Holoscopic optical coherence tomography

Abstract

Method for holoscopic optical coherence tomography of an object (3), in particular an eye, wherein the method comprises - Providing source radiation and splitting the source radiation into illumination radiation (B) and reference radiation (R), - Illuminating a lighting field extending horizontally across the object (3) with the lighting radiation (B), - Collecting backscattered illumination radiation from the object (3) as measuring radiation (M), - Separating the measuring radiation (M) collected by the object (3) from the illumination radiation (B) directed to the object (3), and - Superimposing the measurement radiation (M) with the reference radiation (R) and detecting an interference signal of the superimposed radiations with at least one detector (17), characterized in that - the object (3) in the illumination field is illuminated simultaneously with more than one spatial radiation mode, - where the radiation modes of the illumination in the illumination field are spatially and temporally coherent to each other, but have a fixed phase difference to each other.

Description

[0001] The invention relates to a method for holoscopic optical coherence tomography of an object, in particular an eye, wherein the method comprises providing source radiation and splitting the source radiation into illumination radiation and reference radiation, illuminating an illumination field extending horizontally across the direction of incidence on the object with the illumination radiation, collecting illumination radiation backscattered by the object as measurement radiation, separating the measurement radiation collected from the object from the illumination radiation directed to the object, and superimposing the measurement radiation with the reference radiation and detecting an interference signal of the superimposed radiations with a two-dimensionally extended area detector.

[0002] The invention further relates to an optical coherence tomograph for the holoscopic examination of an object, in particular an eye, comprising: an illumination device for providing source radiation, an illumination and measurement beam path comprising a splitting element for dividing the source radiation into illumination radiation and reference radiation, illuminating an illumination field in the object with the illumination radiation extending horizontally across the direction of incidence and collecting backscattered illumination radiation from the object as measurement radiation, a reference beam path comprising 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 to a superposition point, 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, superimposes them at the point of superposition, and directs them onto a two-dimensionally extended area detector.

[0003] Such methods and coherence tomographs are known from DE 102014115157 A1. Similar approaches can also be found in DE 102014115153, DE 102014115155 A1, DE 102015101251 A1 and WO 2017 / 137567 A1.

[0004] 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 here. In OCT systems, the lateral resolution (x and y) is determined by the numerical aperture of the optics used. The axial resolution, 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.

[0005] Measuring objects such as 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 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.

[0006] 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.

[0007] In a scanning OCT system, the accessible diameter of the 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.

[0008] 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 needed to detect such vessels. The patent patent DE 102014115157 A1 therefore provides a solution that improves lateral resolution in holoscopic optical coherence tomography (OCT) when used on the retina of the human eye, without limiting the accessible depth range.

[0009] OCT systems that utilize a spatially resolved detector array and coherent evaluation algorithms known from digital holography are referred to in the literature as holoscopic optical coherence tomographs.

[0010] Here, the term "holoscopic system" generally refers to an interferometric OCT system with more than one lateral sensor element. In this notation, for example, holoscopic wide-field systems, line-scanning OCT systems, and small-field-scanning OCT systems are holoscopic systems, while confocal systems are not, even if they use balanced detection with two sensors.

[0011] In such holoscopic multi-channel systems, the optics are designed so that the light is diffraction-limited onto the detector elements, so that each detector element is able to detect only a very limited light conductance.

[0012] In confocal systems, illumination is provided by a spatially coherent single-mode source. Fiber-coupled lighting systems are most commonly used. The light exiting this single-mode lighting fiber exhibits a TEM00 mode spectrum with a light conductance of approximately 2πλ. 2 The optical conductance of the detectable light backscattered from the sample is identical to the illumination optical conductance and the mode shape. For this reason, the part of a light wave field that carries the optical conductance of a TEM00 mode is generally referred to as a spatial mode in the following.

[0013] Therefore, confocal systems with only one lateral sensor element are single-mode in detection, while systems with more than one spatial sensor element are multi-mode. The number of detector elements (pixels) is thus a measure of the number of detection modes.

[0014] OCT systems can be illuminated spatially coherently or incoherently. Preferably, "time domain" systems are illuminated at least partially with incoherent light wave fields such as LEDs or thermal lamps. However, in such systems, the measured values ​​cannot be coherently reconstructed. For example, white light interferometers are frequently implemented in this way.

[0015] In contrast, conventional OCT systems (e.g., confocal systems, full field holoscopic systems, line scanning systems, and small field scanning systems) are illuminated with a spatially fully coherent, spatial single-mode source to enable coherent reconstruction.

[0016] In all systems where illumination is performed in a single spatial mode while detection is multi-mode, the field distributions of the illumination and detection modes can be aligned within a reference plane (e.g., the focal plane) through appropriate optical design. Only the parts of the sample that can be detected are then illuminated. However, outside the focal plane, the illumination mode and the detection modes diverge due to their differing optical conductivity. A certain degree of mode mismatch outside the focus is therefore unavoidable in such systems.

[0017] The basic idea of ​​"multi-directional scanning OCT" (WO 2017 / 137567 A1) is to illuminate a sample point with a spatial single-mode and detect it with a corresponding sensor element in a suitable detection mode. This combined illumination / detection mode can then be switched sequentially between different illumination directions. If the data measured from the sample from different directions are thus coherently reconstructed, a higher light conductivity can be achieved compared to a confocal system, and therefore a higher lateral resolution with a sufficient focal depth (of the single mode) can be obtained.

[0018] Currently, commercially available OCT systems used in the eye are typically confocal systems. With these systems, there is a fixed relationship between pupil size, lateral resolution, and the achievable depth range (Rayleigh length). Doubling the pupil size doubles the lateral resolution but reduces the depth range by a factor of four. For this reason, all systems used for retinal imaging have similar pupil sizes and thus similar lateral resolutions in the range of approximately 17 µm.

[0019] One possible way to overcome this limitation is to use multiple sensors in any technical configuration (e.g., wide-field, line-scanning, or small-field scanning). However, as long as these systems are illuminated with only a single lighting mode, which is the case for all systems known from the current state of the art, the "mode mismatch" described above cannot be circumvented, and this mismatch, due to the associated vignetting effect, severely limits the effectively achievable depth range.

[0020] Time-sequential systems, such as multi-directional scanning OCT, require longer acquisition times. All motion artifacts during measurements on, for example, living eyes, alter the phase information of individual measurements, thus complicating coherent reconstruction.

[0021] The invention is based on the objective of expanding the accessible depth range in holoscopic optical coherence tomography.

[0022] The invention is defined in claims 1 and 7. Advantageous embodiments are the subject of the dependent claims.

[0023] The underlying insight of the invention, which is attributable to the inventor for the first time, is that spatial coherence as a necessary property for a coherent reconstruction in holoscopic coherence tomography and a multi-mode illumination wave field are not mutually exclusive, but rather complement each other advantageously.

[0024] Therefore, a method for holoscopic optical coherence tomography of an object, particularly an eye, is provided, in which source radiation is supplied and split into illumination radiation and reference radiation. The illumination radiation is used to illuminate a field on the object that extends horizontally across the direction of incidence. Illumination radiation reflected back from the object is collected as measurement radiation. It is separated from the illumination radiation directed to the object in the beam path and superimposed with the reference radiation. A detector detects the superimposed radiations and thus an interference signal. This signal is then evaluated using known coherent evaluation algorithms. The detector can perform balanced detection. A non-spatially resolved confocal detector in combination with a scanner for scanning the object is suitable as the detector.Furthermore, the detector can be one- or two-dimensionally spatially resolving, e.g. a two-dimensionally extended area detector, whereby the individual detector elements can also be spaced apart.

[0025] Based on this finding, the object in the illumination field is simultaneously illuminated by more than one spatial radiation mode. The different radiation modes of the illumination are spatially and temporally coherent to each other within the illumination field. However, they have a fixed phase difference, meaning they differ by different static initial phases.

[0026] Regarding the optical coherence tomography system for the optical and holoscopic examination of an object, an illumination device is provided which supplies source radiation. An illumination and measurement beam path includes its splitter element for separating the source radiation into illumination radiation and reference radiation. It further illuminates a planar illumination field on the object, extending transversely to the direction of incidence, and collects backscattered illumination radiation from the object as measurement radiation. A detection beam path receives the measurement radiation from the illumination and measurement beam path as well as the reference radiation from a reference beam path and superimposes them at a point of superposition. It then directs them to a detector, e.g., confocal or as a two-dimensionally extended area detector.

[0027] Furthermore, the illumination and measurement beam path illuminates the object in the illumination field simultaneously with more than one spatial radiation mode. The radiation modes of the illumination are temporally and spatially coherent to each other in the illumination field, but have a fixed phase difference. This fixed phase difference is generated because the illumination modes have unequal static initial phases when they are generated.

[0028] In a preferred embodiment of the method and coherence tomograph, the initial phases or phase differences are distributed statistically or quasi-randomly across the illumination field.

[0029] Technically, this can be achieved, for example, by illuminating a static phase plate, such as a lithographic scattering plate, with spatially coherent light from a spatial single-mode source (such as an SLD or swept-source laser). The light transmitted through this element has a significantly increased optical conductivity, but its temporal and spatial coherence remains unaffected because it is static and introduces only small time-of-flight differences. The element is designed and arranged such that the optical conductivity is increased by a specific factor, known as the number of illumination modes.

[0030] A technically equivalent description is to describe a single-mode source with an M 2 to use the beam parameter product of approximately 1 and to transform the beam with the phase element so that an M 2 is generated by a value significantly above 1.

[0031] The optical conductivity is measured in each contiguously illuminated sample zone to distinguish the new method from multispot systems, in which several confocal sensors simultaneously measure separate areas of the sample and in which one M is measured per spot. 2 The total light conductance can be measured at around 1, but also at a significantly higher overall value. The key feature of the new arrangement is that the different illumination modes can overlap within the sample volume and interact with multiple detector channels simultaneously. This is also evident from the fact that the intensity of the illumination wave is modulated in the field plane and thus deviates from typical Gaussian beam profiles.

[0032] In a further development of the method and coherence tomography, it is therefore provided that the area detector has adjacent detector elements in an image plane, between which gaps may also exist. Each detector element corresponds to a detector mode. The number of detector elements and the number of illumination modes are matched, preferably in a 1:1 ratio.

[0033] Even if the number of detector elements or detector modes is lower than the number of illumination modes, information reconstruction can still be performed. This involves evaluating spatial frequencies that are transferred from the illumination to the detection via mode interactions. The coherence tomograph has a control unit that performs the information reconstruction and preferably also carries out the coherent evaluation using the holoscopic coherence tomography principle.

[0034] The invention is based on the understanding that spatial coherence, a necessary property for coherent reconstruction in holoscopic coherence tomography, and a multi-mode illumination wavefield are not mutually exclusive, but can advantageously complement each other. Therefore, embodiments provide that in holoscopic optical coherence tomography, the illumination field contains more than one spatial radiation mode. Illumination is thus multimodal. The different radiation modes are spatially and temporally coherent to each other within the illumination field. They have a fixed phase difference with each other, i.e., they differ by different statistical initial phases. Speckle sometimes occurs in such an illumination field. In embodiments, the effect of this speckle is removed by incoherent averaging. For this purpose, several images are taken that are shifted relative to each other by at least one speckle grain.In this way, the images become mathematically independent of each other, and the speckle structure disappears due to averaging. Alternative designs also capture multiple images that are shifted relative to each other by at least one speckle grain. They compensate for the speckle effect by setting up a system of equations and suppressing multiple scattered signal components via linear regression of these equations. This increases the image contrast.

[0035] In a particularly simple embodiment with regard to speckle suppression, the illumination and measuring beam path generates a speckle pattern in the illumination field that has no axis of symmetry.

[0036] To generate the illumination modes, the illumination and measuring beam path can have a scattering element that is located in a field plane of the beam path.

[0037] It is understood that the features mentioned above and those to be explained below can be used not only in the combinations given, but also in other combinations or on their own, without leaving the scope of the present invention.

[0038] The invention is explained in more detail below with reference to exemplary embodiments and the accompanying drawings, which also disclose essential features of the invention. These exemplary embodiments serve only for illustration and are not to be interpreted as limiting. For example, a description of an exemplary embodiment with a plurality of elements or components is not to be interpreted as meaning that all of these elements or components are necessary for implementation. Rather, other exemplary embodiments may also contain alternative elements and components, fewer elements or components, or additional elements or components. Elements or components from different exemplary embodiments may be combined with one another unless otherwise specified. Modifications and variations described for one of the exemplary embodiments may also be applicable to other exemplary embodiments.To avoid repetition, identical or corresponding elements in different figures are designated with the same reference symbols and are not explained multiple times. The figures show: . Fig. 1A A schematic representation of a holoscopic optical coherence tomography (OCT) scanner with scanning, confocal imaging of an eye, Fig. 1B a schematic representation of a detector device for an OCT with opposite Fig. 1A modified illustration, Fig. 1C a schematic representation of a holoscopic optical OCT with imaging of the eye onto a flat detector, Fig. 2 and Fig. 3 the beam paths with respect to illumination in a conventional holoscopic system ( Fig. 2) and a system according to Fig. 1A, Fig. 1B, or 1C ( Fig. 3) Fig. 4 a schematic representation of an illumination vignetting effect that can occur outside the focal plane, Fig. 5 areas of the sample that are not vignetting for a given distance to the focal plane, Fig. 6 an accessible spatial frequency spectrum for 4-channel detection, Fig. 7 an accessible spatial frequency spectrum for 7-channel detection, Fig. 8A a close-up of a lithographically produced scattering plate, viewed under the microscope of the Fig. 1A, Fig. 1B or Fig. 1C can be used, and Fig. 8B a diffuser plate that produces one unspattered and three scattered lighting modes.

[0039] The following examples refer to holoscopic coherence tomography of the eye. However, they can also be used on other objects; the reference to the eye is purely illustrative.

[0040] The human eye is an optical system that simultaneously encompasses both the retina, and thus a certain field, and the pupil, and thus the far-field distribution. The mathematical pupil, whose field distribution corresponds to the Fourier transform of the retinal field distribution, lies only a few millimeters in front of the eye.

[0041] Fig. 1A and Fig. Figure 1C each shows an OCT 1, which takes three-dimensional images of, among other things, a retina 2 of an eye 3. Fig. Figure 1B shows a modified detector setup for the OCT 1 of the Fig. 1A. The described embodiments are always described using the example of a fiber-based swept-source system, but can also be applied to free-jet setups and / or time-domain systems and frequency-domain systems.

[0042] Source radiation from a wavelength-tunable radiation source 4, for example a corresponding 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. In the design of the Fig. 1A and Fig. In 1B, the OCT 1 performs a scanning confocal imaging procedure, in the design of the Fig. 1C is mapped onto an area detector.

[0043] In all fiber-based designs, the 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 modified with respect to the illumination modes by means of an illumination optic 10 and then directed to a beam splitter 11. From there, it reaches a front optic 12, which focuses the illumination radiation B into a focus located in the eye 3. Between the beam splitter 11 and the front optic 12, in the design of the Fig. 1A and Fig. 1B and optionally also in the embodiment of Fig. 1C a scanner 13 which deflects the focus on the retina 2 in a two-axis manner perpendicular to the direction of incidence, i.e., laterally. 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 optics 12.

[0044] The illumination radiation in the illumination focus on the retina 2 is scattered back from different depths z within the depth of field area.

[0045] The backscattered radiation is collected by the front optics 12 as measurement radiation M and (if present) 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.

[0046] The beam splitter 11 separates the measuring radiation M from the illumination radiation B and directs it to a detector device 17. In the embodiment of the Fig. 1A comprises the detector assembly 17, a beam splitter / combiner 18, and two sensors 19a and 19b. Confocal imaging is performed. The setup of the Fig. 1A implements a well-known "balanced detection" method. Fig. 1B provides a large number of balanced sensor pairs 19a / 19b to perform multimodal detection. Fig. In 1C, this is done in a wide field, and an area sensor 19c has a spatial resolution, i.e., it allows a resolution of the intensity distribution across the beam cross-section. The detector device 17 is preferably located in a pupil plane 30 of the measuring beam path, i.e., in a plane that is conjugate to the plane of the pupil P of the eye 3, through which the measuring radiation M emerges. However, since the light wave field, if its magnitude and phase are known in one plane, can be converted into any other plane by numerical propagation, detection can also be performed in other planes, e.g., conjugate to the retina as in Fig. 1C, will take place.

[0047] From a fiber 23 ( Fig. 1A and B) or by oblique illumination of the area detector ( Fig. 1C) Reference radiation R from the reference arm 8 is also coupled into the detector device 17. This has a fiber 20 after the splitter 6.

[0048] The reference arm 8 has in Fig. 1A and Fig. In the embodiments shown in Figure 1C, a path length adjustment device 21 is provided, which serves 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. The reference radiation R is then superimposed with the measurement radiation M and directed onto the sensors 19a and 19b or the area detector 19c.

[0049] The path length adjustment device 21 is designed as a free-beam path. This is optional, as is the use of a retroreflector 22. Various methods for adjusting the optical path length of a beam are known in the prior art.

[0050] The detector device 17 is in Fig. 1A and B are implemented as so-called "balanced detection". This is also optional. Balanced detection has the advantage that a common-mode component in the superposition of reference radiation R and measurement radiation M can be suppressed in a particularly simple way. Alternatively, such suppression could be omitted if only a single detector, for example sensor 19b, were used and the beam splitter / combiner 18 was implemented as a pure beam combiner.

[0051] The interference between the reference radiation R and the measurement radiation M is used to generate an image, as is known for scanning holoscopic 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.

[0052] In the construction method of the Fig. In 1A, only a single detection mode is evaluated; accordingly, the measurement radiation M is collected with a single lens 24 in a plane conjugated to the plane of the retina 2 and, by means of an optical fiber 25 at the beam splitter / combiner 18, is made to interfere with reference radiation from the optical fiber 23 and directed onto the sensor pair 19a / 19b. In the design of the Fig. In 1B, a multimodal confocal detection is performed by coupling the measurement radiation M from a plane conjugate to the pupil plane into a plurality of optical fibers 25.x using a lens 24 configured as a mini-lens array. Four optical fibers are shown as examples, one of which is designated with the reference symbol 25.x. If n optical fibers with a corresponding number of mini-lenses in the mini-lens array 24 are used, the radiation in optical fiber 23, i.e., the reference radiation, is also split into a total of n optical fibers by a beam splitter 26, of which optical fiber 27.x is designated as an example. The optical fibers 25.x and 23.x are coupled in pairs at beam splitters / combiners 18.x, so that the radiation reaches sensor pairs 19a.x / 19b.x. This procedure will be used for all optical fibers 25.x to measure the measurement radiation M from the mini-lens array 24 and reference radiation R from the optical fiber 27.x was carried out accordingly. In total, each mini-lens of the mini-lens array 24 represents a detection channel, which corresponds to a detector element. Each sensor pair 19.ax / 19.bx corresponds to a detector element in the detector assembly 17.

[0053] To perform image generation, the OCT 1 has a control unit C, which receives a signal via the wavelength matching and the measurement signals from the detector unit 17. Optionally, the control unit C controls the radiation source 4 for wavelength matching, thus knowing the currently dominant wavelength in the system and being able to assign the measurement signals accordingly. The sensors receive measurement radiation M from an object field at the eye 3. In the design according to Fig. In 1A and B, sensors 19a / 19b and 19a.x / 19b.x, respectively, together with the scanner, scan the intensity distribution with spatial resolution. In the design of the Fig. In 1C, this is handled by the area detector 19c. If it is optionally located 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 position information within the object field. If, on the other hand, it is located in the conjugate pupil plane 30, 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 phase information. This, too, can be used for image reconstruction, as is known. Of course, to detect a larger object field, the design of the Fig. 1C a scanner 13 can be used.

[0054] The beam splitter / combiner 18 or 18.x combines in Fig. 1A or B the measurement radiation M from the measurement arm 7 and the reference radiation R from the reference arm 8. The sensors 19a / 19b and 19a.x / 19b.x detect the interference pattern between the measurement radiation M and the reference radiation R through the action of the scanner 13. 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. The same applies to the design of the Fig. 1C. The holoscopic evaluation is described firstly in the patent specifications mentioned at the beginning and secondly also from the publication Hillmann et al., “Efficient holoscopy image reconstruction”, Optics Express, Vol. 20, No. 19, 21247.

[0055] Illumination radiation with a single illumination mode would have such a low optical conductance that it would focus either near the retina or near the anterior chamber. It can also be observed that in single-mode systems, the field distribution in the retinal and iris planes cannot be determined independently. For this reason, it would not be possible to holoscopically image the entire eye, even if the coherence length were sufficient.

[0056] All variants therefore share the feature of coherent multi-mode illumination, the illumination distribution and number of modes of which are preferably adapted to the detection arrangement. This allows the field size on retina 2 and the detectable pupil size in the anterior chamber to be set independently of each other (however, the possible optical resolution in retina 2 depends on the pupil size, and the anterior chamber resolution on the retinal field size). This enables true holoscopic whole-eye imaging.

[0057] The multimodal illumination in the illumination beam B is achieved by the illumination optics 10, which generates an illumination field plane in which a diffusing element 100 is located. This, for example, provides illumination in an illumination field located in the eye 3, which has a speckle field distribution. The diffusing element can optionally be swung out to switch the microscope to conventional illumination.

[0058] Fig. Figure 2 shows the incident illumination radiation B in the conventional single-mode case and the area covered by the measurement radiation M. As can be seen, the illumination radiation B can only capture a small part of the measurement field 104.

[0059] At the in Fig. In contrast, in the multimodal illumination implemented in 1A to C, the measurement field 104 is exactly the same size as the illumination field 102. This applies to the retina 2 as well as to the lens of the eye and in the area of ​​the pupil P.

[0060] From an image analysis perspective, there are several preferred implementations. These are based on the following considerations: When the retina is illuminated with a multi-mode speckle field distribution, this complex speckle field function is multiplied by the complex backscattering function of the retinal tissue, resulting in a "noisy" image of the retina. Conventional confocal imaging would also produce noisy images, as these involve multiplying a Gaussian beam distribution with the complex backscattering function of the retina. When a sample is illuminated with an extended coherent multimode wave, the field strength of this wave is modulated by speckle. The speckles have lateral dimensions determined by the numerical aperture of the illumination, corresponding to the diffraction-limited resolution. The speckles also have axial depths corresponding to the wave-optical depth of field or Rayleigh length.Each of these excitation speckle grains has a random phase and brightness, which is described by speckle statistics, a concept well-known to those skilled in the art. If a plane illumination wave is assumed during data reconstruction and the characteristics of the scattered illumination wave are not taken into account, the modulations of the illumination wave can become visible in the final image, especially at small illumination apertures, and distort the image impression. It is known from the literature that, for example, 3D retinal images acquired with holoscopic wide-field imaging exhibit significantly lower contrast than images acquired with a confocal point scanner due to the considerably more efficient detection of multiple scattering. If holoscopic wide-field systems are combined with a partially confocal method, multiple scattering can be significantly reduced; however, this also negates the advantages of holoscopic imaging, such as…The decoupling of lateral resolution and accessible depth range is limited. Examples of such semi-confocal systems known from the prior art are line scanners or "spinning disc" systems (with so-called Nipkow disks). With line scanners, the image is acquired confocally in one direction, so the accessible depth range and lateral resolution can only be decoupled within the detection lines. With Nipkow disks, the sample is scanned simultaneously with many confocal spots, so the properties are more like those of a confocal system than a holoscopic detection system. In the prior art, especially in microscopy, there are various methods known under the term structured illumination.In contrast to partial confocal imaging, where unwanted extrafocal signals are suppressed, the goal of structured illumination is to specifically measure and computationally subtract the extrafocal signal components. For this purpose, the sample is illuminated with an intensity-modulated or structured light wave. Multiple images are acquired from each sample point using different illumination patterns or pattern positions, and these images are then combined during analysis.

[0061] In image analysis, the multiply scattered sample light is suppressed without compromising the holoscopic advantages of decoupling lateral resolution and accessible depth. For this to be possible with the holoscopic imaging technique used here, the illumination patterns must be structured at all depths of the sample.

[0062] In a first approach, the complex retinal backscatter function is additionally multiplied by the complex illumination field distribution using the multi-mode illumination employed here, thus altering the speckle statistics. However, this effect has no impact on parameters such as lateral resolution or sensitivity. Therefore, this is a first step towards disregarding the multi-mode illumination in the evaluation.

[0063] In a second variant, several complete holoscopic datasets are acquired in which, for example, random eye movements or scanning of the illumination distribution cause the illumination speckle distribution relative to retina 2 to be shifted in a random direction by more than one speckle grain. The image data are then analyzed to further improve image quality by estimating and correcting the 3D displacement vectors between the holoscopic datasets and then incoherently averaging the acquisitions. This preserves the lateral resolution of the images while increasing their sensitivity and removing speckle.If the exact displacement vectors of the holoscopic sub-datasets are known, or estimated through cross-correlations, and if the exact speckle function is also known, then the 3D reconstruction algorithms known from holoscopy (with estimation and correction of ocular aberrations) can be applied, thus doubling the lateral resolution. This allows for a lateral resolution of up to 2.5 µm with an ocular pupil diameter of approximately 4.5 mm.

[0064] In a third variant, the sample is illuminated with a coherent scattered multimode light wave field, and the resulting illumination speckles in all planes of the sample are used for structured illumination. Several 3D datasets of a sample are acquired with scattered coherent illumination. The datasets contain no motion artifacts or optical aberrations, or these have been previously estimated and numerically corrected, e.g., using algorithms otherwise known for holoscopic imaging. Reference is made to the publication Hillmann et al., "Efficient holoscopy image reconstruction," Optics Express, Vol. 20, No. 19, 21247. For example, by exploiting random sample movements (eye movement) or active (e.g.,(Motorized) adjustment changes the position of the illumination speckles relative to a defined sample point such that the illumination speckles in the different images are shifted from each other by at least one speckle grain. If n 3D datasets of the sample are acquired in this way, then for each 3D data point, n measurements Yi are obtained with different but known illumination speckle brightnesses Xi. This allows for n equations of type... Yi=Xi*eS+mS The setup involves acquiring n 3D datasets. Here, i corresponds to the dataset number and ranges from 1 to n, eS represents the singly scattered signal, and mS represents the multiply scattered signal; these are complex quantities. The equation shown can be formulated independently for each sample point within the 3D cube. A linear regression over these equations with complex quantities then yields an image with the desired magnitude of the singly scattered signal eS. Due to the random phase relationships, all speckles average out from the magnitudes of the complex quantities eS and mS. By suppressing the multiply scattered signal mS, the image contrast is significantly better than with holoscopic images and reaches the contrast range of confocal images. For various analyses of the OCT data under coherent multimode illumination, it is necessary to be able to predict the precise functionality of the speckles at different sample depths.A linear scattering model is assumed for this purpose. This means that the wave propagates through an isotropic transparent medium across the different sample depths. This model assumption is also used in the prior art for holoscopic volume reconstruction. Depending on the scattering strength of the sample, this model assumption is only applicable up to certain sample depths, depending on the spatial frequencies used in illumination and detection and the resulting lateral resolution. For very large measurement depths, however, the assumption no longer holds completely, so that the resolution in coherently reconstructed images is then no better than in incoherently reconstructed images.To achieve OCT images with very high penetration depths, as required for retinal imaging of the eye (e.g., for diagnosing choroidal abnormalities), it is particularly advantageous to limit the numerical aperture of the illumination waveform. This results in coarser illumination speckle structures. While this reduces the maximum achievable lateral resolution, it simultaneously improves the suppression of the multiply scattered background at greater measurement depths. Technically, this can be achieved, for example, by selectively shading the illumination waveform with an adjustable (e.g., motorized) pupillary diaphragm. A particularly preferred alternative is to incorporate (e.g., pivot) a diffuser in the intermediate image plane (e.g., a diffuser with smaller diffraction angles). The diffraction angles correspond to illumination apertures."Smaller" means that the illumination apertures are smaller than the detection apertures.

[0065] The basic principles of the three variants can also be applied to line-scanning systems, whereby the known scan movement can preferably be used to selectively shift the illumination speckle distribution across the retina.

[0066] Typical OCT systems are preferably implemented with "shot noise" limitation. This means that the acquisition speed for a given sensitivity is limited only by the illumination intensity and the detection pupil size. The tissue of the iris and retina are similarly sensitive to thermal stress from the illumination light sources used. Therefore, the light limits for both parts of the eye (retina and anterior chamber) are similar according to the ophthalmic device standard DIN EN ISO 15004-2 and the laser standard DIN EN 60825-1, and are in the range of 0.5–5 mW for single-mode illumination. Even if the system is designed so that the illumination focus is in front of the eye, it is technically difficult to ensure that, even with eye movements and misfocusing, the illumination focus never falls on any tissue of the eye and endangers it.

[0067] Multi-mode illumination allows for significantly higher overall illumination intensities to be safely applied, as the higher light transmittance prevents higher peak intensities from reaching the eye. For line scanning, illumination waves of up to approximately 130 mW and for wide-field systems up to 1 W of continuous power can thus be used in compliance with standards.

[0068] The light limits depend only on the illumination conductance and not on the coherence of the light wave field. For this reason, incoherent multi-mode illumination can be implemented with similarly high permissible illumination intensities. However, with these systems, coherent reconstruction outside the focal plane is not possible due to the incoherence of the illumination. Therefore, "time domain" systems are implemented in such a way that the system is always actively focused on the plane to which the delay is set. Thus, these systems do not require coherent reconstruction. Partially coherent systems are also possible, where, for example, detection is performed with 1000 x 1000 pixels, but illumination is only achieved with 100 x 100 modes. With these systems, the images can be coherently reconstructed locally (within a 10 x 10 pixel area). However, this also reduces the illumination conductance by a factor of 100 and thus significantly reduces the permissible illumination intensity.

[0069] However, due to coherent multi-mode illumination, the illumination conductance can be maximized without compromising coherent reconstruction. Therefore, significantly faster and more sensitive OCT measurements can only be achieved with coherent multi-mode illumination.

[0070] In a confocal system, if the lateral resolution needs to be increased, it is possible in certain embodiments to adjust the illumination pupil magnification using a telescope and to illuminate and detect the retina with larger pupil diameters. Alternatively, it is also possible to increase only the illumination aperture / pupil diameter by means of the diffuser 100 in the illumination system.

[0071] In a confocal system, sensitivity outside the focal plane decreases with NA. 4 (NA = numerical aperture). To double the lateral resolution, the NA must be doubled.

[0072] For an asymmetrical system with an illumination aperture enlarged by the diffuser 100 and a detection with a smaller, unchanged numerical aperture, the sensitivity outside the focal plane decreases only with NA. 2 However, to double the resolution, the illumination aperture must be increased approximately fourfold. Therefore, the sensitivity drops off outside the focal plane, similar to a confocal system with a correspondingly higher NA.

[0073] The basic principle of the evaluation can be described as follows.

[0074] The detector covers, for example, a pupil size of 0.65 mm. This corresponds to a focus size of approximately 35 µm on retina 2 for a wavelength of around 1 µm. The illumination pupil is increased to, for example, 4.5 mm by the diffuser 100. The diffuser 100 is designed so that the illuminated retinal area also has a diameter of approximately 35 µm. Due to wave scattering, the 35 µm illumination spot is modulated in intensity by approximately 37 speckle grains. If two images of retina 2 are now taken, shifted from each other by one speckle grain, the two measured images are mathematically independent of each other due to the interaction with retina 2. For this reason, retina 2 is illuminated with strongly overlapping spots, and in the illustrated case, the... Fig. 1A is achieved with a 5 µm scanner feed between two OCT-A scans. The result, after evaluation, is a 3D image with a lateral sampling depth of 5 µm, which, however, does not yet have the specified resolution because the object function is still convolved with the illumination speckle pattern. If this speckle pattern is known for the focal plane, it can also be numerically propagated to all other planes using the waveguide equation. Thus, in the final image, any 2D depth-slice image can be coherently unfolded with the known speckle function. In principle, all spatial frequencies up to the cutoff frequency are reconstructed, but some spatial frequencies are attenuated due to the speckle nature. Since confocally acquired retinal images also exhibit such a speckle structure, this is often acceptable in applications.

[0075] To enable coherent reconstruction of the images, it is preferable that the absolute phase of the signals for capturing all superimposed retinal spots is not distorted by motion artifacts. For typical retinal movements, this can be ensured for approximately 100 µs. With a 100 kHz OCT system (A-scan rate), approximately ten individual measurements can therefore be performed before this phase stability is no longer maintained. In the case described, this means that the approximately seven overlapping spot images of a B-scan are phase-coupled with each other, but the OCT B-scans exhibit statistical relative phases with each other, thus making coherent reconstruction impossible. For this reason, a different design is preferred in the construction of the Fig. 1A The corresponding sample area of ​​the retina is scanned again with reversed lateral directions (perpendicular scan pattern) every approximately 5..10 A-scans after an initial scan, in order to locally estimate the phase relationships of the B-scans with this redundant information and to correct them locally before a coherent reconstruction.

[0076] For systems designed for use on the living eye, the typical optical aberrations of the eye significantly limit the achievable resolution for pupil diameters above approximately 1.3 mm. Prior art in holoscopy has established methods for estimating and subtracting these ocular aberrations from the data. During the evaluation process, the aberrations are estimated and corrected iteratively. Thus, such systems not only provide a sharp image of the retina but also reveal the wavefront errors of the measured eye.

[0077] If one wants to use an OCT system as a wavefront sensor, the optimization goal is to be able to measure the largest possible aberrations for a given pupil diameter. Since the coherent multi-mode illumination method is asymmetrical and uses only one large illumination aperture / pupil diameter, a significantly increased detection range for aberration intensity is obtained compared to a symmetrical confocal high-NA OCT system.

[0078] A method known in the art is multi-beam Doppler. This method allows for the quantitative measurement of the velocity and direction of flows, such as blood flow in the retinal capillaries, by simultaneously illuminating multiple confocal beams from different subpupils onto a single retinal point. These beams are superimposed on the retina in a temporally incoherent manner. However, with a single confocal sensor, regardless of the numerical aperture (NA) used, quantitative 3D blood flow measurement would not be possible due to system symmetry. Thanks to coherent multi-mode illumination, the speckle distribution in and around the focal point now makes it possible to achieve asymmetry-free illumination, enabling quantitative measurement of 3D blood flow velocity with a single confocal sensor.

[0079] According to the invention, the phase element 100 is preferably positioned conjugate to the retina 2 in the illumination beam path. This makes it possible to motorically swivel the diffuser plate 100 out of the beam path, thus enabling confocal OCT, as is standard practice. This allows the system to be flexibly switched between rapid images, lateral high-resolution images, or special imaging techniques such as blood flow measurement or wavefront measurement.

[0080] To measure an OCT image with a specified resolution, all spatial frequencies less than or equal to a corresponding cutoff frequency must be recorded and reconstructed. Therefore, the optical transfer function must be free of gaps, because otherwise the spatial frequencies of the sample falling within these gaps will not be transmitted by the system and consequently cannot be reconstructed. Mathematically, the optical transfer function is the convolution of the illumination pupil function and the detection pupil function. To completely suppress the described vignetting effect, the illumination pupil function and the detection pupil function must be identical. Furthermore, if the transfer function is to be free of gaps, the detection pupil function must also be free of gaps.In systems that sequentially measure multiple subpupils and then coherently reconstruct them as a whole, the transfer function for each subpupil can be represented individually as the convolution of the illumination pupil function and the detection pupil function. The overall transfer function is then the sum of the individual transfer functions. If gaps occur in the detection pupil function during this type of vignetting-free acquisition, spatial frequencies falling within these gaps cannot be transmitted by the system and generate artifacts during reconstruction. Consequently, if there is no mode crosstalk between the different illumination and detection modes, the detection subpupil must completely fill the total pupil to avoid reconstruction artifacts.

[0081] For a given pupil size and an optical aberration level determined by the application, the subpupils / detection channels must not exceed a certain diameter to prevent excessive fringe washout across the subpupils. Furthermore, if there must be no gaps in this detection pupil, this results in a minimum number of detectors that cannot be undercut to achieve a specific lateral resolution.

[0082] If a limited vignetting effect is acceptable for the application, the necessary number of detector elements (19a.x / 19b.x in) can now be determined. Fig. 1B or pixel in Fig. 1C) in the detector device 17 can be significantly reduced if the resulting gaps in the detection pupil function are masked by a gapless illumination pupil function. Due to the interactions of the different illumination modes with the different detection modes, spatial frequencies from the illumination are transferred to the detection via speckle modulation. For this reason, the idea is to accept and / or optimize a certain vignetting effect in exchange for a reduced number of detector elements.

[0083] The laterally high-resolution OCT 1 is designed to detect a pupil size of 4.5 mm (corresponding to a lateral resolution of 2.5 µm for a wavelength of λ=1 µm due to the large illumination and detection pupil) and a retinal field of approximately 60°. Numerical simulations using an eye model demonstrate that, in this case, the size of the subapertures at the pupil margin should not exceed 0.65 mm to limit the fringe washout effect, which would otherwise limit sensitivity. This results in a hexagonal arrangement with 37 detector elements of detector assembly 17 as the minimum gap-free configuration.

[0084] Using the established technique of multi-directional scanning OCT, 37 directions would have to be scanned sequentially to achieve artifact-free imaging at the specified resolution. The effort involved would be considerably greater.

[0085] Fig. Figure 4 schematically shows a representation of the vignetting effect in the so-called “mode mismatch” outside the focal plane. Fig. Figure 5 shows the conditions in the focal plane. The detector is located at position 108, and the illumination at position 110. The hatched areas 112 are not used due to vignetting effects, as the illumination lies outside the dashed area that can be detected at position 108. For clarity, positions 108 and 110 are located at the pupil P, and the vignetting is shown below the retina 2. The areas 112 are illuminated but cannot be detected.

[0086] Fig. Figure 5 shows the distance to the focus at which vignetting does not yet occur. In the central area, which corresponds to the dashed area in Fig. If the value corresponds to 4, no vignetting occurs, regardless of the position relative to the focal plane. Outside the focal plane, vignetting limits the detectable numerical aperture and thus the spatial frequency range accessible to measurement with increasing distance Δz.

[0087] For retinal imaging with such a high lateral resolution of 2.5 µm, only small retinal areas and only the uppermost retinal layers down to the RPE are measured. For this reason, the depth range to be measured can be limited to, for example, Δz = 500 µm. Comparing the area of ​​illumination in Fig. 7 with the detector size, it can be seen that the detector device 17 can also capture a spatial frequency range with vignetting effect, which for a vignetting-free system uses e.g. seven detectors (ring with six detectors around central detector).

[0088] Multi-detector systems with several (2 ... 16) sensor elements 19a.x / 19b.x or pixels are therefore used for the construction of the Fig. 1B or Fig. 1C is preferably combined with coherent multi-mode illumination. Various arrangements are conceivable for this purpose. The specific properties will be discussed below using two different arrangements as examples, which are described in Fig. 6 and Fig. 7 are shown. Fig. Figure 3 shows a system with a multiple detector system that has four detector elements 119.1 to 119.4, Fig. 7 a variant with seven detector elements 119.1 to 119.7. The detector device 17 therefore has, for example, four or seven detector elements.

[0089] The central detector element 119.1 in these arrangements has three important functions. In measurements on living eyes with a small usable iris diameter, the central location (e.g., pixel 119.1) ensures half the possible lateral resolution when the outer locations (pixels) are vignetting by the iris. For all measurements with a sufficient iris diameter, the central location (pixel 119.1) significantly reduces the effective vignetting outside the focus. Near the focus, the data from the central location (detector 119.1) are redundant and can also be independently reconstructed from the other locations (detectors) with their overlapping spatial frequency spectra. Therefore, the data from the central location (pixel 119.1) can be used to reconstruct the absolute phase relationship between the B-scans during overlapping sampling without having to acquire additional perpendicular B-scans for local absolute phase correction.

[0090] In detection with such multi-channel detection, the phase plate 100 should also preferably be arranged conjugate to the retina 2 in order to be able to swivel the scattering plate out motor-driven or to be able to swivel in different scattering plates that are optimized for specific applications as a replacement.

[0091] The in Fig. The illumination shown in Figure 4, with a filled pupil of enlarged diameter, is particularly suitable for lateral high-resolution measurements without reconstruction artifacts caused by gaps in the spatial frequency spectrum. Compared to a confocal system retrofitted with a diffuser 100, the vignetting effect is significantly reduced because the multiple detectors 119.1–119.7 better cover the pupil. However, vignetting effects will still occur far outside the focus.

[0092] For example, if one wants to measure the application-relevant blood flow in the vessels of the choroid and retina, larger measurement depth ranges may be necessary. To completely suppress the vignetting effect, a phase plate optimized for this purpose should be designed so that no filled scattering pupil is generated, but rather a single illumination mode is generated precisely at the assigned detector position (with an initial phase according to the main requirement).

[0093] She is exemplary in Fig. Figure 8A shows a section of the image and produces a fully filled scattering pupil and an asymmetrical field distribution around the focal point. It is particularly suitable for high-resolution applications. If necessary, it may then be required to distribute the detector channels differently, e.g., as asymmetrically as possible, within the detection pupil. Fig.Figure 8B shows a scattering plate that produces one unscattered and three scattered illumination modes. In both images, the phase delays are coded in gray.

[0094] Of course, segmented detector arrays that completely fill the detection pupil, such as a 37-pixel hexagonal detector array, can also be used. Systems with 2 to 100 separate detection channels can also be implemented, particularly favorably, as photonic integrated circuits.

Claims

[1] Method for holoscopic optical coherence tomography of an object (3), in particular an eye, wherein the method comprises - Providing source radiation and splitting the source radiation into illumination radiation (B) and reference radiation (R), - Illuminating a lighting field extending horizontally across the object (3) with the lighting radiation (B), - Collecting backscattered illumination radiation from the object (3) as measuring radiation (M), - Separating the measuring radiation (M) collected by the object (3) from the illumination radiation (B) directed to the object (3), and - Superimposing the measurement radiation (M) with the reference radiation (R) and detecting an interference signal of the superimposed radiations with at least one detector (17), characterized by , that - the object (3) in the illumination field is illuminated simultaneously with more than one spatial radiation mode, - where the radiation modes of the illumination in the illumination field are spatially and temporally coherent to each other, but have a fixed phase difference to each other. [2] Method according to claim 1, characterized by that the phase differences are distributed statistically or quasi-randomly across the illumination field. [3] Method according to claim 1 or 2, characterized by , that confocal detection is performed with the detector (17). [4] Method according to claim 1 or 2, characterized by , that the detector (17) comprises a pair of balanced detectors (19a, 19b). [5] Method according to claim 1 or 2, characterized by , that the detector comprises a one-dimensional or two-dimensional spatially resolving detector (19c). [6] Method according to claim 5, characterized by, that the area detector has adjacent detector elements (119.1-119.7) in an image plane, wherein the number of detector elements (119.1-119.7) and the number of illumination modes are matched. [7] Method according to claim 6, characterized by , that the number of detector elements (119.1-119.7) is equal to the number of illumination modes and the illumination modes are assigned to the detector elements (119.1-119.7) 1:

1. [8] Method according to claim 5, characterized by , that the number of detector elements (119.1-119.7) is less than the number of illumination modes and that information reconstruction is carried out from spatial frequencies that are transferred from the illumination to the detection by a mode interaction. [9] Method according to any one of claims 1 to 8, characterized by, that by detecting the interference signal of the superimposed radiations several individual images are taken by the detector (17), between which the illumination is shifted by at least one illumination speck grain relative to each other, and that the individual images are combined to form a total image, preferably by averaging or by setting up and solving a system of equations via linear regression. [10] Method according to any one of claims 1 to 8, characterized by , that the object (3) in the illumination field is illuminated with a speckle pattern that has no axis of symmetry. [11] Optical coherence tomograph for holoscopic examination of an object (3), in particular an eye, which has: - a lighting device (4, 5) for providing source radiation, - an illumination and measuring beam path (7) which has a splitting element (6) for splitting the source radiation into illumination radiation (B) and reference radiation (R), with which the illumination radiation (B) illuminates an illumination field extending horizontally across the direction of incidence on the object (3) and collects backscattered illumination radiation from the object (3) as measuring radiation (M), - 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 a reference beam path (8) and superimposes them at a superposition point and directs them onto a detector (17), characterized by , that - the illumination and measuring beam path (7) illuminates the object (3) in the illumination field simultaneously with more than one spatial radiation mode, - where the radiation modes of the illumination in the illumination field are spatially and temporally coherent to each other, but have a fixed phase difference to each other. [12] Coherence tomograph according to claim 11, characterized by , that the detector (17) includes a confocal detector (19a, 19b). [13] Coherence tomograph according to claim 11, characterized by , that the detector (17) comprises a two-dimensionally extended area detector (19c), in particular has 2 to 100 detector elements and is optionally implemented as a photonically integrated circuit. [14] Coherence tomograph according to claim 13, characterized by , that the area detector has adjacent detector elements (119.1-119.7) in an image plane, wherein the number of detector elements (119.1-119.7) and the number of illumination modes are matched. [15] Coherence tomograph according to claim 13, characterized by, that the number of detector elements (119.1-119.7) is equal to the number of illumination modes and the illumination modes are assigned to the detector elements 1:

1. [16] Coherence tomograph according to claim 13, characterized by , that the number of detector elements (119.1-119.7) is less than the number of illumination modes and the coherence tomograph has a control unit (C) that performs information reconstruction from spatial frequencies transferred from illumination to detection by mode interaction. [17] Coherence tomograph according to any one of claims 11 to 16, characterized by , that the illumination and measuring beam path (7) illuminates the object (3) in the illumination field with a speckle pattern that has no axis of symmetry. [18] Coherence tomograph according to any one of claims 11 to 17, characterized by, that the illumination and measuring beam path (7) has a scattering element (100) for generating the illumination modes, which is located in the illumination beam path, preferably in a field plane. [19] Coherence tomograph according to claim 18, characterized by , that the spreading element (100) is swivel-out.

Images