination radiation and reference radiation, an illumination field in the eye is illuminated with the illumination radiation, illumination radiation backscattered in the eye is collected as measurement radiation, wherein the lateral position of the illumination field in the eye is adjusted with a scanner, the reference radiation is delayed in a reference beam path and superimposed with the measurement radiation to generate an interference signal which is detected with an area detector.

[0003] Optical coherence tomography (OCT) is an established method in ophthalmology for imaging the eye. It allows for three-dimensional imaging, which is very helpful in diagnosing eye diseases and monitoring their progression. Retinal diseases, such as glaucoma and age-related macular degeneration, are particularly relevant in this context. In OCT systems, the lateral resolution (x and y) is determined by the numerical aperture (NA) of the optics used. The axial resolution (z), on the other hand, is calculated from an interference pattern and is generally much greater than the depth of field of the image, which in turn depends on the numerical aperture, more precisely proportionally to 1 / NA. 2 In the commonly used Fourier-domain OCT, which employs a broadband or wavelength-adjustable radiation source, the depth resolution is inversely proportional to the spectral bandwidth, more precisely proportional to λ. 2 / Δλ, where λ is the mean wavelength and Δλ is the bandwidth.

[0004] Measuring the retina of the human eye requires both high lateral and high axial resolution. Simultaneously, the detectable and thus illuminated volume in depth (along the optical axis) should be as large as possible; this necessitates a small numerical aperture (NA) of the optical system. Lateral resolution requires a large numerical aperture. Therefore, in the prior art, the extent of the accessible depth and the lateral resolution are ultimately linked via the numerical aperture of the optical system and cannot be set independently.

[0005] US Patent 2014 / 0028974 A1 discloses an imaging technique based on optical coherence tomography (OCT). In this technique, a line is projected onto an object by an imaging system. The backscattered radiation is combined with reference radiation via interference and directed to a detector, where confocal filtering is performed in one direction. Astigmatic optics are used for this purpose. Depth resolution is achieved using optical coherence tomography. In the case of spectroscopic analysis of the radiation, a two-dimensional detector is used. One dimension of this detector serves for confocal filtering with respect to the line-shaped illuminated area, while the other dimension resolves the spectral information. The correlation between lateral resolution and accessible depth range is also present in the approach according to US Patent 2014 / 0028974 A1.

[0006] WO 2014 / 0179465 A1 describes an OCT that operates in the spectral domain, meaning it analyzes the interference of radiation with a spectrometer. The light source emits a beam of light consisting of numerous parallel individual beams, which are focused onto the sample by the objective lens. The reference arm also guides several parallel individual beams, so that ultimately each individual beam is guided through the device according to the OCT measurement principle and also analyzed by the spectrometer. This device is very complex to adjust.

[0007] US patent 2013 / 0003077 A1 discloses an optical coherence tomography (OCT) scanner for measuring the eye at three locations simultaneously. Radiation from a laser source is directed onto the eye at these three locations, superimposed with individually delayed reference radiation for each location, and then analyzed in a spectrometer. This is known as the standardized optical coherence tomography (SD-OCT) principle.

[0008] DE 102010055350 A1 describes an optical coherence tomograph for measuring the length of the eye and for measuring the anterior segment of the eye.

[0009] The DE 102005058220 A1 patent deals with capturing different axially spaced sections of the eye using SD-OCT technology.

[0010] In a scanning OCT system, the accessible diameter of the eye's pupil is typically between 1 mm and 1.5 mm. This results in a lateral resolution of approximately 15 µm and a depth of view of 3 mm. A higher numerical aperture of the optical system would improve the lateral resolution, but this would simultaneously reduce the depth of view. Furthermore, aberrations increase with increasing numerical aperture. While defocusing, a higher aberration, can usually be neglected in known OCT systems that utilize pupil diameters of up to 1.5 mm, astigmatism and coma increase with larger pupils. Therefore, diffraction-limited resolution cannot be achieved.

[0011] For certain applications, particularly for diagnosing age-related macular degeneration (AMD), high lateral resolution is desirable. To detect the early stages of this disease, a lateral resolution of approximately 5 µm is required. Simultaneously, a scannable depth of approximately 3 mm is necessary, as AMD is thought to involve blood vessel formation in deeper tissue layers. Furthermore, a good signal-to-noise ratio is required to detect such vessels.

[0012] The invention is based on the objective of providing an optical coherence tomograph for measurements on the retina of the human eye, in which the lateral resolution is improved without simultaneously limiting the accessible depth range.

[0013] The invention is defined in claims 1 and 9. Advantageous embodiments are the subject of dependent claims 2 to 8 and 10 to 15.

[0014] The invention combines several features to obtain a three-dimensional image using optical coherence tomography, which has a particularly good resolution laterally, i.e., perpendicular to the optical axis, and at the same time can cover a very large depth range axially, i.e., along the optical axis, without having to adjust focusing elements or lenses during the measurement process.

[0015] The invention implements multi-spot holoscopy. One aspect of the invention is that the object is illuminated and imaged simultaneously at a multitude of object spots, with the imaging occurring in parallel by filtering the radiation during detection using a multi-hole aperture. Each object spot is imaged onto a detector spot, the intensity distribution of which is recorded. Preferably, oversampling is used, meaning that the resolution of the intensity distribution of each detector spot corresponding to each object spot is higher than the resolution of the optical imaging of the illumination would normally allow. In preferred embodiments, an image correction is then derived from the intensity distribution.

[0016] The radiation from the imaged object spots is superimposed with reference radiation according to the OCT principle, so that the detector spots ultimately record an interference signal between measurement radiation and reference radiation.

[0017] For this purpose, an area detector is used that scans object spots on the retina. The multi-hole aperture of the optical imaging defines these object spots, and the area detector is calibrated to the overall size of the area covered by the aperture with holes and, in terms of its spatial resolution, to the size of the holes.

[0018] The beam path is preferably designed such that illumination with illuminating radiation and the collection of backscattered measurement radiation can be performed with different numerical apertures. This allows a numerical aperture to be set for illumination that illuminates a large axial area, so that the collected measurement radiation originates from a comparatively large depth range and, consequently, an image over a large depth range is obtained using the OCT principle. The numerical aperture for collecting the measurement radiation, i.e., for imaging an object region, is then independent of the numerical aperture of the illumination and can therefore be set larger, for example. This combines high lateral resolution with a large illuminated depth range.

[0019] The area detector is a two-dimensional detector. The pixel count ranges from 4 to 100 pixels per direction and per aperture of the multi-hole aperture, preferably between 5 and 40 pixels. This pixel count proves advantageous for scanning each object spot in terms of resolution, signal-to-noise ratio, and potential image error correction.

[0020] One image error correction technique of particular importance is the correction of aberrations produced by the eye. Since the numerical aperture of illumination and detection are decoupled, it is possible to perform detection—that is, the imaging of the object area onto the retina—with a very high numerical aperture, so large that eye aberrations play a significant role. The spatial resolution provided by the area detector for each object spot allows, as explained below, correction of these aberrations when the area detector is located in a conjugate pupil plane of the image. If the area detector is not located in a pupil plane, aberration correction is equally possible if the detected signal is converted to a pupil plane, as is known for holograms in the prior art.

[0021] In the object plane and image planes of a beam path, the image information is purely spatial. Imaged structures are also found as intensity differences in intermediate image planes. In the pupil plane, the image information is purely angular. The angles of the incident rays encode the image information. This is the well-known effect that a change in the cross-sectional area of ​​a pupil affects only the image brightness, not the image size. For this reason, in the human eye, the iris lies in the pupil plane, so that by constricting or dilating the iris, the human eye adjusts with respect to brightness. Whenever this description refers to the plane of the pupil of the eye, it means the iris plane. An imaging beam path maps an object from the object plane onto an image in the image plane (e.g., the location of a detector).Between, for example, the object plane and an intermediate image plane, a pupil is always present due to the laws of imaging. Similarly, an intermediate image plane always exists between two pupil planes. Likewise, in this description, planes located between the plane of the eye's pupil and the detector are referred to as conjugate pupil planes, since they are conjugate to the plane of the eye's pupil by the optical imaging elements. The fact that the retina is mentioned here as the object is not intended to limit the invention. Other structures of the eye can also be imaged as objects.

[0022] The invention combines the advantages of a confocal scanning system with those of a spatially scanning detector. The confocal principle of a scanning system lies in the highly effective suppression of scattered radiation, resulting in a measurement signal with a high signal-to-noise ratio. Simultaneously, the lateral resolution can be increased by enlarging the aperture at the eye. The invention provides that the numerical aperture of the illumination is decoupled from the numerical aperture of the detection. This enables high lateral resolution without compromising the detectable depth range. The conflicting design objectives of the prior art (high lateral resolution requires a high numerical aperture, a large detectable depth range requires a small numerical aperture) are thus resolved.

[0023] The embodiments of the invention preferably use a confocal multi-hole aperture. In this description, the term "confocal" refers not only to an aperture that lies exactly in an intermediate image plane conjugate to the object plane, but also encompasses an arrangement of the aperture that lies within a certain error range in front of or behind an intermediate image plane. If the confocal aperture is not located exactly in the intermediate image plane, but near it, stray light suppression may be reduced, but the function as a confocal aperture, which defines the object field from which the measurement radiation is collected, is equally fulfilled. The aperture is in or near an intermediate image plane if it is spaced from the intermediate image plane by a maximum of three times the image depth; a spacer of at most one image depth is preferred.The depth of focus, also known as the image depth, defines an axial region in image space, i.e., at the intermediate image plane of an optical system, where a sufficiently sharp image is formed in every focal plane. Within this depth of focus, circles of confusion are registered as points. The region in object space conjugate to the depth of focus is the depth of field. The depth of field is a measure of the extent of the sharp area in object space and is given by λ / (NAo). 2 Given, where NAo denotes the numerical aperture in object space. The image depth at the intermediate image plane is obtained analogously to the depth of field from the numerical aperture by Lambda / (NAz). 2NAz is the numerical aperture at the intermediate image plane, which can be calculated, for example, from NAo using the image scale. In the above consideration, the wavelength can be considered to be the maximum wavelength of the measurement radiation at the intermediate image plane.

[0024] The invention has the advantage that higher light intensities can be directed into the human eye because they are distributed over a larger area in the anterior chamber. Therefore, the illumination beam path includes a structure that illuminates the retina with a multi-spot pattern, thereby shifting the pupil of the illumination into the eye and preferably illuminating the full diameter of the iris (approx. 4.2 mm). It is particularly preferred to shift the effective position of the illumination pupil directly into the plane of the iris. This can be achieved by a field lens, or alternatively by a corresponding design of the optics between the eye and a multi-spot aperture of the illumination beam path.

[0025] A pupil position outside the eye would mean that the illumination beam would not pass through the iris on the optical axis for all spots. This results in a small angle between illumination and detection, which can lead to vignetting effects across the retinal measurement depth. These effects can be detrimental for high-precision applications.

[0026] The invention generates multiple illumination spots on the object, e.g. the retina, from a homogeneous, planar, extended illumination wave without losses, and simultaneously utilizes all pixels on the matrix-shaped area detector on the detection side.

[0027] Several principles are suitable for detection. Detection with a single detector, balanced detection, or off-axis detection are all possible.

[0028] The invention will now be explained in more detail, for example with reference to the accompanying drawings, which also reveal essential features of the invention. They show: Fig. 1 a schematic representation of an optical coherence tomograph (OCT) in a first embodiment, Fig. 2 a schematic representation of an OCT in a second embodiment, Fig. 3a-d Schematic representations of different variants of the detection beam path of the OCT of the first or second embodiment, Fig. 4 illustrations demonstrating aberration correction performed during one of the OCT scans Fig. 1 or Fig. 2 can be used, Fig. 5 A top view of a detector used in one of the OCTs of the Fig. 1 or Fig. 2 can be used, Fig. 6. A diagram illustrating a depth correction performed during one of the OCT scans. Fig. 1 or Fig. 2 can be applied, Fig. 7 Signal intensities of different channels of a detector of an OCT according to one of the Fig. 1 or Fig. 2, Fig. 8 a representation similar to the Fig. 7, Fig. 9 A schematic representation to illustrate the scanning principle of an OCT according to one of the Fig. 1 or Fig. 2, and Fig. 10 and Fig. 11 schematic representations similar to the Fig. 9 to illustrate the creation of a three-dimensional image.

[0029] Fig. Figure 1 shows an OCT 1 that acquires three-dimensional images of a retina 2 of an eye 3. Source radiation from a wavelength-tunable radiation source 4, for example, a suitable laser, is coupled into a fiber 5. The source radiation lies, for example, in the infrared wavelength range. In the following description, this wavelength range is also referred to as "light." This term encompasses all radiation of the electromagnetic spectrum that obeys the laws of optics.

[0030] Fiber 5 terminates in a splitter 6, which divides the source radiation into a measuring arm 7 and a reference arm 8. A fiber 9 connects to the splitter 6 in the measuring arm 7, and the illumination radiation B exiting the fiber end is guided by an illumination optic 10 to a beam splitter 11. From there, it passes to a front optic 12, which focuses the illumination radiation B onto a focus located on the retina 2 of the eye 3. The illumination optic 10 and the front optic 12, among other things, set the numerical aperture NA with which the eye 3 is illuminated. A scanner 13 is located between the beam splitter 11 and the front optic 12, deflecting the focus on the retina 2 laterally in two axes perpendicular to the direction of incidence. The coordinates of this deflection are subsequently denoted by x and y. The z-position of the focus can be adjusted by moving the front optic 12. This is schematically indicated by a double arrow.

[0031] The illumination radiation at the illumination focus on the retina 2 is backscattered from various depths within the depth of field. This depth of field is defined by the numerical aperture NA, which is determined by the interaction of the front optics 12 and illumination optics 10, as well as the optical properties of the eye 3.

[0032] The backscattered radiation is collected by the front optics 12 as measurement radiation M. To distinguish the incident illumination radiation B and the backscattered measurement radiation M collected by the front optics 12, these are in Fig. 1. The illumination radiation is shown with solid lines in the figure, the measurement radiation M with dotted lines. The measurement radiation collected by the front optic 12 is directed to the scanner 13. Here it is scanned, so that after the scanner 13 the measurement radiation M is present as a stationary beam.

[0033] The collection of the measurement radiation M is an image of the retina 2. The beam splitter 11 separates the measurement radiation M from the illumination radiation B and directs it to a detector device 17. The detector device 17 will be described later with reference to the Fig. 3 is explained in more detail. It features, among other things, optics which, together with the front optics 12 and the optical properties of the eye 3, determine the numerical aperture NA of the image of the retina 2. In this way, illumination and detection have different numerical apertures. The numerical aperture of the illumination is determined by the combination of the illumination optics 10 and the front optics 12, the numerical aperture of the detection by the front optics 12 and the detector assembly 17.

[0034] Reference radiation R from reference arm 8 is also coupled into detector device 17. This arm has a fiber 20 after splitter 6. Reference arm 8 has, in the case of the Fig. In the embodiment shown in Figure 1, a path length adjustment device 21 is used to adjust the length of the reference arm 8 to match the position of the retina 2 of the eye 3. For this purpose, the radiation is coupled out of the fiber 20 and guided via a retroreflector 22, the position of which can be adjusted, as indicated by the double arrow in Figure 1. Fig. 1 indicates. Via a further deflecting mirror 23 and optics 24, 25, the reference radiation R is directed to the detector device 17, which superimposes the reference radiation R with the measuring radiation M onto a surface detector 19.

[0035] The path length adjustment device 21 is in Fig. 1 is designed as a free beam path. This is optional, as is the use of a retroreflector 22. Various measures for adjusting the optical path length of a beam are known in the prior art.

[0036] The interference between reference radiation R and measurement radiation M is used to generate an image, as is known for optical coherence tomography. Since the wavelength of the source radiation is tuned, the Fourier domain principle, which is fundamentally known from the prior art, is applied during image generation.

[0037] To perform image generation, the OCT 1 has a control unit C, which receives a signal via the wavelength tuning and the measurement signals from the area detector 19. Optionally, the control unit C controls the radiation source 4 for wavelength tuning, thus knowing the currently dominant wavelength in the system and being able to assign the measurement signals accordingly. The area detector 19 receives measurement radiation M from an object field on the retina 2, which is filtered by an aperture in the detector assembly 17 (see Fig. 3) is defined. The area detector 19 is adapted in size to this aperture and scans the intensity distribution with spatial resolution using individual pixels. If the area detector 19 lies in an image plane, i.e., in a plane that, taking into account the imaging performed by the front optics 12, detector optics 14, and the other intermediate optical elements, is conjugate to the plane of the retina 2, the individual pixels already contain the spatial information in the object field. If, on the other hand, the area detector lies in a conjugate pupil plane, which is conjugate to the plane in which the pupil P of the eye 3 lies, the pixels detect the intensity distribution in the pupil plane and thus the angular information. This can also be used for image reconstruction, as will be explained below.

[0038] Essential to the invention is that the scanner 13 shifts the object field in the retina 2, since it acts not only on the illumination radiation B but also on the collection of the measurement beams M. At each position of the scanner 13, a partial image of the retina is thus created. These partial images are, as will be explained below, combined to form a complete image that has a significantly higher resolution than that known from wide-field OCT.

[0039] In the construction method of the Fig. 1. The detector assembly 17 combines the measurement radiation M from the measurement arm 7 and the reference radiation R from the reference arm 8. The area detector 19 detects the interference pattern between the measurement radiation M and the reference radiation R. The corresponding measures for generating such interference, in particular the necessary properties of the radiation source 4 and the path length matching, are known in the prior art for optical coherence tomographs.

[0040] The complex amplitudes of the measurement radiation and the reference radiation can be written as: Usample=us∗eiφs and Ureference=ur∗eiφr, if you are with u s and u r the amplitudes and φ s and φ r the phases of the signals in the two arms are designated (the indices “sample” and “s” refer to the measuring arm, the indices “reference” and “r” to the reference arm).

[0041] The detector detects a signal I1 and, in the case of "balanced detection", which will be discussed later, a signal I2: I1=|Usample+Ureference|2=|Usample|2+|Ureference|2+2Re{Usample*U¯reference}I2=|Usample+Ureference∗eiπ|2=|Usample|2+|Ureference|2+2Re{Usample∗U¯reference∗e−iπ}.

[0042] The signal amplitude of the interference signal is located on a common-mode component |U sample | 2 and |U reference | 2It is modulated and filtered out through appropriate data analysis, balanced detection, or off-axis detection.

[0043] Fig. Figure 2 shows another embodiment for the OCT 1, in which the path length adjustment device is arranged not in the reference arm 8, but in the measuring arm 7. A path length adjustment device 29 is located after the fiber 9 and the illumination optics 10, again purely by way of example by means of a movable retroreflector 30. The embodiment of Fig. Figure 2 shows that it is irrelevant whether the path length adjustment device is located in the reference arm 8 or in the measuring arm 7. It is also possible to provide a path length adjustment device in both. The essential point is that the interference state between the reference radiation R from the reference arm 8 and the measuring radiation M can be adjusted so that it is adapted to the current measurement task, i.e., the actual position of the object to be measured, in the embodiments described here, exemplified by the retina 2 of the eye 3.

[0044] In Fig. Figure 2 shows an option in which the front optics 12 are formed in two parts by two imaging elements 12a and 12b.

[0045] Regarding illumination brightness for applications on the eye (3), it is important to note that there are specifications for a maximum permissible radiation intensity at the cornea. Increasing the illuminated field allows more illumination radiation to be coupled to the eye (3) without exceeding a threshold for illuminance density. For ophthalmic devices, in the infrared radiation range in which OCT typically operates, there is a maximum luminance of approximately 1 mW / mm². 2 in the anterior chamber of the eye. If a pupil diameter of 4.5 mm is illuminated homogeneously, a total power of approximately 16 mW would be permissible. However, to prevent the depth of detection from becoming too small, not the entire pupil P of eye 3 is used for illumination.

[0046] Rather, a NA of approximately 0.035 (or a pupil diameter of 1.2 mm) is preferably used as the upper limit for usable depth perception.

[0047] For retinal tissue, the maximum permissible power is 1.5 mW for each spot smaller than 1.5 mrad and for a wavelength of 1060 nm. This means that the 16 mW permissible for the pupil must be distributed over an angle of 15 mrad in at least one direction to avoid exceeding the limit at retina 2. These values ​​would maximize the overall signal intensity, but at the expense of image contrast, since stray light is to be expected under normal wide-field illumination due to the high-intensity radiation.

[0048] In the OCT 1, this conflict of objectives is resolved by simultaneously illuminating and detecting the retina at several spaced-apart spots. The spacing of the spots minimizes the problem of stray light. Fig. Figures 3a-d schematically show the relevant elements of the OCT 1, whereby all components relating to the generation and coupling of the reference radiation R have been omitted for the sake of simplicity. Elements of the Fig. Figures 3a-d, which correspond to the elements already explained with reference to the previous figures, are labeled with the same reference symbols, so that their description need not be repeated. The coupling of the reference radiation R is shown in the representation of the Fig. 3a at the point of the dashed double line. The Fig. Figures 3b-d show different variants for the implementation of coupling and detection.

[0049] Illumination and detection take place in the OCT of the Fig. 1 and Fig. 2 in a multi-spot principle, as Fig. Figures 3a-d show this. However, this does not apply to reference beam path 8, whose radiation is not penetrated by multiple spots. The guidance of the reference radiation is shown in Fig. 3 is not shown. It is located at the point of the dashed double line, i.e., in the direction of the image, after the beam splitter 11 and before the in Fig. Optics 14 shown in 3a is coupled in. For the sake of simplicity, this coupling is shown in Fig. 3a not shown. Optics 14 is the one based on Fig. The optics of the detector assembly 17 mentioned above, which together with the front optics 12 and the optical properties of the eye 3 determine the numerical aperture NA of the retina 2. The optics 14 are therefore a detector optic.

[0050] The detector optics 14 focus the measurement radiation M into an intermediate image plane in which an aperture 15 is located. The aperture 15 determines the size of the object field from which the measurement radiation M is detected at the retina 2. Taking into account the image scale of the detector optics 14, the front optics 12, and the eye 3, the size of the aperture 15 corresponds exactly to the size of the object field at the retina 2 from which the measurement radiation M is collected.

[0051] As will be explained below, aperture 15 is designed as a multi-hole aperture, which, together with subsequent components (to be explained in more detail later), maps a multitude of object spots on the retina onto a corresponding number of detector spots on the area detector 19. The area detector is designed such that each detector spot has a spatial resolution of 4 to 100 pixels in one direction, preferably 5 to 50 or 5 to 40 pixels. The area detector thus scans each spot with respect to its intensity distribution using the individual detector areas. The significance of this scanning for the holographic OCT will be discussed below.

[0052] According to Fig. In steps 3a-d, the illumination radiation B provided at the end of fiber 9 is distributed into spots and coupled across the entire usable pupil P (diameter 4.5 mm) of eye 3. This ensures that both the power limits of the cornea of ​​eye 3 and the power limits of the retina 2 are respected. If the spots are configured so that they have a minimum spacing of 2 mm at the retina 2 (more generally, a spacing that roughly corresponds to the detectable depth range of the OCT), multiply scattered photons are suppressed just as effectively as in a confocal OCT that deflects a single spot. This eliminates the problem of stray light.

[0053] The illumination radiation B coupled out from fiber 9 is collimated by a collimator lens 31 and then focused onto a multi-hole aperture 34 by a multi-lens array 32 and a field lens 33. The multi-hole aperture 34 determines the pattern, spacing, and size of the illumination spots on the retina 2, as it lies in a plane that is conjugate to the object plane on the retina 2 due to the subsequent optics 35, 12a, 12b. The optics are preferably designed such that both the beam splitter 11 and the scanner 13 are located near a pupil of the illumination beam path. The optional field lens 33 in front of the multi-hole aperture 34 ensures that in the plane of the pupil P of the eye 3 the radiation is spread uniformly over the entire pupil P, i.e. over the diameter of 4.5 mm, so that there are no points where the permissible radiation intensity is exceeded.

[0054] The measurement radiation M backscattered at the retina 2 is focused by the front optics with optics 12a, 12b via the intermediate image plane 26 and the scanner 13 as well as the beam splitter 11, both of which are located near or in a conjugate pupil plane that is conjugate to the plane of the pupil P of the eye 3, and after coupling of the reference radiation (in the section between the dotted double line) from the detector optics 14 onto the area detector 19. The multi-hole aperture 15 in Fig. 3 is the one to Fig. The aperture 1 is referred to as the multi-hole aperture 15. A downstream multi-lens array 36 focuses the measurement radiation M from the individual aperture openings of the multi-hole aperture 15 onto the detector 19. The multi-lens array 36 ensures that the measurement radiation M, which passes through the individual aperture openings of the multi-hole aperture 15, does not mix before it falls onto the respective assigned pixels of the area detector 19. Thus, each hole of the multi-hole aperture 36 illuminates several pixels of the detector 19.

[0055] Fig. Figures 3b-d show possible variations for coupling the reference radiation R. The reference radiation R is exemplified by an optical fiber 70, which forms the end of the additional path length through which the reference radiation R is sent. Free-space guidance is equally possible for all variations.

[0056] According to Fig. 3b The reference radiation is superimposed on the measurement radiation in such a way that balanced detection is achieved. In the design of the Fig. 3c also performs balanced detection and in Fig. In 3D, off-axis detection is performed.

[0057] How Fig. As shown in Figure 3b, the measurement radiation is superimposed on the reference radiation R, which originates from the optical fiber 70 and is first split into a multitude of spots by a lens 72 and a multi-lens array 36c, with the focal plane of the multi-lens array 36c lying in an intermediate image plane 26. A multi-hole aperture 15c can optionally be located there. A corresponding optic 14c collimates the radiation and directs it to the beam superimposition device 71. This superimposes the measurement radiation M with the reference radiation R onto two detectors 19a, 19b, each of which is preceded by a corresponding optic 14a, 14b, a multi-hole aperture 15a, 15b, and a multi-lens array 36a, 36b. The multi-lens arrays 36a, 36b in front of the detectors 19a, 19b and the multi-lens array 36c are aligned with each other and with the multi-lens array 32 and the multi-hole aperture 34 of the illumination beam path.

[0058] The in Fig. The setup shown in Figure 3b can be used for all multispot systems of the invention equipped with balanced detection. The measurement radiation M is coherently superimposed with the reference radiation R near the pupil in the detection beam path and then imaged into two independent beam paths by the two area detectors 19a and 19b and detected. The multi-hole apertures 15a and 15b are conjugated to the multi-hole aperture 34 of the illumination beam path. The measurement radiation M passing through the multi-hole aperture 15a or 15b is collimated by the multi-lens array 36a or 36b and registered by the area detector 19a or 19b.

[0059] If the illumination beam path uses a pupil in the eye with a diameter of approximately 1.2 mm and the detection beam path uses a pupil with a diameter of 4.5 mm, the microlenses of the multilens arrays 36a-b in the detection section have a focal length 4.5 / 1.2 = 3.75 times shorter than the microlenses of the multilens array 32. The angular spectrum of the radiation at the area detectors 19a, 19b of the various spots then fills the sensor exactly, without any overlap or gaps. The magnification between the image plane of the retina 2 and the multi-hole apertures 15a, 15b of the detection section is chosen such that a desired number of pixels are detected for each individual spot generated by a microlens of the multilens array 36a, 36b, for example, ten pixels per spot. Detection takes place near the pupil, i.e. the area detectors 19a, 19b lie in a plane that is conjugate to the pupil plane P.The multi-hole apertures 15a, 15b, on the other hand, lie in an intermediate image plane 26 conjugate to the image plane (plane of the retina 2).

[0060] For the detection to be coherent, each measuring beam must be superimposed with a reference beam of the same aperture. This is achieved by collimating the reference radiation R, which emerges from the optical fiber 70, with the lens 72 and focusing it into the intermediate image plane 26 with the multi-lens array 36c. There, a reference wave is formed as a multi-spot pattern, which is imaged onto the multi-hole apertures 15a, 15b in the superimposed beam path of reference radiation R and measuring radiation M using the additional lens 14c and the lenses 14a, 14b. Preferably, the lens 14c forms a 4f arrangement with the lenses 14a and 14b, respectively.

[0061] If each spot is intended to illuminate a field of approximately 20 µm in diameter on retina 2, and these spots are to be spaced approximately 2 mm apart, the multi-lens arrays 36a and 36b utilize comparatively small effective field angles. In this case, it is not essential that the area detectors 19a and 19b are positioned directly in the focal planes of the microlenses of the multi-lens arrays 36a and 36b; they can also be located further away. This results in phase variations across the area detectors, which, however, can be numerically compensated after coherent detection.

[0062] If the distance between the microlenses of the multilens array 36a, 36b and the area detector 19a, 19b can be larger, this enables a particularly simple detection arrangement for balanced detection, which is described in Fig. Figure 3c shows that the reference radiation R is collimated with lens 72 so that the area detectors 19a and 19b are illuminated over their entire area. Only a single multi-hole aperture 15 with a multi-lens array 36, positioned upstream of the beam splitter 71, is then required.

[0063] Fig. Figure 3D shows an off-axis detection where there is sufficient space between the multi-lens array 36 of the detection beam path 17 and the area detector 19 to achieve superposition with the reference radiation R. In this case as well, if the distance between the multi-lens array 36 and the area detector 19 is greater than the focal length of the microlenses of the multi-lens array 36, corresponding phase variations across the area detector 19 can be numerically compensated during evaluation.

[0064] For the principle of off-axis detection, it is generally preferred to implement the multi-lens array 36 using anamorphic cylindrical lenses on the front and back sides of a thick, plane-parallel substrate layer. This arrangement, together with a rectangular arrangement of the microlenses in the multi-lens array 36, also makes it possible to illuminate the camera pixels of the area detector 19 without loss for off-axis detection, even if 2-3 times more pixels are required in the off-axis direction to image the same aperture values.

[0065] For off-axis detection, the angle to the optical axis must be adapted to various detection parameters. The smaller the angle, the greater the distance between the multi-lens array 36 and the area detector 19. Distances that are too large and angles that are too small mean that the resulting phase variations can no longer be adequately corrected numerically. Conversely, an angle that is too large leads to a loss of coherence in the superposition. The use of a TIRF prism as a beam splitter 71 represents a particularly good compromise. This prism consists of two glass prisms with a small air gap between them, which is schematically shown in Fig. Figure 3d shows the beam interpolation device 71. The glass prisms are dimensioned such that the measurement radiation M strikes the air gap at an angle of 45° and can still be transmitted, whereas the reference radiation R, due to the angular offset, strikes at an angle greater than 45° and is totally reflected, thus reaching the area detector 19 completely. The angle of 45° is only mentioned here as an example. Crucially, the critical angle for total internal reflection is not exceeded for the measurement beam, so that the measurement radiation M can pass through the TIRF prism, while the light from the reference strikes the glass-air interface at an angle larger than the critical angle, causing the reference light R to be reflected and therefore also reach the detector. The critical angle for prisms (with refractive index n(prism)>1) and an air gap (with refractive index n(air)=1) is given by Theta.Critical = arcsin(n(air) / n(prism)).Simultaneously, the angle of the reference light R on the sensor must of course correspond to the angle for off-axis detection. This is determined by the pixel width and the number of pixels across which a phase shift of 2π is distributed.

[0066] As explained above, in a pupil of the beam path, the image information is present in the form of angular information, and the intensity distribution in the pupil is generally completely uniform. It is therefore preferable to arrange optical elements that are intended to act uniformly on all structures to be imaged in a single pupil. Such elements include, for example, the scanner 13 and the beam splitter 11. However, it is not essential that these elements be arranged completely and exclusively in a conjugate pupil plane. In the design of the Fig. 3. For example, it is sufficient to arrange these elements in such a way that the rays emanating from adjacent holes of the multi-hole aperture 34 are already superimposed. This is the case when the corresponding marginal rays intersect.

[0067] It is likewise preferable to arrange lenses or other elements that can produce reflections as far as possible outside a conjugate pupil plane. Again, this requirement is not to be interpreted strictly. It is sufficient to arrange such elements in areas where the beams of rays from adjacent holes of the multi-hole aperture 34 have not yet begun to overlap, i.e., where their marginal rays have not yet intersected. In the case of the embodiment of Fig. 4 thus the term “near pupil” or “far from pupil” refers to the place along the optical axis where the marginal rays of adjacent holes of the multi-hole aperture 34 intersect.

[0068] Scanner 13 is located at OCT 1 of the Fig. 1 to 3 preferably in or near a pupil plane of the detection beam path as well as the illumination beam path. This pupil plane is conjugate to the plane of the pupil P of eye 3.

[0069] The front optics 12 optionally include, as exemplified by the embodiment of the Fig. Figure 2 shows optics 12a and 12b, which together form a 4f optic. Optic 12a is an ophthalmoscopic lens and optic 12b is a scanning lens. This 4f optic images the pupil P of eye 3 in a pupil plane conjugate to the plane of pupil P, in which scanner 13 is located. It is not essential to position scanner 13 exactly in this conjugate pupil plane, but doing so has advantages. An intermediate image plane 26 lies between the plane of pupil P of eye 3 and its conjugate pupil plane. Due to its proximity to scanner 13, beam splitter 11 is also located near the conjugate pupil plane. It is also possible to place the beam splitter 11 in this conjugate pupil plane if the scanner 13 is moved out of the conjugate pupil plane.

[0070] In one embodiment, the beam splitter 11 is designed as a polarization splitter. A lambda / 4 plate 27 is then arranged in front of it in the imaging direction (see figure). Fig. 2) This embodiment will be discussed in more detail below.

[0071] The detector optics are preferably also designed as 4f optics. They provide a further intermediate image plane 26 in which the aperture 15 is located. The intermediate image plane 26 is conjugate to the object plane in which the retina 2 to be imaged is located.

[0072] The aperture 15, 15a, 15b has two functions in all embodiments. Firstly, it suppresses stray light, thereby improving the contrast at the detector device 17.

[0073] Ultimately, in this respect, the aperture acts similarly to a confocal aperture for confocal OCT. Due to the effect of the detector optics, the area detector 19 is preferably located in a plane conjugate to, or near, the pupil plane of the eye. This arrangement is advantageous, but not essential. It has the advantage that the phase function of the electromagnetic field can be easily scanned. The maximum spatial frequency in the plane of the area detector 19 is determined by the object field size on the retina 2 and thus ultimately by the size of the aperture 15 in the intermediate image plane 26. The aperture 15 therefore also ensures particularly favorable signal acquisition.

[0074] In all embodiments of the OCT, the area detector has a pixel count of 4 to 100, preferably 5 to 50, particularly preferably 5 to 40 pixels in each direction per hole of the multi-hole aperture 15. Fig. Figure 5 shows a top view 43 of the detector, indicating that the pixel arrangement does not necessarily have to be rectangular, but that a hexagonal arrangement of the pixels is also possible. The pixel pattern is therefore freely selectable.

[0075] Holoscopic OCT systems with detectors featuring 100 to 4000 pixels per direction are known in the prior art. These pixel counts are deliberately not used here.

[0076] The number of pixels is linked to the required illumination brightness, the measurement speed, and the suppression of multiple scattering.

[0077] In a preferred embodiment of the OCT 1, aberrations are corrected. The pixels of the area detector 19 are hereinafter also referred to as channels. The measurement signal is distributed across these multiple channels. If, according to a preferred embodiment, the detector 19 is located in a conjugate pupil plane, each channel of the detector contains measurement radiation M from different angles, which has been scattered within the retina 2. The spatial resolution of the area detector 19 allows the distribution of the measurement radiation in the pupil P to be recorded for each spot. The following explanation refers only to one of these spots. Aberrations affect this distribution. Aberrations caused by the eye 3 often reach an intolerable level when an area larger than 1.5 mm in diameter is used in the plane of the pupil P of the eye 3. However, such a larger area would be desirable with regard to lateral resolution.Without spatial resolution in the conjugate pupil plane, phase differences would mix and average out in the single detection channel if the pupil were used more extensively in eye 3.

[0078] The corresponding Zernike polynomials that describe these aberrations are in Fig. Figure 4 shows top views 37 to 42 of a conjugate pupil plane. Also shown is the grid of a detector with 5 x 5 channels (or pixels) per spot. The pixels scan the pupil P and thus allow phase differences within the pupil P to be distinguished.

[0079] The maximum resolvable phase differences depend on the number of channels per spot. It was found that the number of distinguishable phase differences in this plane is the number of channels per direction multiplied by pi. With five channels per direction, as in Fig. As shown in section 4, polynomials up to 4 can be used per spot. Z4mA distinction is made where m can take the values ​​0 (sphere), ±2, and ±4. This applies to infinitesimally small channels in terms of area. In reality, of course, they have a definite size. The measurement signal captured in a channel therefore corresponds to an average of the interference signal over the area of ​​the respective channel (pixel area). The theoretically possible maximum order of the Zernike polynomial can thus only be achieved if the phase of the signal within a channel varies by less than Pi for each spot. It was shown that with an OCT mean wavelength of 1060 nm, the phase differences are discernible for uniformly spatially distributed channels for astigmatism caused by the eye, provided that the condition 2Pi / (5 channels per aberration period) ≤ Pi is met for five channels per spot. Then, one period of minima and maxima lies within the aperture.For higher orders: 0.6*2Pi / (5 channels per period of aberration) = 1.2*Pi / (5 / 1.5) ≤ Pi for the third order and 0.5*2Pi / (5 channels per period of aberration) = 1.0*Pi / (5 / 2) ≤ Pi for the fourth order.

[0080] These considerations show that an area detector with at least five channels per direction and spot is capable of resolving at least third-order astigmatism and aberrations. A higher number of channels allows for the detection of even higher orders of aberration.

[0081] The above considerations only took one spatial direction into account. How Fig. Figure 4 shows that aberrations usually have a two-dimensional pattern. Fig. Top view 37 shows the first-order aberration, also known as "piston"; top view 38 shows the "tilt" aberration; top view 39 shows the "tip" aberration; top view 41 shows the "defocus" aberration; and top views 40 and 42 show "astigmatism" aberrations. As can be seen, most aberrations have a two-dimensionally distributed pattern, which also means that the phase variation is two-dimensional. This pattern can be detected and corrected for each spot using the spatially resolved area detector 19.

[0082] The aberrations cause a phase for each detector channel c. θc:Usample,c:=Usample∗eiθc. It arises from the thickness δd and refractive index δn of the material of the eye (e.g. cornea, aqueous humor, lens, vitreous body) that passes through, which in reality differs from a theoretical, aberration-free eye: θc(k)=δn(k)∗k∗δdc

[0083] Thus, the detected signal is shifted by the aberration-induced phase: Idc,c(k)=4∗us∗cos(k∗Δz−δn(k)∗k∗δdc) =4∗us∗ur∗cos(k∗(Δz−δn(k)δdc))

[0084] For monochromatic radiation of 780 nm, the eye causes wavefront aberrations of up to 0.7 µm, leading to a phase shift of 2π (ignoring defocus). Such a phase shift corresponds to a thickness difference between the lens and the aqueous humor (the elements with the greatest refractive index differences in the eye), which takes the following value: δd=2Pi∗780nm2Pi∗δn(780nm)=780nmnlens(780nm)−naqueous(780nm)≈780nm1.415−1.334≈10μm.

[0085] Using known dispersion data, we obtain: θc(λ0=1060nm)=2Pi1060nm∗(nlens(1060nm)−naqueous(1060))∗δdc =2Pi1060nm∗(1.4104−1.3301)∗10μm=1.516Pi or θc(k0)=k0∗0.8034 μm.

[0086] When a wavelength range of Δλ = 50nm is traversed, the phase differences of the corresponding wavenumbers (k0 ± Δk) are: θc(k0+Δk)=2Pi1110nm∗(nlens(1110nm)−naqueous(1110nm))∗δdc=2Pi1060nm∗(1.4099−1.3297)∗10μm=1.445Pi=(k0+Δk)∗0.8022μm and θc(k0+Δk)=2Pi1010nm∗(nlens(1010nm)−naqueous(1010nm))∗δdc=2Pi1060nm∗(1.4098−1.3305)∗10μm=1.594Pi=(k0−Δk)∗0.8048μm.

[0087] These calculations show that, to a sufficiently accurate approximation, the phase shifts caused by the aberrations vary linearly with the wavenumber k within a wavelength tuning range. Therefore, the detected measurement signal can be written as follows: Ibd,c(k)=4∗us∗ur∗cos(k∗(Δz−δn(k0)δdc)).

[0088] A Fourier transform for the measured wavenumbers k reveals the axial distribution, i.e., the distribution in the z-direction for the scattering tissue. Compared to an aberration-free system, the axial distribution is shifted by the value δn(k0)δdc for each channel c of the area detector. A corresponding simulation example shows Fig. Figure 8, in which the z-coordinate is plotted on the abscissa and the signal intensity on the ordinate. Curves 55 to 58 correspond to four channels c of the area detector. It can be assumed that in most areas of the tissue, the variation of the axial scattering profile within a pupil size of 5 mm of the eye 3 is small. Therefore, the profile differences for channels c mainly result from aberrations that shift the profile axially. It is therefore intended to determine the aberration-induced phases Θ. c(k0) from the channels relative to a central channel (for example, the channel located in the center of the detector, which corresponds to a perpendicular incidence on the sample). The measured intensities for a frequency determination are multiplied by the phase factor to correct for aberrations. The phase factor is e -iθc(k0) .

[0089] Each detector channel has a specific orientation relative to the retina. The interference signal can be recorded during wavelength adjustment of the laser for the respective wavenumber k = 2πn / λ, where n is the refractive index of the medium and λ is the wavelength. As in a classical OCT system, the measurement signals are Fourier-transformed with respect to the wavenumbers, and the depth distribution of the scattering layers is calculated. The relationship Δφ = kΔz is used, where Δz is the distance of a scattering layer to a layer from which the measurement radiation traveled a path length to the detector that is identical to the path length of the reference radiation.

[0090] However, due to the lateral extent of the area detector 19 per spot, the optical path length for the individual pixels of a spot is not identical, as Fig. Figure 6 shows that for the area detector 19, the five exemplary pixels or channels 46, 47, 48, 49, and 50 for each spot differ with respect to the optical path length to a specific point in the tissue 44 scanned by the spot. The wavefronts for the central channel 48 are shown with solid lines. They are perpendicular to the optical axis of the beam path between the point under consideration, which is shown at the very bottom of the structure 44, and channel 48. For this central channel 48, the radiation travels along the optical axis. For a more outwardly located channel, for example, channel 50, the principal beam travels at an angle α to the optical axis, so that the path length, which has the value d for the central channel 48, has the value d*cos(α) for the outer channel 50. The corresponding wavefronts and principal beams are shown in Fig. 6 for the outer channel 50, shown with a dashed line. Further details are in Fig. Figure 6 shows the eye lens 45 as an example. The depth is referenced to the principal plane of the eye lens, since its refractive index step can be used as a reference point during measurement. How Fig. As Figure 6 clearly shows, pixels / channels located further out collect radiation that has traveled a longer path through the medium. This has an effect on how the image information is reconstructed. Fig. Figure 7 shows an example. Signal curves 51 to 54 for four channels of a spot are shown there. The plot corresponds to that of the Fig. 8, i.e., the depth coordinate z is plotted on the abscissa, and the intensity on the ordinate. As can be seen, the individual curves are not only shifted in the z-direction, but they are also compressed for pixels located further out. Curve 54 is the measurement signal of the central pixel 48, and curves 53, 52, and 51 are measurement signals from channels located further out.

[0091] In a preferred embodiment, the measurement error caused by this effect is corrected to obtain a particularly good image. The geometric effect is preferably corrected by rescaling z to z cos(α) for each spot. c ) is carried out, whereby a c The angle α is that which the c-th channel has to the optical axis. The angle α is related to a virtual position of the area detector 19, which, taking the magnification into account, places it directly in front of the eye. For an area detector that lies exactly in a plane conjugate to the pupil plane of the eye, the area detector thus reaches the plane of the pupil P of the eye 3 with an extent that is modified by the magnification.

[0092] In aberration reconstruction, different channels are reconstructed independently. Subsequently, cross-correlation is performed in the axial direction, i.e., in the depth direction, to determine the relative phase shift between the individual channels. Reconstructing the lateral image for each channel (possibly taking the scanning process into account, as described below) and then the phase gradient yields a lateral shift in the image obtained for a given scanner position. This image is subsequently referred to as the pupil canal partial image. In one embodiment, the aberration is determined by means of a lateral cross-correlation of the pupil canal partial image, and the entire aberration phase distribution is thus determined and numerically corrected.

[0093] The quality of these approaches depends on the sample structure. In the human eye, a clearly recognizable axial layered structure is present. Laterally, the structures are relatively rough, for example, due to blood vessels or the optic disc combined with very fine structures such as photoreceptors, with hardly any structures of intermediate size and roughness. Therefore, in a preferred embodiment, a depth correlation correction is first performed by using the axial layered structure to correct the majority of pupil phase aberrations. Optionally, a lateral correlation correction follows, utilizing lateral structures, such as photoreceptors, that became visible as a result of the first correction.

[0094] The aberrations of the eye differ at various locations on the retina. In principle, it is possible to calculate the aberration-induced phase changes in each channel for all locations in a lateral image. In a simplified approach, it is assumed that the aberrations do not vary greatly laterally, and the aberrations are calculated only for a few lateral locations on the retina, with interpolation used for intermediate locations.

[0095] When a comparatively large wavelength range is traversed, it is preferable to consider the dispersion of aberrations. In this embodiment, it is not assumed that the phase shifts change linearly with the wavenumber k. Therefore, a peak in the profiles, which in the OCT image originates from the retina 2 at the fundus of the eye 3, is used to compensate for the shift of the profiles relative to each other. Thus, for example, one searches in curves 51 to 54 of the Fig. 7. A structure (in the form of a peak) is identified, and the curves are corrected relative to each other using this reference structure. In this way, the aberrations Θ can be corrected. c (k0) is determined and corrected as described above. Alternatively, a complex correlation algorithm can be applied to the profiles of the different channels. In addition to a shift, scaling (compression or stretching) of the measurement signals can also be corrected.

[0096] In one position of the scanner 13, a partial image of the retina is obtained, the size of which is determined by the aperture 15 (dimensions, aperture size, and number of apertures) and the front optics 12 and detector optics 14, which contribute to imaging the measuring light. A Fourier transform of the signal from the channels yields the image of the sample, but for each spot, only the portion corresponding to the size of the detected spots in the pupil. To generate a larger image, the scanner 13 is provided, which shifts the position of the imaged object field, i.e., the object spots, on the retina 2. The image area of ​​each spot corresponds to a partial image 59, which has a center 60. For the sake of simplicity, it suffices to refer to the center 60 of the partial image 59 for the actual deflection by the scanner 13. Scanning multispot images is generally known in the prior art, for example, from confocal microscopy. In this respect, reference is made to the corresponding techniques.However, it should be added here that not only lateral information is recorded by adjusting the scanner, but also depth information is recorded by tuning the wavelength of the radiation source.

[0097] This now enables various scanning approaches. For example, the scanner can be left stationary while the wavelength of the light source 4 is being tuned. Before retuning, the scanner is moved so that the next layer of the spot pattern is shifted to match the previous layer. In this way, the layers of the spot patterns can collectively determine a larger overall image 61 of the retina. This approach is in Fig. Figure 9 shows a depth plane. As a result, individual partial images 59 (corresponding to a position of the spot pattern) are combined to form the overall image 61. The images from the individual planes then produce a three-dimensional image of a cuboid area in the retina 2. This shows Fig. 10, in which three levels 62, 63 and 64 are shown as examples. The sub-images 59, which are in the representation of the Fig. The images 10, which are related to each other by a dashed double arrow, each originate from a wavelength sweep at the light source 4. Since the scanner 13 remains stationary during each wavelength sweep and is only adjusted in between, the partial images 59 generated from one wavelength sweep in planes 62 to 64 all lie exactly on top of each other with their centers 60.

[0098] For certain embodiments of the scanner 13, it is preferred to operate it continuously, i.e., to adjust it while the wavelength is being adjusted in a single pass. This approach requires synchronization of the scanner 13 and the wavelength adjustment at the light source 4. It is preferred to adjust the lateral adjustment speed of the scanner 13 such that, during a single wavelength pass, a maximum of one partial image 59 is scanned in one direction, preferably even less. The partial image 59 then differs from the partial image of the Fig. 9. In this approach, the adjustment of the scanner 13 can, for example, correspond to the spacing of the object spots, so that a partial image 59 corresponds exactly to the shift of the spot pattern by one period, i.e., by the distance between two adjacent individual object spots. This changes the position of the centers 60 for the individual planes 62, 63, and 64, since the partial images 59 in the planes are transformed into the different wavelengths by the Fourier transform. The result is a preliminary overall image 61, which, unlike in the embodiment of the Fig. 10 is not a rectangular prism, but, for example, due to the adjustment of the scanner 13 during wavelength tuning, a parallelepiped that is no longer rectangular. For imaging, this effect is preferably corrected by cropping the parallelepiped to a rectangular prism.

[0099] There are various ways to account for the simultaneity of wavelength tuning and lateral shift. If the detector is located near an intermediate image plane, i.e., in a plane conjugate to the retina, the data of the three-dimensional parallelepiped are shifted relative to each other. For each wavenumber k i Can an image of the sample be joined together, where I holds true? i =I(k i ,x,y). These images l i The values ​​are slightly shifted relative to each other. Since the relationship between lateral scan position and wavenumber is known, the entire wavelength shift can be combined accordingly for each location (x, y) in the sample. In this way, the three-dimensional data are easily assembled.

[0100] In embodiments where the detector is located in or near the conjugate pupil plane, it measures the Fourier transform of the intensity distribution in the object plane (retina 2). A displacement in the object plane leads to a phase ramp in the detector plane; therefore, the correction for the simultaneous lateral adjustment by the scanner 13 and the wavelength adjustment by the light source 4 is a multiplication of the detector signal by a time-dependent phase ramp that is proportional to the scan speed and the distance between the pupillary subchannel and the optical axis in the pupil plane.

[0101] The optical structure of the Fig. In steps 1 to 3, the illumination and the reception of the measuring light are no longer coupled with respect to their optical properties, and in particular the pupil size. This allows for adjustment of the illumination. For example, Bessel-type illumination can be combined with a top-hat cross-sectional profile for detection. In one embodiment, this achieves a high depth of illumination, i.e., illumination that remains unchanged over a large z-range, while simultaneously maintaining a high numerical aperture of the image. With the same numerical aperture, for example, a Gaussian beam would achieve an illumination focus with a diameter of 1 mm in the z-direction. With Bessel-type illumination, a diameter of 2 to 3 mm in the z-direction is obtained. In this way, the optical resolution can be increased by 10 to 30% when detection is performed with a top-hat profile.

[0102] In another embodiment of the OCT, polarization splitting takes place at the beam splitter 11. Such splitting is usually disadvantageous in the prior art, and intensity splitting is typically used. Surprisingly, this is advantageous for the described OCT because the polarization state of the polarized radiation entering the eye is altered. Different structures of the eye have varying effects, so that the polarization state of the backscattered signal is not unique or clearly defined, but rather consists of components with different polarization states. This consideration was also known in the prior art and led to the conclusion that intensity splitting is necessary precisely because the backscattered radiation does not have a clearly defined polarization state.However, it now turns out that the measurement light is superimposed on the reference light, and only components of the beam with the same polarization state can interfere with each other. Ultimately, the polarization state of the reference light determines which portion of the measurement light can be used. Non-interfering components fall on the detector and form a disturbing background.

[0103] The polarization division is described below with reference to the embodiment of the Fig. 2 explained, but is not limited to the features otherwise realized there and can, for example, also be applied to the embodiment according to Fig. 1. The illumination radiation B is linearly polarized after the polarization splitter 11. The lambda / 4 plate 27, as shown in Fig.The device shown in figure 2 provides circularly polarized illumination radiation B at the eye. 3. Backscattered measurement radiation M, which is also circularly polarized, is re-linearly polarized by the lambda / 4 plate, with the polarization direction rotated by 90 degrees relative to the polarization direction of the illumination radiation B emitted by the polarization splitter 11. Thus, the measurement radiation M passes through the polarization splitter 11 without deflection and interferes with the reference radiation R if it has the same polarization. This is the case if the reference radiation R and the illumination radiation B are identically linearly polarized after being split from the source radiation. This is also the case if the reference radiation R and the illumination radiation B are circularly polarized after being split from the source radiation and the reference radiation is linearly polarized identically to the measurement radiation M before being superimposed.Ultimately, it is important that the polarization splitting (e.g. by polarization splitter 11 and plate 27) conditions the measurement radiation M and the reference beam path conditions the reference radiation R so that both radiations have the same polarization state at the detector.

[0104] This measure thus increases the signal-to-noise ratio, since only those parts of the measuring light capable of interfering with the reference light are directed through the beam splitter 11 to the detector device 17. Ultimately, the polarization splitting and the rejection of a portion of the measuring radiation M at the beam splitter 11, which is inherently disadvantageous, improves the signal quality.

[0105] In another embodiment of the OCT, the fact that the illumination optics 10 allow the focus of the illumination beam B to be placed at a different z-position than the focus determined by the detector optics 14 for collecting the measurement beam M is utilized. Due to multiple scattering in the retina, measurement beam M from the retina may have a path length suitable for interference but propagate in a different direction, which would limit the lateral resolution in depth. By using different depth levels for illumination and detection, this effect can be compensated for, and the resolution in depth is optimized.

[0106] To reconstruct the image from the detector signals, the current wavelength must be known according to the FD-OCT principle. This wavelength, or the corresponding wavenumber k, can be derived from the control signal of light source 4. Alternatively, it is possible to extract a portion of the beam and measure its wavelength to better understand the currently set wavelength or the progression of a wavelength sweep.

[0107] Detector channels can be grouped perpendicular to the scanning direction to reduce speckle. This is particularly advantageous when only z-slices through the retina are desired.

[0108] For a high-resolution image, e.g., as a preview, it is possible to combine all or several detector channels for each spot. This is done after the corrections (e.g., aberration, z-position, overall image generation). The result is a resolution comparable to known OCT systems, but with a higher signal-to-noise ratio and improved speckle behavior, precisely because the combining occurs after one or more of the corrections and thus goes beyond normal pixel binning.

[0109] If a detector with only one spatial resolution is used, aberrations can only be corrected in that direction. This may be sufficient for certain applications.

[0110] In one embodiment, an iris camera is provided to assist the user in adjusting the device to the eye position.

[0111] The following further developments can be advantageously used for all embodiments of the described optical coherence tomograph or optical coherence tomography method: Phase errors that arise when the area detector 19, 19a, 19b is not exactly in the focal plane of the microlenses of the multilens array 36, 36a, 36b can be corrected numerically.

[0112] The microlenses of the multilens arrays, and thus ultimately the illumination spots on the retina, can be arranged in a square or hexagonal grid. Since round apertures are preferred for the multi-hole apertures, and the pupil or detection aperture is generally approximately round, a hexagonal grid allows for a further reduction in the number of detection pixels, i.e., area detectors with fewer pixels.

[0113] It is preferred to have exactly one pixel of the area detector 19, 19a, 19b at the center of each imaged spot, regardless of the grid of the illumination spots on the retina 2. Therefore, with a hexagonal grid of illumination spots in combination with a rectangular grid of pixels of the area detector 19, 19a, 19b, the size of the holes of the multi-hole aperture 34, and thus also of the multi-hole apertures 15, 15a, 15b, should be adapted to the pixel size and ultimately the resolution of the area detector 19, 19a, 19b such that this condition is sufficiently, i.e., at least approximately, e.g., to + / - 10% of the spot diameter, fulfilled.

[0114] Insofar as the above-mentioned process steps and / or signal corrections have been described, these are carried out in the OCT 1 by the control unit C, which is connected to the detector, reads out its measurement signals and receives further data about the operation of the scanner 13 and the wavelength tuning and / or controls these components accordingly.

Claims

[1] Optical coherence tomograph for examining an eye (3) which features: - a lighting device (4, 5) for providing source radiation, - an illumination and measurement beam path (7) comprising a splitting element (6) for splitting the source radiation into illumination radiation (B) and reference radiation (R), with the illumination radiation (B) illuminating an illumination field in the eye (3) and collecting backscattered illumination radiation in the eye (3) as measurement radiation (M), wherein the illumination and measurement beam path (7) comprises a scanner (13) for adjusting the lateral position of the illumination field in the eye (3) and a front optic (12), - a reference beam path (8) which provides for the reference radiation (R) an optical path length (21) corresponding to an optical path length from the splitting element (6) to the illumination field and back to a superposition point (71), - a detection beam path (14, 15, 17) which receives the measuring radiation (M) from the illumination and measuring beam path (7) and the reference radiation (R) from the reference beam path (8) and superimposes them at the superposition point (71) and directs them onto a detector (19), - the illumination and measuring beam path (7) further exhibits - a beam splitter (11) for separating the measuring radiation (M) collected by the eye (3) from the illumination radiation (B) directed to the eye (3), wherein the beam splitter (11) directs the separated measuring radiation (M) to the detection beam path (14, 15, 17), and - a light splitting element (32, 34) that splits the illumination radiation (B) into spots to illuminate the retina (2) with a multi-spot pattern, characterized by , that - the lighting device (4, 5) is tunable with respect to the wavelength of the source radiation, - the detector is an area detector (19), - the detection beam path further - an optical element (14) acting only on the measuring radiation (M), which interacts with the front optics (12) and adjusts the numerical aperture with which measuring radiation (M) is collected in the eye (3), and - an aperture (15) which is arranged in front of the area detector (19), in or near an intermediate image plane and which defines the size of an object field from which the measuring radiation (M) reaches the area detector (19), and - wherein the aperture prior to the area detector is designed as a first multi-hole aperture (15) and a first multi-lens array (36) is arranged between this multi-hole aperture and the area detector (19), which focuses the radiation emanating from each hole of the first multi-hole aperture (15, 15a, 15b) onto a pixel area of ​​the area detector (19) which has a spatial resolution of 4 to 100 pixels in one direction, preferably as a 2D pixel area with 5 to 50 pixels or 5 to 40 pixels per direction. [2] Optical coherence tomograph according to claim 1, characterized by , that the light splitting element is designed as a second multi-lens array (32) and a second multi-hole aperture (34) located in an intermediate image plane. [3] Optical coherence tomograph according to claim 2, characterized by, that the area detector (19) lies in a plane which is conjugate to a plane in which the pupil (P) or iris of the eye (3) lies, and that the second and the first multi-hole aperture (32, 15) each lie in a plane which is conjugate to a plane in which the retina (2) of the eye lies. [4] Optical coherence tomograph according to claim 1, 2 or 3, characterized by , that the beam splitter (11) is a polarization beam splitter and that a lambda / 4 plate (27) is arranged between the eye (3) and the beam splitter (11), which filters the measurement radiation with respect to a polarization state that is adapted to a polarization state of the reference radiation (R). [5] Optical coherence tomograph according to any of the above claims, characterized by, that the illumination and measuring beam path (7) has an optical element (10) acting only on the illumination radiation (B), which interacts with the front optics (12) and sets the numerical aperture of the illumination of the illumination field in the eye (3). [6] Optical coherence tomograph according to any of the above claims, characterized by , that the reference beam path (8) has a reference radiation multi-lens array (36c) that focuses the reference radiation (R) into a multi-spot pattern corresponding to the second multi-hole aperture (15, 15a, 15b). [7] Optical coherence tomograph according to any of the above claims, characterized by , that a jet splitter (71) is arranged at the point of overlap. [8] Optical coherence tomograph according to any of the above claims, characterized bya control unit (C) that controls the scanner (13) for deflection during wavelength tuning and generates or receives a scan signal indicating a deflection state of the scanner (13), and is connected to the radiation source (4) for reading a wavelength signal indicating the wavelength of the source radiation and thus of the illumination radiation (B), and to the area detector (19) for reading measurement signals for each pixel, wherein the control unit generates partial images (59) of the retina (2) from the wavelength signal and the measurement signals and evaluates the scan signal to combine the partial images (59) into a 3D overall image (61). [9] Method for optical coherence tomography for examination of an eye (3) wherein the method has - Providing source radiation, matching its wavelength and splitting the source radiation into illumination radiation (B) and reference radiation (R), - Illuminating an illumination field in the eye (3) with the illumination radiation (B) and collecting backscattered illumination radiation in the eye (3) as measurement radiation (M), wherein a scanner (13) is used to adjust the lateral position of the illumination field in the eye (3) and a front optic (12) in the eye (3) are used, - Separation of the measuring radiation (M) collected by the eye (3) from the illumination radiation (B) directed to the eye (3), - Determining the numerical aperture of the illumination of the illumination field in the eye (3) by using an optical element (10, 35) acting only on the illumination radiation (B) and cooperating with the front optics (12), and determining the numerical aperture with which measuring radiation (M) is collected in the eye (3) by using an optical element (14) acting only on the measuring radiation (M) and cooperating with the front optics (12), - Superimposing the measurement radiation (M) with the reference radiation (R) and detecting an interference signal of the superimposed radiations with a surface detector (19), and - Using an aperture positioned upstream of the area detector (19) and located in or near an intermediate image plane to define the size of an object field from which the measurement radiation (M) reaches the area detector (19), - wherein the retina (2) is illuminated with a multispot pattern and the aperture arranged upstream of the area detector (19, 19a, 19b) is designed as a multi-hole aperture (15, 15a, 15b) and the radiation emanating from each hole of this multi-hole aperture (15, 15a, 15b) is focused onto a pixel area of ​​the area detector (19) which has a spatial resolution of 4 to 100 pixels in one direction, preferably as a 2D pixel area with 5 to 50 pixels or 5 to 40 pixels per direction. [10] Method according to claim 9, characterized by, that the area detector (19) is arranged in a plane which is conjugate to a plane in which the pupil (P) or iris of the eye (3) lies, that the multispot pattern is generated by means of an illumination multi-hole aperture (34), and that the multi-hole aperture (15, 15a, 15b) arranged upstream of the area detector (19, 19a, 19b) lies in a plane which is conjugate to a plane in which the illumination multi-hole aperture (34) lies. [11] Method according to claim 9 or 10, characterized by , that the separation of the measurement radiation (M) collected by the eye (3) from the illumination radiation (B) directed to the eye (3) is carried out by means of a polarization division, wherein the measurement radiation (M) is filtered with respect to a polarization state which is adapted to a polarization state of the reference radiation (R) during superposition, and wherein portions of the measurement radiation (M) not corresponding to this polarization state are discarded. [12] Method according to any one of claims 9 to 11, characterized by , that the illumination radiation (B) in the pupil (P) of the eye (3) is uniformly distributed within a cross-section covered by the illumination radiation (B) except for intensity fluctuations of + / -10%. [13] Method according to any one of claims 9 to 12, characterized by , that the reference radiation (R) is focused by a reference radiation multi-lens array (36c) into a multi-spot pattern corresponding to that of the multi-hole aperture (15, 15a, 15b). [14] Method according to any one of claims 9 to 13, characterized by , that the reference radiation (R) and the measurement radiation (M) are superimposed with a beam merging device (71). [15] Method according to any one of claims 9 to 14, characterized by, that the scanner (13) is controlled for deflection during wavelength tuning and partial images (59) of the retina (2) are generated from measurement signals of the area detector (19) and a wavelength signal and, taking into account the deflection state of the scanner (13), the partial images (59) are combined to form a 3D overall image (61).

Images