DE102014115155A1 - Optical coherence tomography for measurement at the retina - Google Patents

Abstract

An optical coherence tomograph for examining an eye (3) an illumination device (4, 5) for providing source radiation whose wavelength is tunable, an illumination and measurement beam path (7) having a distribution element (6) for splitting the source radiation into illumination radiation (B ) and reference radiation (R), illuminated with the illumination radiation (B) an illumination field in the eye (3) and in the eye (3) backscattered illumination radiation as measuring radiation (M), wherein the illumination and measuring beam path (7) has a front optics (12 ) and the pupil (P) of the eye (3) is a pupil of the illumination and measurement beam path (7), a reference beam path (8) which passes the reference radiation (R) through a delay path (21), 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 superimposed a at least one detector (19, 19a, 19b) conducts, wherein a beam splitter (11) for separating the measuring radiation (M) collected by the eye (3) from the illumination radiation (B) guided to the eye (3) transmits the separated measuring radiation (M) leads to the detection beam path (14, 15, 17), an only on the illumination radiation (B) acting optical element (10.1, 10.2), the pupil (P) of the eye (3) a first pupil area (72) illuminates and one only on the Measuring radiation (M) acting optical element (14.1, 14.2) the measuring radiation (M) from a second pupil area (70) of the pupil (P) of the eye (3) collects, wherein the first and second pupil area (72, 70) do not overlap and are separated from each other by a neutral region (71) of the pupil (P) of the eye (3) which belongs neither to the first nor to the second pupil region (72, 70).

Description

The invention relates to an optical coherence tomograph for examining an eye, which has an illumination device for providing source radiation whose wavelength can be tuned, an illumination and measurement beam path having a distribution element for splitting the source radiation into illumination radiation and reference radiation with the illumination radiation Illuminated illumination field in the eye and in the eye backscattered illumination radiation as measuring radiation, wherein the illumination and measuring beam path has a front optics and the pupil of the eye is a pupil of the illumination and measuring beam path, a reference beam, which passes the reference radiation through a delay line, 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 superimposed on at least one detector passes.

The invention further relates to a method for optical coherence tomography for examining an eye, wherein source radiation is provided and tuned in wavelength and divided into illumination radiation and reference radiation, an illumination field in the eye is illuminated with the illumination radiation, backscattered in the eye illumination radiation is collected as measurement radiation wherein a front optic is used and the pupil of the eye is a pupil of the illumination and the collection of the measurement radiation, and the reference radiation is superimposed with the measurement radiation to produce an interference signal which is detected by a detector.

Optical coherence tomography (OCT) is an established method of eye imaging in ophthalmology. It allows a three-dimensional image, which is very helpful in the diagnosis of eye diseases and their course. In particular, diseases of the retina, such as glaucoma or age-related macular degeneration, may be mentioned here. 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 usually much larger than the depth of field of the image, which in turn depends on the numerical aperture, more precisely proportional to 1 / NA ² . In the commonly used Fourier domain OCT employing 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.

Measurement of the retina of the human eye requires both high lateral and high axial resolution. At the same time the detectable and thus illuminated volume in the depth (along the optical axis) should be as large as possible; This causes a small numerical aperture (NA) of the optical system. The lateral resolution requires a large numerical aperture. Thus, in the prior art, ultimately, the extent of the depth-accessible region and the lateral resolution are linked by the numerical aperture of the optical system and can not be adjusted independently.

From the US 2014/0028974 A1 For example, an imaging method based on OCT is known. A line is projected onto an object through an imaging system. The backscattered radiation is interferingly combined with reference radiation and directed to a detector with confocal filtering in one direction. For this purpose, an astigmatic optics is used. The depth resolution takes place by means of optical coherence tomography. In the case of spectroscopic analysis of the radiation, a two-dimensional detector is used whose expansion serves the purpose of confocal filtering with respect to the line-shaped illuminated area and whose other extent dissolves the spectral information. The combination of lateral resolution and accessible depth range is also in the approach according to US 2014/0028974 A1 given.

OCT imaging is affected by essentially three noise effects. A first noise source is abbreviated RIN, which stands for "relative intensity noise" and refers to incoherent modulations / interactions at different frequencies in an optical wave. Ultimately, the cause of this noise is in the radiation source, and one way to suppress the RIN is to use a special detector to measure the intensity noise and correlate it to the signal. A structural common solution for this is the so-called "balanced detection". Next, the so-called off-axis detection is known as a possible solution. Ultimately, these approaches improve the dynamic range of the measurement signal and thus the signal-to-noise ratio.

A second noise source is the so-called "beat noise". It can be considered as a special form of RIN, which is caused by incoherent interaction between the reference radiation and the incoherent radiation of the measurement radiation. Even with a complete suppression of the RIN, for example, by a balanced detection, this noise is present.

A third noise source is the so-called coherence revival, which can be annoying with certain tunable sources. Due to secondary maxima (so-called side lobes) of the coherence function, partially coherent interactions between the reference radiation and actually incoherent parts of the measuring radiation or reflected illumination radiation can occur. This can be avoided in the prior art only by a very complex construction of the optical structure for suppressing optical reflections in the measurement beam path. Care must be taken to ensure that reflections that could fall into the window of the secondary maxima of the tuned radiation source do not occur.

The object of the invention is to disclose an optical coherence tomograph, in particular for measuring at the retina of the human eye, in which the noise behavior is improved.

The invention is defined in claims 1 and 9. Advantageous developments are the subject of the dependent claims 2 to 8 and 10 to 16.

According to the pupil of the eye is not uniformly illuminated, but there is a certain type of pupil separation instead. In the pupil, two areas are provided. A first pupil area is used for illumination, and a second pupil area for collecting the measuring radiation. The pupil areas do not overlap and are separated from each other by a neutral area in the pupil, which is used neither for the illumination, nor for the collection of the measuring radiation, ie the imaging of the object in the eye, for example the retina.

In a preferred embodiment of these pupil regions, the second pupil region provided for imaging surrounds the neutral region, which in turn surrounds the first pupil region of the illumination. As a result, the illumination is guided through inner regions of the pupil and the measuring radiation through outer regions of the pupil. This improves the imaging quality and (for example in the case of a scanning holographic OCT system) the lateral resolution and at the same time achieves illumination over a large depth range.

The neutral region, which separates the first pupil region from the second pupil region, makes it possible to effectively suppress noise effects, in particular the said "beat noise" and the "coherence revival". The neutral region automatically leads to the fact that even at a certain distance from the pupil the illumination radiation which passes through the first pupil region does not overlap with measurement radiation which is collected from the second pupil region. Without the neutral range, such an overlap would be present at any arbitrarily small distance from the plane in which the pupil lies. The neutral region thus ensures an axial section in which illumination radiation and measuring radiation do not yet overlap laterally. In this axial section optical elements may preferably be arranged, which guide the beam path. Due to the lack of overlap between illumination radiation and measuring radiation in these axial sections, it is almost impossible for illumination radiation to erroneously pass through reflections into the measurement radiation. The axial section in which illumination radiation and measuring radiation do not yet overlap laterally is referred to as a pupil-near section of the beam path or a section near the pupil plane in the sense of this description. In a preferred embodiment, it is provided to arrange optical elements which form the beam path, in particular the frontal optics, in such pupil near axial sections of the beam path.

The illumination pattern of the pupil can be generated in a variety of ways, for example by corresponding apertures in one or more conjugate pupil planes that are conjugate to the pupil plane of the eye. Since the optical coherence tomograph has sections in its illumination and measuring beam path in which exclusively illumination radiation or measuring radiation propagates, it is particularly preferred to create a conjugate pupil plane in each of these sections and to arrange here the necessary diaphragm for defining the first pupil region and the second pupil region ,

An area detector, preferably a two-dimensional detector, which scans a partial area of the retina, is preferably used. Preferably, a diaphragm in an intermediate image of the optical image defines this partial region, and the surface detector is matched to the size of the diaphragm. The beam path is designed so that the illumination with illumination radiation and the collection of the backscattered measuring radiation with different numerical apertures. Thus, a numerical aperture can be set for the illumination, which illuminates an axially large area, so that measuring radiation is reflected from a comparatively large depth range and consequently an image over a large depth range is obtained by means of the OCT principle. Irrespective of this, the numerical aperture of the collection of the measuring radiation, that is to say the imaging of an object area, is independent of the numerical aperture of the illumination and thus larger, for example, so that a high lateral resolution is obtained.

The area detector is preferably a two-dimensional detector. The number of pixels is preferably between 4 and 100 pixels per direction, more preferably between 5 pixels and 40 pixels. This number of pixels proves to be advantageous for scanning the field both in terms of resolution and in terms of signal-to-noise ratio and possible aberration corrections.

Image aberration of particular importance is the aberrations generated by the eye. Since the numerical aperture of illumination and detection are decoupled, it is possible for the detection, i. H. To perform the imaging of the retina object area with a very high numerical aperture, which is so large that aberrations of the eye play a significant role. The spatial resolution of the area detector, as explained below, allows correction of the aberrations when the area detector is arranged in a conjugate pupil plane of the image. If the area detector is not in a pupil plane, an aberration correction is equally possible if the detected signal is converted to a pupil plane, as is known to a person skilled in the art for holograms.

In the object plane and the image planes of a ray path, the image information is pure location information. Shown structures can be found as intensity differences also in intermediate image planes again. In pupil planes, the image information is pure angle information. The angles of the incident rays encode the image information. This is the known effect that a cross-sectional change in a pupil affects only the image brightness, but not the image size. For this reason, in the human eye, the iris is in the pupil plane, so that by narrowing or widening the iris, the human eye adapts itself to the brightness. As far as in this description of the level of the pupil of the eye is mentioned, the iris plane is meant. An imaging beam path images an object from the object plane onto an image in the image plane (eg the location of a detector). Between, for example, the object plane and an intermediate image plane, there is always a pupil due to the mapping laws. Likewise, there is always an intermediate image plane between two pupil planes. Similarly, in this specification, planes located between the pupil plane of the eye and the detector are referred to as conjugate pupil planes, as they are predefined by the optically imaging elements to the plane of the pupil of the eye. As far as the retina is mentioned here as object, this is not intended to limit the invention. Other structures of the eye can also be imaged as an object.

Some embodiments of the invention use a confocal aperture. In this description, the term "confocal" not only designates a diaphragm which lies exactly in an (intermediate image) plane conjugate to the object plane, but also detects an arrangement of the diaphragm which lies within a certain error range in front of or behind an intermediate image plane. If the confocal aperture is not exactly in the intermediate image plane, but close to the intermediate image plane, then stray light suppression may possibly be reduced, but the function as a confocal diaphragm, which defines the object field from which the measurement radiation is collected, is likewise fulfilled. The aperture is at or near an intermediate image plane when spaced by a maximum of three times the imaging depth from the intermediate image plane; a spacing of at most the simple depth of picture is preferred. Depth of field is referred to in the English literature as "depth of focus" and defines an axial region in the image space, ie at the intermediate image plane of an optical system, in which a sufficiently sharp image is formed in each capture plane. In the area of picture depth, circle of confusion is registered as point. The area in the object space conjugate to the image depth is the depth of field. The depth of field is a measure of the extent of the sharp area in the object space and is given by the lambda / (NAo) ² , where NAo denotes the numerical aperture in the 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) ² ; where NAz is the numerical aperture at the intermediate image plane, which is z. B. from NAo calculated by means of the magnification. In the above consideration, the maximum wavelength of the measuring radiation at the intermediate image plane can be used as the wavelength.

The features of the optical coherence tomography described below can be used for different embodiments alone or in different combinations. As far as the following embodiments describe certain combinations of features, the invention is not limited to such combinations.

It is understood that the features mentioned above and those yet to be explained below can be used not only in the specified combinations but also in other combinations or alone, without departing from the scope of the present invention.

The invention will be explained in more detail for example with reference to the accompanying drawings, which also disclose characteristics essential to the invention. Show it:

1 a schematic representation of an optical coherence tomograph (OCT) in a first embodiment,

2 a schematic representation of an OCT in a second embodiment,

3 a schematic representation of an OCT in a third embodiment,

4 Illustrations to illustrate a pupil division, which in one of the OCT of 1 to 3 is used,

5 and 6 Top views of the pupil of the eye with a representation of pupil areas formed there,

7 a schematic representation for illustrating the scanning principle in an OCT according to one of 1 to 3 and

8th a schema representation similar to the 7 to illustrate the generation of a three-dimensional image.

1 shows an OCT 1 , the three-dimensional images of a retina 2 one eye 3 receives. Source radiation of a wavelength-tunable radiation source 4 , For example, a corresponding laser, is in a fiber 5 coupled. The source radiation is, for example, in the infrared wavelength range. In the following description also this wavelength range is referred to as "light". This term subsumes all radiation of the electromagnetic spectrum which satisfies the laws of optics.

The fiber 5 flows into a splinter 6 that the source radiation into a measuring arm 7 and a reference arm 8th divides. To the splinter 6 closes in the measuring arm 7 a fiber 9 on, and exiting the fiber end illumination radiation B is by means of a lighting optical system 10.1 to a beam splitter 11 directed. From there it reaches a front optic 12 which focuses the illumination radiation B into a focus on the retina 2 of the eye 3 lies. The illumination optics 10.1 and the front optics 12 Among other things, they set the numerical aperture NA, with which the eye 3 is illuminated. Between the beam splitter 11 and the front optics 12 there is a scanner 13 who has the focus on the retina 2 biaxially perpendicular to the direction of incidence, ie deflected laterally. The coordinates of this deflection are denoted below by x and y. The z-position of the focus can be adjusted by adjusting the front optics 12 be set. This is indicated schematically by a double arrow.

The illumination radiation in the illumination focus on the retina 2 is scattered back from different depths z within the depth of field. This depth of focus range is defined by the numerical aperture NA, which is due to the interaction of front optics 12 and illumination optics 10.1 as well as the optical properties of the eye 3 is fixed.

The backscattered radiation is from the front optics 12 collected as measuring radiation M. To distinguish the incident illumination radiation and that of the front optics 12 collected, backscattered measuring radiation M are these in 1 registered differently. The illumination radiation is drawn with solid lines in the figure, the measuring radiation M with dotted lines. The of the front optics 12 Collected measuring radiation becomes a scanner 13 directed. Here it is descanned, so after the scanner 13 the measuring radiation M is present as a stationary beam.

The collection of the measuring radiation M is an image of the retina 2 , The beam splitter 11 separates the measuring radiation M from the illumination radiation B and passes them to a detector optics 14.1 , The detector optics 14.1 lays down with the front optics 12 and the optical properties of the eye 3 as well as any further imaging elements in the imaging beam path (eg lens 16 ) the numerical aperture NA of the image of the retina 2 firmly. 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 10.1 and the front optics 12 established. The numerical aperture of detection by the detector optics 14.1 and the front optics 12 ,

The detector optics 14.1 focuses the measuring radiation M in an intermediate image plane, in which a diaphragm 15 located. This aperture 15 Sets the size of the object field from which it attaches to the retina 2 the measuring radiation M is detected. Taking into account the imaging scale of detector optics 14.1 , Front optics 12 and eye 3 corresponds to the size of the aperture 15 exactly the size of the object field at the retina 2 from which measurement radiation M is collected, which is thus imaged.

Another look 16 after the aperture 15 directs the measuring radiation M on a detector device 17 , In the embodiment of the 1 includes the detector device 17 a beam splitter / combiner 18 as well as two area sensors 19a and 19b , The area sensors 19a . 19b are in size to match the aperture 15 and the intermediate optics 16 designed. They have a spatial resolution, ie they allow a resolution of the intensity distribution over the beam cross section.

At the beam splitter / combiner 18 will also reference radiation R from the reference arm 8th coupled. This points after the splitter 6 a fiber 20 on. The reference arm 8th has at the in 1 In the embodiment shown, a path length adjusting device 21 which serves the length of the reference arm 8th suitable for the position of the retina 2 of the eye 3 adjust. For this, the radiation from the fiber 20 decoupled and via a retroreflector 22 whose position can be adjusted, such as the double arrow in 1 suggests. About another deflection mirror 23 as well as optics 24 . 25 the reference radiation R becomes the beam splitter / combiner 18 passed, the reference radiation R with the measuring radiation M superimposed on the surface sensors 19a and 19b passes.

The path length adjustment device 21 is in 1 designed as a free jet. This is just as optional as using a retroreflector 22 , In the prior art, various measures are known to adjust the optical path length of a beam.

The detector device 17 is in 1 executed as so-called "balanced detection". Again, this is optional, as shown below 2 will be explained. The balanced detection has the advantage that a common mode component in the superimposition of reference radiation R and measurement radiation M can be suppressed in a particularly simple manner. Alternatively, such a suppression could be dispensed with, if only one of the detectors, for example the detector 19b would be used and the beam splitter / combiner 18 is executed as a pure Strahlvereiniger.

The interference between reference radiation R and measurement radiation M is converted to produce an image, as is known for optical coherence tomography. Since the wavelength of the source radiation is tuned, the Fourier domain principle is used in the image generation, which principle is known from the prior art.

To perform the imaging, the OCT points out 1 a controller C, which receives a signal about the wavelength tuning and the measurement signals of the area detectors 19a . 19b receives. Optionally, the controller C controls the radiation source 4 For wavelength tuning, therefore, knows the currently prevailing in the system wavelength and can thus assign the measurement signals accordingly. The area detectors receive measurement radiation M from an object field at the retina 2 passing through the aperture 15 is fixed. The area detectors 19a . 19b are in size to the aperture 15 adapted accordingly and scan with their individual pixels, the intensity distribution spatially resolved. Are the area detectors 19a . 19b in an image plane, ie in a plane taking into account the image, which of front optics 12 , Detector optics 14.1 and the other interposed optical elements conjugated to the plane of the retina 2 is, the individual pixels already contain the location information in the image field. On the other hand, the area detectors lie in a conjugated pupil plane that leads to the plane in which the pupil P of the eye 3 lie, conjugate, the detectors detect the intensity distribution in the pupil plane and thus phase information. This too can be used for image reconstruction, as will be explained below.

Essential for the invention is that the scanner 13 the object field in the retina 2 shifts, since it acts not only on the illumination radiation B, but also on the collection of the measurement beam M. At every position of the scanner 13 This creates a partial image of the retina, its resolution by the number of pixels and distribution of area detectors 19a . 19b is determined. As will be explained below, these partial images are combined to form an overall image.

In the construction of the 1 the beam splitter / combiner combines 18 the measuring radiation M from the measuring arm 7 and the reference radiation R from the reference arm 8th , The area detectors 19a . 19b detect the pattern of interference between measuring radiation M and reference radiation R. The corresponding measures for generating such interference, in particular the necessary properties of the radiation source 4 and path length adaptation are known in the art for optical coherence tomographs. Due to the balanced detection, a relative phase difference of Pi lies between the sum of the two signals of the two area sensors 19a . 19b in front.

wenn man mit u_s und u_r die Amplituden und φ_s und φ_r die Phasen der Signale in den beiden Armen bezeichnet (die Indices ”sample” und ”s” beziehen sich auf den Messarm, die Indices ”reference” und ”r” auf den Referenzarm).The complex amplitudes of the measuring radiation and the reference radiation can be written as:

if u _s and u _{r are} the amplitudes and φ _s and φ _{r are} the phases of the signals in the two arms (the indices "sample" and "s" refer to the measuring arm, the indices "reference" and "r" on the reference arm).

The of the two sensors 19a . 19b detected signals 11 and 12 are then I ₁ = | U _sample + U _reference | ² = | U _sample | ² + | U _reference | ² + 2Re {U _sample · U _reference } and I ₂ = | U _sample + U _reference · e ^iπ | ² = | U _sample | ² + | U _reference | ² + 2Re {U _sample · U _reference · e ^-iπ }.

U is complex conjugate to U, and Re is an operator that returns the real part of a complex value. For the difference _signal I _{bd of} the two detectors 19a . 19b you get I _bd : = I ₁ -I ₂ = {U _sample U _reference } - 2Re {U _sample · U _reference ^-iπ · e} = 2 · _s · u u _r · cos (Δφ) - 2 · _s · u u _r · cos (Δφ - π) = 4 · _s · u u _r · cos (Δφ) where Δφ: = φ _s - φ _r denotes the relative phase between the measuring and reference arm.

The formulas show that in the difference signal of the two detectors 19a . 19b only the interference pattern of the two signals cos (Δφ) is present and the common-mode components | U _sample | ² and | U _reference | ^{2 are} suppressed.

In this way, in particular in the case of a possible subsequent analog-to-digital conversion of the difference signal, the dynamic range of the signal is used to a maximum for information evaluation.

In another embodiment, there is no balanced detection; the signal swing of the interference signal is then modulated on a common mode component and is filtered out by appropriate data analysis.

2 shows a modified construction of the OCT 1 that in many aspects of the 1 like. Like elements bear the same reference numerals as in FIG 1 , The main difference lies in the design of the detector device 17 that are in the construction of the 2 only a single area detector 19 having. On this surface detector 19 the measurement radiation M and the reference radiation R are brought to interference at an angle to each other. The angular offset results in a phase shift between pixels lying along the plane which is spanned by the optical axis along which the measuring radiation M is incident and by the optical axis along which the reference radiation R is incident. This phase shift can be evaluated to suppress the common mode component. Such a detection arrangement is referred to as off-axis detection and is known in the art for common-mode rejection.

3 shows a further embodiment for the OCT 1 Here, the Weglängenanpasseinrichtung not in the reference arm 8th but in the measuring arm 7 is arranged. After the fiber 9 and the illumination optics 10.1 there is a Weglängenanpasseinrichtung 29 , again purely by way of example by means of a movable retroreflector 30 realized. The embodiment of the 3 shows that it does not matter if the path length adjustment device in the reference arm 8th or in the measuring arm 7 lies. It is also possible to provide in both a Weglängenanpasseinrichtung. It is only essential that the interference state between the reference radiation R from the reference arm 8th and the measurement radiation M can be set so that it is the retina, for example, the retina, for the current measurement task, ie the actual position of the object to be measured, in the exemplary embodiments described here 2 of the eye 3 , is adjusted.

In 3 Two additional features are shown, each for embodiments of the OCT 1 can be used. The front optics 12 is two-part by two imaging elements 12a and 12b educated.

The scanner 13 is located at the OCT 1 of the 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 to the pupil P plane of the eye 3 conjugated.

The front optics 12 includes the optics 12a and 12b , which together form a 4f optic. Thus, the optics 12a an ophthalmoscopic lens and the optics 12b a scan lens. This 4f optic forms the pupil P of the eye 3 in a pupil plane conjugate to the plane of the pupil P, in which the scanner 13 lies. It is not mandatory to use the scanner 13 However, to place exactly in this conjugated pupil plane has advantages. Between the level of the pupil P of the eye 3 and the conjugate pupil plane is an intermediate image plane 26 , The beam splitter 11 is due to its proximity to the scanner 13 also close to the conjugated pupil plane. It is also possible to use the beam splitter 11 to put in this conjugated pupil plane when the scanner 13 is moved from the conjugated pupil plane.

In one embodiment, the beam splitter 11 designed as polarization divider. He is then in the imaging direction a lambda / 4-plate 27 upstream. This embodiment will be discussed below.

The detector optics are also designed as 4f optics. It represents another intermediate image plane 28 ready, in which the (in 3 not shown) aperture 15 lies. The intermediate image plane 28 is to the object plane, in which the retina to be imaged 2 lies, conjugates. The (not shown) aperture 15 is in its size taking into account the magnification, which for the generation of the intermediate image plane 28 relevant to the size of the imaged area on the retina 2 ,

The aperture 15 has two functions in all embodiments. First, it suppresses stray light, causing the contrast at the detector device 17 is improved. Ultimately, the aperture in this regard is similar to a confocal aperture for confocal scanning OCT 17 due to the effect of the detector optics is preferably in a plane which is conjugate to the pupil plane of the eye, or in the vicinity of this plane. This arrangement is advantageous, but not mandatory. 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 or the area detectors 19a . 19b is due to the object field size on the retina 2 and ultimately the size of the aperture 15 in the intermediate image plane 28 specified. The aperture 15 On the other hand, this ensures a particularly favorable signal acquisition.

In preferred embodiments of the OCT, the area detector has a pixel number of 4 to 100, preferably 5 to 50, particularly preferably 5 to 40 pixels in each direction. In the prior art holoscopic OCT systems are known which have detectors with 100 to 4000 pixels per direction. These pixel numbers are deliberately not used here. The number of pixels is linked to the necessary illumination brightness, the measuring speed and the suppression of multiple scattering.

As already explained above, the image information is present in the form of angle information in a pupil of the beam path, and the intensity distribution in the pupil is generally completely uniform. The OCT 1 of the 1 to 4 Now deviate from this rule to the effect that the pupil is not illuminated in the same way when the illumination (introduction of the illumination radiation B) and imaging (collection of the measurement radiation M). For this purpose, for example, a diaphragm 10.2 after the illumination optics 10.1 and a panel 14.2 in front of the detector optics 14.1 arranged as exemplified in the 1 to 3 is drawn. An alternative for such diaphragms is, for example, a suitably designed beam splitter 11 , The In the following, structuring of the pupil P of the eye with regard to illumination radiation B and measuring radiation M explained in greater detail is not only possible by means of diaphragms or beam splitters.

4 shows in each case the eye in a left and a right part of the representation 3 , from its retina 2 a section with the extent x is mapped. The dash of length d in the pupil illustrates that rays emanating from the top of the stroke (dotted) overlap rays emanating from the bottom of the stroke (dashed) only when a minimum distance from the pupil plane has been exceeded. In front of the pupil there is no overlap in an axial section y, which lies in front of the pupil P. In this axial section of length y, rays emanating from the upper point of the line of length d are laterally separated from rays emanating from the lower end of the line of length d. The same applies to an axial section z within the eye 3 , The line of length d illustrates that in the pupil P a lateral distance between illuminating radiation B and measuring radiation M results in that they do not overlap in an axial section of length y in front of the pupil plane and an axial section of length z behind the pupil plane. In the section y arranged optical elements, such as a front optics 12 , For physical reasons, it is thus not able to reflect undesired illumination radiation B in the area from which measurement radiation M has been collected. To clarify this effect was in 5 the radiation emitted by the upper end of the line d is deliberately drawn in dotted lines, as is the measuring radiation M in the representations of the 1 and 2 ,

However, the described antireflection effect is not limited only to portions y in front of the pupil plane, but also results for the axial portion z within the eye 3 , In this axial section, for example, lies the eye lens (in 4 not shown). Also on her no illumination radiation B can be reflected so that it is collected as measurement radiation M.

The axial sections y and z are, of course, not only on both sides of the plane in which the pupil P of the eye 3 is located, they are also found in the beam path on both sides of each conjugated pupil plane. Thus, it is preferred optical elements of the observation and measuring beam path 7 and in total the OCT 1 if possible, place it in areas that lie within the zone around a conjugated pupil plane defined by the distances y and z from the conjugated pupil plane - of course taking into account a possible magnification which alters the magnitudes for z and y. In particular, it is preferable to the scanner 13 in such a zone to place a conjugated pupil plane.

The length of the axial sections z and y, in which no overlap of the radiation beam takes place, results as follows: z = (d * L) / (d + x) or by y = (d * L) / x (at an approximate assumption of the refractive index of all elements of n = 1). Of course, the line of length d is only symbolic of the distance between two points in the plane of the pupil P. In the described embodiments, it is now implemented in such a way that a first pupil region through which the illumination radiation B falls and a second pupil region through which the measurement radiation M is collected, ie the retina 2 is provided, a neutral range is provided, which is functionally the distance d between the upper point of the stroke and the lower point of the stroke in 4 equivalent. The 5 and 6 show examples of such a pupil separation. There is a first pupil area in each case in the left part of the figures 72 to see through which the illumination radiation B falls. The pupil area 72 is characterized by hatching running from bottom left to top right. Next is a second pupil area 70 provided by the image of the retina 2 he follows. Between the first pupil area 72 and the second pupil area 70 there is a neutral area 71 through which neither illumination radiation B falls nor the image of the retina 2 , So the collection of the measuring radiation M takes place. The width of the neutral region, measured as the distance between the nearest edges of the first pupil region 72 and the second pupil area 70 corresponds to the stroke of length d in 4 , The right parts of the 5 and 6 show in plan view the retina 2 and the area pictured there 74 , Its shape is due to the shape of the first pupil area 72 given that its Fourier transform defines the illuminated area on the object - and only from the illuminated area measurement radiation M can be backscattered. The backscatter is not spatially directed and therefore can through the second pupil area 70 to be picked up.

To be at an OCT 1 It is preferable to couple the illumination radiation as close as possible to the optical axis so as to provide a sufficiently large section with illumination radiation in an axially sufficient manner, that is to say to have the greatest possible detectable depth range. It is therefore provided, the illumination radiation B through the inner, first pupil area 72 respectively. At the same time, one would like to have the highest possible lateral resolution, especially in the case of a laterally scanning system. For this purpose, a large numerical aperture is advantageous, which is achieved when the image of the retina, ie the collection of the measuring radiation M through the second pupil area 70 which is further away from the axis than the first pupil area 72 ,

The embodiment of the 5 is particularly advantageous for a line-shaped illumination of the retina. It is particularly preferred if the first pupil area 72 has a length of 0.5 to 2 mm and a width of a few microns in the pupil P of the eye and perpendicular to a line of illumination of the retina 2 lies. The second pupil area 70 then consists of two half moon-shaped sections and the neutral area is (except for the first pupil area 72 occupied area) a strip which is at least 400 microns wider than the first pupil area 72 is high.

For symmetrical illumination and imaging of the retina 2 is the arrangement according to 6 advantageous in which the first pupil area 72 a circular disc with a diameter of 0.5 to 2 mm. The neutral range 71 has a width of at least 400 μm, so that the second pupil area 70 has a correspondingly larger inner diameter.

An alternative to using a pupil-blanking element is a detector device 17 , the one or two spatially resolved area detectors 19 ( 2 ) respectively. 19a . 19b ( 1 ) has to arrange this surface detector (s) in a conjugated pupil plane and to read out only those sections of the spatially resolving area detector / the spatially resolving area detectors which are the second pupil area 70 correspond. If the spatially resolving area detector / the spatially resolving area detectors are close or in an intermediate image plane, the electromagnetic field vector can be determined by read-out of the area detector and calculated back into the pupil plane P, whereupon data-filtering is generated in order to limit this to the second pupil area 70 in terms of imaging of the retina 2 to reach. This also suppresses reflexes. In this embodiment, the detector device is part of the element which acts only on the measurement radiation M.

In a preferred embodiment of the OCT 1 aberrations are corrected. The detector device 19 has, as already mentioned, one or two area detectors which have a spatial resolution in the form of pixels. These pixels will also be referred to as channels below. The measuring signal is distributed over these several channels of the detector (s). In a preferred embodiment, if the detector is located in a conjugate pupil plane, each channel of the detector contains measurement radiation M from different angles that are within the retina 2 was scattered. The spatial resolution, which the area detector 19 . 19a . 19b has, allows to detect the distribution of the measuring radiation in the pupil P. Aberrations affect this distribution. From the eye 3 aberrations caused often take an intolerable degree when in the plane of the pupil P of the eye 3 an area is used which is greater than 1.5 mm in diameter. However, such a larger area would be desirable in terms of lateral resolution. Without spatial resolution in the conjugated pupil plane, phase differences would occur with greater pupil utilization on the eye 3 in the then single detection channel mix and average.

The maximum resolvable phase differences depend on the number of channels. It was found that the number of distinguishable phase differences in this plane is given by the number of channels per direction multiplied by Pi. At five channels per direction, polynomials can go up Z m / 4 where m can take the values 0 (sphere), ± 2 and ± 4. This applies to area infinitesimal small channels. In reality, of course, they have a certain size. The measurement signal detected in a channel therefore corresponds to an averaging 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 less than Pi. It was found that at a center wavelength of OCT of 1060 nm for the eye-caused astigmatism evenly distributed channels, the phase differences are recognizable, if five channels the condition 2Pi / (5 channels per aberration period) ≤ Pi is maintained. Then there is a period of minima and maxima 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.

These considerations indicate that an area detector with at least five channels per direction is capable of resolving at least astigmatism and third order aberrations. A higher number of channels allows even higher orders of aberration to be detected.

The above considerations considered only one spatial direction. The aberrations usually have a two-dimensional pattern.

The aberrations cause for each detector channel c a phase θc: U _{sample, c} : = U _sample · e ^iθc . It results from a thickness δd and a refractive index δn of the transmitted material of the eye (eg cornea, aqueous humor, lens, vitreous body), which in reality differs from a theoretical, aberrationless eye: θ _c (k) = .DELTA.n (k) · k · _c .DELTA.D

Thus, the detected signal is shifted by the aberration-related phase: I _{bd, c} (k) = 4 · u _s · u _r · cos (k · Δz - δn (k) · k · δd _c ) = 4 · u _s · u _r · cos (k · (Δz - δ n) k) δd _c ))

For monochromatic radiation of 780 nm, the eye causes wavefront aberrations of up to 0.7 μm, which result in a phase shift of 2 * Pi (leaving aside the defocus). Such a phase shift corresponds to a thickness variation between the lens and the aqueous humor (these are the elements with the largest refractive index differences in the eye), which assumes the following value:

With known dispersion data results: θ _c (λ ₀ = 1060 nm) = 2Pi / 1060 nm x (n _lens (1060 nm) -n _aqueous (1060 nm)) x δd _c = 2Pi / 1060 nm x (1.4104 - 1.3301) x 10 μm = 1.516 Pi or θ _c (k ₀₎ = k ₀ · 0.8034 microns.

If a wavelength range of Δλ = 50 nm is traversed, the phase differences of the associated wavenumbers (k ₀ ± Δk) are: θ _c (k ₀ + .DELTA.k) = 2Pi / 1110 nm · (n _lens (1110 nm) - n _Aqueous (1110 nm)) · .DELTA.D _c = 2Pi / 1060 nm · (1.4099 - 1.3297) x 10 microns = 1.445Pi = (k ₀ + Δk) x 0.8022 μm and θ _c (k ₀ - .DELTA.k) = 2Pi / 1010 nm · (n _lens (1010 nm) - n _Aqueous (1010 nm)) · .DELTA.D _c = 2Pi / 1060 nm · (1.4098 - 1.3305) x 10 microns = 1.594Pi = (k ₀ -Δk) x 0.8048 μm.

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 sweep. Thus, one can write the detected measurement signal as follows: I _{bd, c} (k) = 4 · u _s · u _r · cos (k · (Δz - δn (k ₀ ) δd _c )).

A Fourier transformation for the measured wavenumbers k shows the axial distribution, ie 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 (k ₀ ) δd _c for each channel 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 the channels c are mainly due to the aberrations that displace the profile axially. It is therefore intended to obtain the aberration-caused phases Θ _c (k ₀ ) from the channels relative to a predetermined channel. The measured intensities for a frequency determination are multiplied by the phase factor for correcting the aberrations. The phase factor is

Each channel of the detector has a specific position to the retina 2 , The interference signal can be recorded during the wavelength adjustment of the laser for the respective wavenumber k = 2 * Pi * n / λ, where n is the refractive index of the medium and λ is the wavelength. As in a classical OCT system, the measuring signals are Fourier-transformed with respect to the wavenumbers and the depth distribution of the scattering layers is calculated. In this case, the relationship .DELTA..phi. = K .DELTA.z is used, where .DELTA.z is the distance of a scattering layer to a layer from which the measuring radiation has traveled a path length to the detector which is identical to the path length of the reference radiation.

Due to the lateral extent of the area detector 19 however, the optical path length is not identical for each pixel. For a central channel, the radiation passes along the optical axis. For a further outward channel, the principal ray passes at an angle α to the optical axis, so that a path length which has the value d for the central channel has the value d · cos (α) for an outer channel, each with respect to one Main plane of the eye lens, since its refractive index jump can be used in the measurement as a reference point. Pixels / channels further out collect radiation that traveled a longer distance through the medium. In a reconstruction of the image information affects that. The individual signals are shifted in the z-direction and compressed for further outlying pixels.

The measurement error caused by this effect is corrected in a preferred embodiment in order to obtain a particularly good image acquisition. The geometric effect is preferably corrected by rescaling from z to z cos (α _c ), where α _{c is} the angle that the c-th channel has to the optical axis. The angle α is on a virtual position of the area detector 19 referring to this, taking into account the magnification directly in front of the eye. In an area detector which lies exactly in a plane conjugate to the pupil plane of the eye, the area detector in this way, with an extent modified by means of the magnification, passes exactly into the plane of the pupil P of the eye 3 ,

In the aberration reconstruction, different channels are independently reconstructed. Subsequently, one forms the cross-correlation in the axial direction, d. H. in the depth direction to determine the relative phase offset between each channel. Reconstruction of the lateral image for each channel (possibly under, as will be described below, consideration of the scan) and then the phase gradient provides a lateral offset in the image obtained for a given location of the scanner. This image is also referred to below as the pupil channel sub-image. By means of a lateral cross-correlation of the pupil channel partial image, the aberration is determined in one embodiment, and in this way the entire aberration phase distribution is determined and numerically corrected.

The quality of these approaches depends on the sample structure. The human eye has an easily recognizable axial layer structure. Laterally, the structures are relatively rough, for example, by blood vessels or the optic disc combined with very fine structures, such as photoreceptors, with hardly any structures in terms of size and roughness between them. It is therefore provided in a preferred embodiment that a depth correlation correction is first performed by using the axial layer structure to correct for the largest proportion of the pupil phase aberrations. Optionally, a lateral correlation correction follows, exploiting lateral structures, such as photoreceptors, which became visible as a result of the first correction.

The aberrations of the eye are different at different parts of the retina. In principle, it is possible to calculate the aberration-caused phase changes in each channel for all locations in a lateral image. In a simplified embodiment, it is assumed that the aberrations do not vary very much laterally, and the aberrations are calculated for only a few lateral locations of the retina and interpolated for intermediate locations.

If a relatively large wavelength range is traversed, it is preferable to consider the dispersion of the aberrations. In this embodiment, it is not assumed that the phase shifts vary linearly with the wave number k. It will therefore be a peak in the profiles that in the OCT image of the retina 2 at the fundus of the eye 3 used, to compensate for the displacement of the profiles to each other. One looks for a striking structure and corrects the position of the measurement signals based on this reference point to each other. In this way, the aberrations Θ _c (k ₀ ) can be determined and corrected as described above. Alternatively, a complex correlation algorithm is also possible, which is applied to the profiles of the different channels. In addition to a shift, a scaling (compression or stretching) of the measurement signals can also be corrected.

In a position of the scanner 13 will receive a partial image of the retina, the size of the iris 15 and the front optics involved in imaging the measuring light 12 and detector optics 14.1 is predetermined. A Fourier transform of the signal of the channels provides the image of the sample, but only in a part corresponding to the size of the detector in the pupil. To create a larger image, the scanner is 13 provided the object area on the retina 2 shifts. The image area corresponds to a partial image 59 that is a center 60 Has. For the current distraction by the scanner 13 it is enough to simplify the center 60 of the drawing file 59 To refer to.

For scanning you can the center 60 of the drawing file 59 while tuning the wavelength of the light source 4 leave unchanged. Before a re-tuning, the center becomes 60 shifted so that a new subpicture 59 directly to the previously recorded partial image 59 borders. This way, you can get a bigger picture 61 the retina can be determined. This approach is in 7 shown. As a result, individual partial images 59 to the overall picture 61 together. The images from the individual layers then yield a three-dimensional image of a cuboid area in the retina 2 , this shows 8th , in the example of three levels 62 . 63 and 64 you can see. The drawing files 59 that in the representation of the 8th are associated with a dash-dotted double arrow, each derived from a wavelength sweep at the light source 4 , Because the scanner 13 during each wavelength sweep, and is only shifted therebetween, the fields generated from a wavelength sweep lie 59 in the levels 62 to 64 all with their centers 60 exactly above each other.

The optical structure of the 1 to 3 ensures that the illumination and the recording of the measuring light are no longer coupled with respect to the optical properties and in particular the utilization of the pupil. In this way, it is optionally possible to further adjust the lighting. For example, a Bessel-type lighting can be combined with a top-hat cross-sectional profile for detection. In this way, in one embodiment, a high illumination depth, ie illumination which is unchanged over a large z-range, is achieved with a simultaneously high numerical aperture of the image. For the same numerical aperture, for example, a Gaussian beam would achieve an illumination focus of 1 mm in the z-direction. With a Bessel-type illumination, one obtains 2 to 3 mm extension in the z-direction. In this way, the optical resolution can be increased by 10 to 30% when the detection is done with a top hat-like profile.

In a further embodiment of the OCT takes place at the beam splitter 11 a polarization division. Such is usually disadvantageous in the art, and one uses an intensity division. This is surprisingly advantageous for the OCT described, since polarized radiation entering the eye is changed with regard to its polarization state. Different structures of the eye have a different effect, so that the polarization state of the backscattered signal is not clearly or clearly defined, but consists of components with different polarization states. This consideration was also known in the art and led to the consequence of making an intensity division, just because the backscattered radiation has no clearly defined polarization state. However, it now appears that the measurement light is superimposed with the reference light and can only interfere with each other beam components that have the same polarization state. Ultimately, the reference light, with its state of polarization, therefore specifies which portion of the measuring light can be exploited. Noninterfering components fall on the detector and form a disturbing background.

The illumination radiation B is after the polarization splitter 11 linearly polarized. The lambda / 4 plate 27 as they are in 3 is drawn, ensures circularly polarized illumination radiation B on the eye 3 , Backscattered measuring radiation M, which is also circularly polarized, is again linearly polarized by the lambda / 4 plate, the polarization direction being rotated by 90 degrees with respect to the polarization direction which the illumination radiation B has, that of the polarization splitter 11 was delivered. 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 when reference radiation R and illumination radiation B are identically linearly polarized after being split from the source radiation. This is also the case when reference radiation R and illumination radiation B are circularly polarized after splitting from the source radiation and the reference radiation is linearly polarized before being superimposed identically to the measurement radiation M. Ultimately, it is important that the polarization division (eg by polarization splitter 11 and plate 27 ) the measuring radiation M and the reference beam path condition the reference radiation R such that both radiations at the detector have the same polarization state.

This measure thus increases the signal / noise ratio, since only those parts of the measuring light through the beam splitter 11 to the detector device 17 be guided, who are able to with the Reference light to interfere. Ultimately, the inherently disadvantageous polarization division and the rejection of a portion of the measurement radiation M at the beam splitter increases 11 thus the signal quality.

In a further embodiment of the OCT, use is made of the fact that the illumination optics 10 it allows the focus of the illuminating radiation B to be placed at a different z-position than the focus passing through the detector optics 14 is predetermined for the collection of the measuring radiation M. Due to multiple scattering in the retina, measurement radiation M from the retina may have the distance suitable for interference, but propagate in another direction, which would limit the lateral resolution in depth. Different depth levels for illumination and detection compensate for this effect and the resolution in depth is optimized.

For image reconstruction from the detector signals, one has to know the current wavelength according to the FD-OCT principle. This wavelength or the corresponding wave number k can be obtained from the control of the light source 4 be derived. Alternatively, it is possible to decouple a beam component and detect it with regard to the wavelength in order to better know the currently set wavelength or the course of a wavelength tuning.

Perpendicular to the scan direction, detector channels can be grouped together to reduce speckle. This is particularly advantageous if one wishes only z-sections through the retina.

For a coarsely resolved image, z. B. as a preview, it is possible to combine all or more detector channels. This is done after the corrections (eg aberration, z-position, overall image generation). However, one obtains a resolution as in known OCT systems, however, with a higher signal-to-noise ratio and improved speckle behavior, precisely because combining takes place after one or more of the corrections and thus exceeds normal pixel binning.

If one uses a detector which is only spatially resolving in one direction, aberrations can only be corrected in this direction. This may be sufficient for certain applications.

As far as procedural steps and / or signal corrections have been described above, they are in the OCT 1 performed by the control unit C, which is connected to the detector / detectors whose measurement signals omits and further data about the work of the scanner 13 and receives the wavelength tuning and / or controls these components accordingly.

QUOTES INCLUDE IN THE DESCRIPTION

This list of the documents listed by the applicant has been generated automatically and is included solely for the better information of the reader. The list is not part of the German patent or utility model application. The DPMA assumes no liability for any errors or omissions.

Cited patent literature

• US 2014/0028974 A1 [0005, 0005]

Claims (16)

Optical coherence tomograph for examining an eye ( 3 ), comprising - a lighting device ( 4 . 5 ) for the provision of source radiation whose wavelength is tunable, - an illumination and measuring beam path ( 7 ), which is a distribution element ( 6 ) for splitting the source radiation into illumination radiation (B) and reference radiation (R), with the illumination radiation (B) having an illumination field in the eye ( 3 ) and in the eye ( 3 ) backscattered illumination radiation as measurement radiation (M), wherein the illumination and measurement beam path ( 7 ) a front optics ( 12 ) and the pupil (P) of the eye ( 3 ) a pupil of the illumination and measuring beam path ( 7 ), - a reference beam path ( 8th ), for the reference radiation (R) an optical path length ( 21 ), which corresponds to an optical path length from the splitting element ( 6 ) is equal to the illumination field and back to an overlay location, - a detection beam path ( 14 . 15 . 17 ), the measuring radiation (M) from the illumination and measuring beam path ( 7 ) and the reference radiation (R) from the reference beam path ( 8th ) and superposed on the superimposition site and on at least one area detector ( 19 . 19a . 19b ), characterized in that - the illumination and measuring beam path ( 7 ) - a beam splitter ( 11 ) for separating the from the eye ( 3 ) collected measuring radiation (M) from the eye ( 3 ) guided illumination radiation (B), wherein the beam splitter ( 11 ) the divided measuring radiation (M) to the detection beam path ( 14 . 15 . 17 ), and - an optical element acting only on the illumination radiation (B) ( 10.1 . 10.2 ), which with the front optics ( 12 ) and the passage of the illumination radiation (B) through the pupil (P) of the eye ( 3 ) only in a first pupil area ( 72 ), and - the detection beam path continues to act only on the measuring radiation (M) optical element ( 14.1 . 14.2 ), which with the front optics ( 12 ) and the collecting of the measuring radiation (M) only by a second pupil area ( 70 ) of the pupil (P) of the eye ( 3 ), where - first and second pupil area ( 72 . 70 ) do not overlap and pass through a neutral area ( 71 ) of the pupil (P) of the eye ( 3 ) are separated from each other, neither to the first nor to the second pupil area ( 72 . 70 ) belongs. Optical coherence tomograph according to claim 1, characterized in that the at least one detector ( 19 . 19a . 19b ) is in or near a plane conjugate to a plane in which the pupil (P) of the eye ( 3 ) lies. Optical coherence tomograph according to one of the preceding claims, characterized in that the beam splitter ( 11 ) is a polarization beam splitter and that between the eye ( 3 ) and the beam splitter ( 11 ) a lambda / 4 plate ( 27 ) is arranged, which filters the measuring radiation with respect to a polarization state, which is adapted to a polarization state of the reference radiation (R). Optical coherence tomograph according to one of the above claims, characterized in that the detection beam path ( 14 . 15 . 17 ) for balanced detection, a beam splitter / combiner ( 18 ), which superimposes the reference radiation (R) with the measuring radiation (M) in two different phase positions on two detectors ( 19a . 19b ). Optical coherence tomograph according to one of the preceding claims, characterized in that optical elements, in particular lenses, are arranged in a zone around a conjugated pupil plane, which is at a distance of z = m · (d · L) / (d + x) is defined at the conjugate pupil plane and y = m * (d * L) / x behind the conjugate pupil plane, with "before" and "behind" in the imaging direction, m is an imaging factor of the collection of the measurement radiation (M), L a length of the eye ( 3 ), d is a width of the neutral region, x is an extension of an eye ( 3 ) pictured area ( 2 ). Optical coherence tomograph according to one of the preceding claims, characterized in that the detection beam path is an aperture ( 15 ) provided to the at least one detector ( 19 . 19a . 19b ) is arranged in or near an intermediate image plane and determines the size of an object field from which the measurement radiation (M) to the detector ( 19 . 19a . 19b ), wherein preferably the at least one detector is an area detector ( 19 . 19a . 19b ) having a spatial resolution of 4 to 100 pixels in one direction, preferably 2D area detector having 5 to 50 pixels or 5 to 40 pixels per direction. Optical coherence tomograph according to one of the above claims, characterized in that illumination and measuring beam path ( 7 ) a scanner ( 13 ) for adjusting the lateral position of the illumination field in the eye ( 3 ) having. Optical coherence tomograph according to claims 6 and 7, characterized by a control device (C) which controls the scanner ( 13 ) for deflection during wavelength tuning and generates or receives a scan signal indicative of a deflection condition of the scanner ( 13 ) and with the radiation source ( 4 ) for reading out a wavelength signal which indicates the wavelength of the source radiation and thus of the illumination radiation (B), and the at least one area detector ( 19 . 19a . 19b ) is connected to read out measurement signals for each pixel, wherein the control unit from the wavelength signal and the measurement signals fields ( 59 ) of the retina ( 2 ) and evaluates the scan signal to the fields ( 59 ) to a 3D overall picture ( 61 ). Method for optical coherence tomography for examining an eye ( 3 ), the method comprising - providing source radiation, tuning 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 in the eye ( 3 ) backscattered illumination radiation as measuring radiation (M), wherein a front optics ( 12 ) and the pupil (P) of the eye ( 2 ) is a pupil of the illumination and the collecting of the measuring radiation (M), - separating the from the eye ( 3 ) collected measuring radiation (M) from the eye ( 3 ) guided illumination radiation (B), - defining a first pupil area ( 72 ) of the pupil (P) of the eye ( 3 ), by the illumination radiation (B) on the eye ( 3 ) by using an optical element acting only on the illumination radiation (B) ( 10.1 . 10.2 ), which with the front optics ( 12 ) and defining a second pupil area ( 70 ), by the eye ( 3 ) Measuring radiation (M) is collected by using an optical element acting only on the measuring radiation (M) ( 14.1 . 14.2 ), which with the front optics ( 12 ), wherein first and second pupil area ( 72 . 70 ) do not overlap and pass through a neutral area ( 71 ) of the pupil (P) of the eye ( 3 ) are separated from each other, neither to the first nor to the second pupil area ( 72 . 70 ), and - overlaying the measurement radiation (M) with the reference radiation (R) and detecting an interference signal of the superimposed radiation with at least one detector ( 19 . 19a . 19b ). Method according to claim 9, characterized in that the at least one detector ( 19 . 19a . 19b ) is placed in or near a plane conjugate to a plane in which the pupil (P) of the eye ( 3 ) lies. Method according to one of claims 9 or 10, characterized in that the separation of the eye ( 3 ) collected measuring radiation (M) from the eye ( 3 ) guided illumination radiation (B) by means of a polarization division, wherein previously the measurement radiation is filtered with respect to a polarization state, which is adapted to a polarization state of the reference radiation (R) in the superposition, and not portions of the measurement radiation (M) are discarded corresponding to this polarization state , Method according to one of claims 9 to 11, characterized in that the detection of the interference signal by means of a balanced detection takes place. Method according to one of claims 9 to 12, characterized in that optical elements, in particular lenses, are arranged in a zone around a conjugated pupil plane, which is separated by a distance of z = m · (d · L) / (d + x) is defined before the conjugate pupil plane and y = m * (d * L) / x behind the conjugate pupil plane, where "before" and "behind" is related to the imaging direction, m is an imaging factor of the collection of the measuring radiation (M), L is a length of the eye ( 3 ), d is a width of the neutral region, x is an extension of an eye ( 3 ) pictured area ( 2 ). Method according to one of claims 9 to 13, characterized in that the at least one detector is an area detector ( 19 . 19a . 19b ) with a spatial resolution of 4 to 100 pixels in one direction, preferably a 2D area detector with 5 to 50 pixels or 5 to 40 pixels, and that by means of a diaphragm ( 15 ) associated with the at least one area detector ( 19 . 19a . 19b ) is arranged upstream and in an intermediate image plane is arranged, the size of an object field is determined, from which the measuring radiation (M) to the area detector ( 19 . 19a . 19b ). Method according to one of claims 9 to 14, characterized in that in the illumination and measuring beam path ( 7 ) by means of a scanner ( 13 ) the lateral position of the illumination field in the eye ( 3 ) is adjusted. Method according to claims 14 and 15, characterized in that the scanner ( 13 ) for deflection during the wavelength tuning and from measurement signals of the area detector ( 19 . 19a . 19b ) and a wavelength signal fields ( 59 ) of the retina ( 2 ) and taking into account the deflection condition of the scanner ( 13 ) the partial images ( 59 ) to a 3D overall picture ( 61 ).

Images