fig1 depicts an oct 1 , which captures three - dimensional images of a retina 2 of an eye 3 . source radiation of a radiation source 4 which can be tuned with respect to its wavelength , for example a suitable laser , is coupled into a fibre 5 . the source radiation is , for example , in the infrared wavelength range . in the following description , this wavelength range is also referred to as “ light ”. under this term , all radiation of the electromagnetic spectrum which satisfies the optical laws is subsumed . the fibre 5 terminates into a splitter 6 , which splits the source radiation into a measurement arm 7 and a reference arm 8 . at the splitter 6 , in the measurement arm 7 , a fibre 9 is attached , and the illumination radiation b emerging at the end of the fibre is conducted to a beam splitter 11 by an illumination optical system 10 . from there , it reaches a front optical system 12 , which bundles the illumination radiation b in a focus , which lies on the retina 2 of the eye 3 . the illumination optical system 10 and the front optical system 12 thus set , among other things , the numerical aperture na with which the eye 3 is illuminated . between the beam splitter 11 and the front optical system 12 there is located a scanner 13 , which shifts the focus over the retina 2 biaxial and perpendicular to the direction of incidence , i . e . lateral . the coordinates of such deflection are labelled x and y in the following . the z position of the focus can be set by adjusting the front optical system 12 . 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 field depends from the numerical aperture na , which is defined by the combination of front optical system 12 and illumination optical system 10 as well as the optical properties of the eye 3 . the backscattered radiation is collected by the front optical system 12 as measurement radiation m . for differentiation of the incident illumination radiation and the backscattered measurement radiation m collected by the front optical system 12 , these are plotted differently in fig1 . the illumination radiation is drawn in the figure with continuous lines ; the measurement radiation m with dotted lines . the measurement radiation collected by the front optical system 12 is guided to the scanner 13 . here , it is descanned with the result that , following the scanner 13 , the measurement radiation m is present as a resting beam . the collection of the measurement radiation m is an imaging of the retina 2 . the beam splitter 11 separates the measurement radiation m from the illumination radiation b and guides it to detector optical system 14 . the detector optical system 14 defines , together with the front optical system 12 and the optical properties of the eye 3 as well as any further imaging elements in the imaging beam path ( e . g . a lens 16 ), the numerical aperture na of the imaging of the retina 2 . in this way , illumination and detection have different numerical apertures . the numerical aperture of the illumination is defined by the combination of the illumination optical system 10 and the front optical system 12 . the numerical aperture of the detection is defined by the detector optical system 14 and the front optical system 12 . the detector optical system 14 focuses the measurement radiation m to an intermediate image plane , in which a diaphragm 15 is located . this diaphragm 15 defines the size of the object field in which measurement radiation m is detected from the retina 2 . taking into consideration the magnification factor of detector optical system 14 , front optical system 12 and eye 3 , the size of the diaphragm 15 corresponds exactly to the size of the object field on the retina 2 , from which measurement radiation m is collected . a further optical system 16 following the diaphragm 15 directs the measurement radiation m onto a detector device 17 . in the embodiment of fig1 , the detector device 17 comprises a beam splitter / combiner 18 as well as two 2d sensors 19 a and 19 b . the size of the 2d sensors 19 a , 19 b is designed to match the diaphragm 15 and the intermediary optical system 16 . they have a spatial resolution , i . e . they resolve an intensity distribution over beam cross section . the detector device 17 for example lies in a pupil plane of the measurement beam path , i . e . in a plane which is conjugate to the plane of the pupil p of the eye 3 , through which the measurement radiation m emerges . at the beam splitter / combiner 18 , reference radiation r from the reference arm 8 is also coupled in . following the splitter 6 , the reference arm comprises a fibre 20 . in the embodiment shown in fig1 , the reference arm 8 has a path length adjusting device 21 , which serves to set 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 fibre 20 and conducted via a retroreflector 22 , the position of which can be adjusted , as the double arrow in fig1 indicates . via a further deflecting mirror 23 as well as optical systems 24 , 25 , the reference radiation r is conducted to the beam splitter / combiner 18 , which conducts the reference radiation r with the measurement radiation m in a superposed manner onto the area sensors 19 a and 19 b . in fig1 , the path length adjusting device 21 is designed as a free beam path . this is optional , as is the use of a retroreflector 22 . in the state of the art , various measures are known for adjusting the optical length of a beam path . in fig1 , the detector device 17 is designed for so - called “ balanced detection ”. this is also optional , as is further explained in the following with reference to fig2 . balanced detection has the advantage that a common mode portion in the superposition 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 19 b , was used and the beam splitter / combiner 18 is designed as a simple beam combiner . the signal of interference between reference radiation r and measurement radiation m is converted to an image , as is known for optical coherence tomography . since the wavelength of the source radiation is tuned , the fourier domain principle of image generation is utilized , which is generally known from the state of the art . for image generation , the oct 1 comprises a control device c , which receives a wavelength tuning signal and the measurement signals of the 2d detectors 19 a , 19 b . optionally , the control device c controls the radiation source 4 for wavelength tuning , therefore knows the wavelength currently prevailing in the system and can thus assign the measurement signals accordingly . the 2d detectors receive measurement radiation m from an object field on the retina 2 , which is defined by the diaphragm 15 . the size of the 2d detectors 19 a , 19 b is adapted to correspond to the diaphragm 15 and they sense the intensity distribution in a spatially resolved manner with their individual pixels . if the 2d detectors 19 a , 19 b lie in an image plane , i . e . in a plane which is conjugate to the plane of the retina in the imaging process carried out by front optical system 12 , detector optical system 14 , and further intermediary optical elements , the individual pixels already contain the location information from the object field . if , on the other hand , the 2d detectors lie in a conjugate pupil plane , which is conjugate to the plane in which the pupil p of the eye 3 lies , the detectors detect the intensity distribution in the pupil plane and thus the phase information . this can also be used for image reconstruction , as will be explained later . it is important for the invention that the scanner 13 shifts the object field over the retina 2 since it acts not only on the illumination radiation b but also on the collection of the measurement radiation m . at each position of the scanner 13 , a single image of the retina forms , the resolution of which is determined by the number of pixels and the arrangement of the 2d detectors 19 a , 19 b . as will be explained in the following , these single images are assembled to form a total image , which has a considerably higher resolution than known from standard widefield oct . in the design of fig1 , the beam splitter / combiner 18 combines the measurement radiation m from the measurement arm 7 and the reference radiation r from the reference arm 8 . the 2d detectors 19 a , 19 b detect the pattern of the interference between measurement radiation m and reference radiation r . the requirements necessary to generate such interference , in particular the properties of the radiation source 4 and of the path length adaption , are known in the state of the art of optical coherence tomography . because of the balanced detection , there is a relative phase difference of pi between the sum of the two signals of the two 2d sensors 19 a , 19 b . the complex amplitudes of the measurement radiation and of the reference radiation can be written as : if the amplitudes are labelled with u s and u r and the phases of the signals in the two arms are labelled with φ s and φ r ( the indices “ sample ” and “ s ” relate to the measurement arm , the indices “ reference ” and “ r ” relate to the reference arm ). the signals i 1 and i 2 detected by the two sensors 19 a , 19 b are then : i 2 =| u sample + u reference | 2 =| u sample 2 +| u reference | 2 + 2 re { u sample * u reference } and i 2 =| u sample + u reference * e iπ | 2 =| u sample | 2 +| u reference | 2 + 2 re { u sample * u reference * e − iπ . ū is complexly conjugate to u , and re is an operator , which supplies the real part of a complex value . as differential signal i bd of the two detectors 19 a , 19 b , the following is obtained : wherein δφ := φ s − φ r refers to the relative phase between measurement and reference arm . the formulae show that only the interference pattern of the two signals cos ( δφ ) is still present in the differential signal of the two detectors 19 a , 19 b and the common mode components | u sample | 2 and | u reference | 2 are suppressed . in this way , in particular in the case of any subsequent analog - digital conversion of the differential signal , the dynamic range of the signal is used to the maximum for information extraction . another embodiment works without balanced detection ; the signal amplitude of the interference signal is then modulated onto a common mode portion and is filtered out by suitable data analysis . fig2 shows a modified design of the oct 1 , which resembles that of fig1 in many respects . the same elements bear the same reference numbers as in fig1 . the essential difference consists in the design of the detector device 17 , which , in the design of fig2 , comprises only one single 2d detector 19 . the measurement radiation m and the reference radiation r strike this 2d detector 19 at an angle and interfere with each other . through the angular offset , a phase shift occurs between pixels which lie in a plane which is spanned by the optical axis , along which the measurement 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 portion . such a detection arrangement is referred to as off - axis detection and is known for common mode suppression to a person skilled in the art . fig3 shows a further embodiment for the oct 1 , wherein here the path length adjusting device is not arranged in the reference arm 8 , but in the measurement arm 7 . following the fibre 9 and the illumination optical system 10 there is a path length adjusting device 29 , again purely by way of example in form of a movable retroreflector 30 . the embodiment of fig3 shows that it does not matter whether the path length adjusting device lies in the reference arm 8 or in the measurement arm 7 . it is also possible to provide a path length adjusting device in both of them . it is only necessary that the interference state between the reference radiation r from the reference arm 8 and the measurement radiation m can be set such that it is adapted to the current measuring task , i . e . the actual position of the object to be measured , in the embodiment examples described here by way of example the retina 2 of the eye 3 . in fig3 , two further features are represented which can be used individually for all embodiments of the oct 1 . the front optical system 12 is formed in two parts by two imaging elements 12 a and 12 b . the scanner 13 is for example located , in the case of the oct 1 of fig1 to 3 , in or near to a pupil plane of the detection beam path as well as also of the illumination beam path . this pupil plane is conjugate to the plane of the pupil p of the eye 3 . the front optical system 12 comprises the optical sub - systems 12 a and 12 b , which together form a 4f optical system . thus , the optical sub - system 12 a is an ophthalmoscopic lens and the optical sub - system 12 b is a scan lens . this 4f optical system images the pupil p of the eye 3 to a pupil plane conjugate to the plane of the pupil p , in which pupil plane the scanner 13 lies . the scanner 13 does not have to be placed exactly in this conjugate pupil plane but that does have advantages . between the plane of the pupil p of the eye 3 and the pupil plane conjugate thereto there is an intermediate image plane 26 . because of its proximity to the scanner 13 , the beam splitter 11 is also located near to the conjugate pupil plane . it is also possible to place the beam splitter 11 in this conjugate pupil plane if the scanner 13 is moved out of the conjugate pupil plane . in one embodiment , the beam splitter 11 is formed as a polarizing splitter . then , a lambda / 4 plate 27 is arranged in front of it when seen in imaging direction . this embodiment will be discussed later . the detector optical system is likewise formed as a 4 f optical system . it provides a further intermediate image plane 28 , in which lies the diaphragm 15 ( not drawn in in fig3 ). the intermediate image plane 28 is conjugate to the object plane , in which lies the retina 2 to be imaged . the size of the diaphragm 15 ( not shown ) determines the size of the imaged area on the retina 2 , of course under consideration of the magnification factor which is relevant to the generation of the intermediate image plane 28 . in all embodiments , the diaphragm 15 has two functions . on the one hand , it suppresses scattered light , whereby the contrast is improved at the detector device 17 . ultimately , in this respect the diaphragm acts in a similar way to a confocal diaphragm of confocal scanning octs . because of the effect of the detector optical system , the detector device 17 for example lies in a plane which is conjugate to the pupil plane of the eye , or near to this plane . this arrangement is advantageous but not mandatory . it has the advantage that the phase function of the electromagnetic field can be sensed easily . the maximum spatial frequency in the plane of the 2d detector 19 or of the 2d detectors 19 a , 19 b is predetermined by the size of the object field on the retina 2 and thus , ultimately , the size of the diaphragm 15 in the intermediate image plane 28 . the diaphragm 15 thus , on the other hand , facilitates a particularly favourable signal detection . in all embodiments of the oct , the number of pixels of the 2d detector is 4 to 100 , for example 5 to 50 , in another example 5 to 40 pixels in each direction . fig5 shows a top view 43 of the detector in that the arrangement of the pixels does not necessarily have to be rectangular but that a hexagonal arrangement of the pixels is also an option . the pixel pattern can be chosen freely , thus . in the state of the art , holoscopic oct systems are known which comprise detectors with 100 to 4000 pixels per direction . these numbers of pixels are deliberately not used here . the number of pixels is linked to the required illumination brightness , the measurement rate and the suppression of multiple scatterings . in an example embodiment of the oct 1 , aberrations are corrected . the detector device 19 comprises , as already mentioned , one or two 2d detectors which have a spatial resolution in the form of pixels . these pixels are also referred to as channels in the following . the measurement signal is distributed over these several channels of the detector ( s ). if , in an example embodiment , the detector lies in a conjugate pupil plane , each channel of the detector receives measurement radiation m from different angles , which radiation was scattered from the retina 2 . the spatial resolution of the 2d detector 19 , 19 a , 19 b allows it to detect the distribution of the measurement radiation in the pupil p . aberrations have an effect on this distribution . aberrations caused by the eye 3 often assume a level which is no longer acceptable when an area in the plane of the pupil p of the eye 3 is utilized , which is larger than 1 . 5 mm in diameter . however , such a larger area would be desirable with respect to the lateral resolution . without spatial resolution in the conjugate pupil plane , a larger pupil utilization in the eye 3 would mix and average phase differences in the then single detection channel . the corresponding zernike polynomials which describe these aberrations are represented in fig4 , which shows top views 37 to 42 of a conjugate pupil plane . further , the grid of a detector with 5 × 5 channels ( or pixels ) is shown . the pixels sense the pupil p and thus allow phase differences within the pupil p to be distinguished . the maximum resolvable phase differences depend on the number of channels . it was found that the number of distinguishable phase differences in this plane results from the number of channels per direction multiplied by pi . in the case of five channels per direction , as is represented in fig4 , polynomials to z ″ 4 can be distinguished , wherein m can assume the values 0 ( sphere ), ± 2 and ± 4 . this applies to channels with an infinitesimally small surface area . in reality , of course , they have a particular size . the measurement signal detected in a channel therefore corresponds to an averaging of the interference signal over the surface area of the respective channel ( pixel area ). the maximum theoretically possible order of the zernike polynomial can thus only be achieved when the phase of the signal within a channel varies by less than pi . it was found that , at an average wavelength of the oct of 1060 nm , with uniformly spatially distributed channels , the phase differences of astigmatism caused by the eye can be detected , if , in the case of five channels , the condition 2 pi /( 5 channels per aberration period )≦ pi is met . then , a period of minima and maxima lies within the aperture . for higher orders , the following applies : 0 . 6 * 2 pi /( 5 channels per period of the aberration )= 1 . 2 * pi /( 5 / 1 . 5 )≦ pi for the third order and 0 . 5 * 2 pi /( 5 channels per period of the aberration )= 1 . 0 * pi /( 5 / 2 )≦ pi for the fourth order . these findings show that a 2d detector with at least five channels per direction is capable of resolving at least astigmatism and the aberrations of third order . a higher number of channels allows even higher orders of the aberration to be detected . the above analysis considered only one spatial direction . as fig4 shows , the aberrations generally have a two - dimensional pattern . in the top view 37 , fig4 shows the aberration of first order , which is also referred to as “ piston ”, the top view 38 shows the aberration “ tilt ”, the top view 39 shows the aberration “ tip ”, top view 41 shows the aberration “ defocus ” and the top views 40 and 42 show aberrations of the “ astigmatism ” type . as can be seen , most aberrations have a two - dimensionally distributed pattern , whereby the phase variation is also two - dimensional . this pattern can be detected and corrected by the spatially resolving 2d detector 19 , 19 a , 19 b . for each detector channel c , the aberrations produce a phase θ c : u sample , c = u sample * e iθ c . it results from a thickness δd and a refractive index δn of the passed - through material of the eye ( e . g . cornea , aqueous humour , lens , vitreous body ), which in reality differs from a theoretical , aberration - free eye : for monochromatic radiation of 780 nm , the eye causes wavefront aberrations of up to 0 . 7 nm , which lead to a phase shift of 2 * pi ( if defocus is disregarded ). such a phase shift corresponds to a deviation in thickness between lens and aqueous humour ( these are the elements with the greatest differences in refractive index in the eye ), which assumes the following value : if a wavelength range of δλ = 50 nm is covered , the phase differences of the associated wave numbers ( k 0 ± δk ) are : these calculations show that , in sufficiently exact approximation , the phase shifts which are caused by the aberrations vary linearly with the wave number k within a wavelength sweep . the detected measurement signal can thus be expressed as follows : i bd , c ( k )= 4 * u s * u r * cos ( k *( δ z − δn ( k 0 ) δ d c )). a fourier transformation for the measured wave numbers k yield the axial distribution , i . e . the distribution of the scattering tissue in z direction . compared with an aberration - free system , the axial distribution is shifted by the value δn ( k 0 ) δd c for each channel c of the 2d detector . fig8 shows a corresponding simulation example , in which the z coordinate is plotted on the x axis and the signal intensity is plotted on the y axis . the curves 51 to 54 correspond to four channels c of the 2d detector . it can be assumed that , in most areas of the tissue , the variation of the axial scatter profile is small within a pupil size of 5 mm of the eye 3 . the differences in profile for the channels c are therefore caused mainly by the aberrations , which shift the profile axially . it is therefore provided to relate the aberration - caused phases θ c ( k 0 ) of the channels to a central channel ( for example the channel lying in the center of the detector , which channel corresponds to a perpendicular incidence on the sample ). the measured intensities for a frequency determination are multiplied by the phase factor to correct the aberrations . the phase factor is e − iθ c ( k 0 ). each channel of the detector has a particular position in relation to the retina 2 . the interference signal can be captured during the wavelength adjustment of the laser for the respective wave number k = 2 * pi * n / λ , wherein n is the refractive index of the medium and λ is the wavelength . as in conventional oct systems , the measurement signals are fourier - transformed with respect to the wave numbers , and the depth distribution of the scattering layers is calculated . here , the relationship δφ = k * δz is used , wherein δz is the distance of one scattering layer to a layer , from which the measurement radiation passed through a certain path length to the detector , which certain path length is identical to the path length of the reference radiation . because of the lateral extension of the 2d detector 19 , however , the optical path length for the individual pixels is not identical , as fig6 shows . the five pixels or channels 46 , 47 , 48 , 49 and 50 drawn in by way of example for the 2d detector 19 . they differ with respect to the optical path length to a particular point in the tissue 44 . the wavefronts for the central channel 48 are drawn in with continuous lines . they are perpendicular to the optical axis of the beam path between the channel 48 and the observed point , which is drawn in at the bottom of the structure 44 . for the central channel 48 , the radiation runs along the optical axis . for a channel situated further out , for example the channel 50 , the chief ray runs at an angle α to the optical axis such that the path length , which has the value d for the central channel 48 , takes the value d * cos ( α ) for the outer channel 50 . the corresponding wavefronts and chief rays for the outer channel 50 are drawn in with dashed lines in fig6 . further , in fig6 the lens of the eye 45 is shown by way of example . the depth is based on the principal plane of the lens of the eye since the refractive index step thereof can be used as a reference point during the measurement . as fig6 clearly shows , pixels / channels which lie further out collect radiation which has covered a longer path through the medium . in a reconstruction of the image information , this has the effect shown in fig7 by way of example . there , signal curves 51 to 54 are shown for four channels . the plot corresponds to that of fig8 , i . e . the depth coordinate z is plotted on the x axis , the intensity on the y axis . as can be seen , the individual curves are not only shifted in the z direction , they are also concentrated for pixels lying further out . the curve 54 is the measurement signal of the central pixel 48 , and the curves 53 , 52 and 51 are measurement signals from channels lying further out in each case . the measurement error caused by this effect is corrected in an example embodiment , in order to obtain a particularly good image capture . the geometric effect is for example corrected by carrying out a rescaling from z to z cos ( α c ), wherein as is the angle which the c th channel has to the optical axis . the angle α is based on a virtual position of the 2d detector 19 in which the detector is placed directly in front of the eye and by taking into consideration the magnification factor . a 2d detector which lies exactly in a plane conjugate to the pupil plane of the eye virtually extends in the plane of the pupil p of the eye 3 at a size which is modified by the magnification factor . in the reconstruction with respect to the aberration , different channels are reconstructed independently . subsequently , the cross - correlation is formed in axial direction , i . e . in depth direction , in order to determine the relative phase offsets between the individual channels . a reconstruction of the lateral image for each channel ( optionally , as will be described in the following , taking into consideration the scanning procedure ) and then of the phase gradient yields a lateral offset in the image which is obtained for a given position of the scanner . this image is also referred to in the following as a pupil channel partial image . by a lateral cross - correlation of the pupil channel partial image an embodiment determines the aberration and , in this way , determines and numerically corrects the entire aberration phase distribution . the quality of these approaches depends on the sample structure . in case of the human eye , an easily recognizable axial layer structure is available . lateral thereto , the structures are relatively rough , for example on account of blood vessels or the papilla , combined with very fine structures , such as photoreceptors , wherein few structures lie in between in terms of size and roughness . an example embodiment , therefore , first of all , carries out a depth correlation correction by using the axial layer structure in order to correct the greatest proportion of the pupil phase aberrations . optionally , a lateral correlation correction follows , which utilizes lateral structures , such as for example photoreceptors , which became visible because of the first correction . the aberrations of the eye are different at different points of the retina . in principle , it is possible to calculate the aberration - caused phase changes in each channel for all points in a lateral image . in a simplified embodiment , it is assumed that the aberrations do not vary very greatly laterally , and the aberrations are calculated only for a few lateral locations of the retina and are interpolated for locations in between . if a relatively large wavelength range is covered , it is preferred to take into consideration the dispersion of the aberrations . in this embodiment , it is not assumed that the phase shifts change linearly with the wave number k . a peak in the profiles , which originates at the fundus of the eye 3 in the oct image of the retina 2 , is therefore used in order to balance the shift of the profiles with respect to each other . for example therefore , a structure ( in the form of a peak ) is sought , in the curves 51 to 54 of fig7 , and the curves are corrected relative to each other on the basis of this reference structure . 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 expansion ) of the measurement signals can also be corrected . in one position of the scanner 13 , a single image of the retina is obtained , the size of which is predetermined by the diaphragm 15 and the front optical system 12 and the detector optical system 14 co - operating during the imaging of the measurement light . a fourier transformation of the signal of the channels provides the image of the sample , but only in a part which corresponds to the size of the detector in the pupil . in order to generate a larger image , the scanner 13 is provided , which shifts the position of the imaged object field , i . e . the image section on the retina 2 . the image section corresponds to a single image 59 , which has a centre 60 . for the current deflection by the scanner 13 , it is sufficient for simplification to refer to the centre 60 of the single image 59 . various scanning approaches are possible now . for instance , the center 60 of the single image 59 can be left unchanged during the sweep of the wavelength of the light source 4 . before a new sweep is started , the center 60 is shifted such that a new single image 59 borders directly on the previously captured single image 59 . in this way , a larger total image 61 of the retina can be detected . this approach is shown for a depth plane in fig9 . as a result , individual single images 59 are assembled to form the total image 61 . the images from the individual planes then result in a three - dimensional image of a cuboid zone in the retina 2 . fig1 shows this , in which three planes 62 , 63 and 64 can be seen by way of example . the single images 59 , which are referenced to each other in the representation of fig1 with a dot - dashed double arrow , in each case originate from one wavelength sweep of the light source 4 . since the scanner 13 rests during each wavelength sweep and is only adjusted between them , the single images 59 generated from one wavelength sweep lie in the planes 62 to 64 all with their centres 60 exactly one above the other . for particular embodiments of the scanner 13 it is preferred to operate the scanner continuously , i . e . to adjust while the wavelength is adjusted in a sweep . this approach requires synchronization of the scanner 13 and wavelength adjustment at the light source 4 . here , it is preferred to set the lateral adjustment speed through the scanner 13 such that , during one wavelength sweep , at most one single image 59 is passed in one direction , preferably even fewer . thereby , for the individual planes 62 , 63 and 64 , the position of the centers 60 changes since the single images 59 in the planes come from different wavelengths of the fourier transformation . as a result , a temporary total image 61 is obtained , which , unlike in the embodiment of fig1 , is not a right - angled cuboid , but is e . g . a no longer right - angled parallelepiped , depending on the adjustment of the scanner 13 during the wavelength sweep . for the imaging , this effect is for example corrected in that the parallelepiped is cropped to form a cuboid . there are various possibilities for taking into consideration the synchronicity of wavelength sweep and lateral displacement . if the detector lies close to an intermediate image plane , i . e . in a plane conjugate to the retina , the data of the three - dimensional parallelepiped are shifted with respect to each other . for each wave number k 1 , an image of the sample can be added , wherein i i = i ( k i , x , y ). these images i i are shifted somewhat with respect to each other . since the assignment between lateral scan position and wave number is known , the entire wavelength adjustment can be assembled correspondingly for each location ( x , y ) in the sample . in this way , the three - dimensional data are assembled simply . in example embodiments in which the detector is located in or near the conjugate pupil plane , it measures the fourier transformation of the intensity distribution in the object plane ( retina 2 ). a displacement in the object plane leads to a phase ramp in the detector plane . the correction of the simultaneous lateral adjustment by the scanner 13 and of the wavelength adjustment by the light source 4 is therefore a multiplication of the detector signal by a time - dependent phase ramp , which is proportional to the scan speed and the distance between pupil partial channel and optical axis in the pupil plane . the optical set - up of fig1 to 3 de - couples the illumination and the capture of the measurement light from each other with respect to the optical properties and , in particular , the pupil size . in this way it is possible to optimize the illumination . for example , a bessel - type illumination can be combined with a top hat cross - sectional profile for detection . in this way , in one embodiment achieves a high illumination depth , i . e . an illumination which is unchanged over a large z - range , and a high numerical aperture of the imaging at the same time . in the case of identical numerical aperture , an illumination focus of an extent of 1 mm would be achieved in z direction , for example with a gaussian beam . in the case of a bessel - type illumination , 2 to 3 mm extent is obtained in z direction . in this way , the optical resolution can be increased by 10 to 30 % when the detection is done with a top hat - type profile . a further example embodiment of the oct uses a polarization splitting at the beam splitter 11 . such a thing is usually disadvantageous in the state of the art , and an intensity splitting is generally used . polarization splitting is , surprisingly , advantageous for the described oct , since polarized radiation entering the eye is changed with respect to its polarization state . different structures of the eye have a different effect such that the polarization state of the backscattered signal is not unambiguously or clearly defined , but consists of components of different polarization state . this was also known in the state of the art and resulted in an intensity splitting being carried out , precisely because the backscattered radiation does not have a clearly defined polarization state . however , it was found that the measurement light is superimposed with the reference light and that only the components of the beams can interfere with each other , which have the same polarization state . ultimately , the reference light therefore determines with its polarization state which portion of the measurement light can be utilized . non - interfering portions fall on the detector and form an interfering background . after the polarizing splitter 11 , the illumination radiation b is linearly polarized . the lambda / 4 plate 27 , as is shown in fig3 , provides circularly polarized illumination radiation b at the eye 3 . backscattered measurement radiation m , which is likewise circularly polarized , is again linearly polarized by the lambda / 4 plate , with the polarization direction being rotated by 90 degrees compared with the polarization direction of the illumination radiation b , which was emitted by the polarizing splitter 11 . the measurement radiation m then passes through the polarizing splitter 11 without deflection and interferes with the reference radiation r , if this has the same polarization . this is the case when reference radiation r and illumination radiation b are identically linearly polarized after dividing from the source radiation . this is also the case when reference radiation r and illumination radiation b are circularly polarized after dividing from the source radiation and the reference radiation is linearly polarized identically to the measurement radiation m before superposition . generally , it is of importance that the polarization splitting ( e . g . by polarizing splitter 11 and plate 27 ) conditions the measurement radiation m and the reference beam path conditions the reference radiation r such that both radiations have the same polarization state at the detector . this measure increases the signal - to - noise ratio , since only those parts of the measurement light are guided by the beam splitter 11 to the detector device 17 which are capable of interfering with the reference light . thus , the polarization splitting which is actually disadvantageous and the discarding of a part of the measurement radiation m at the beam splitter 11 increases the quality of the signal . in a further embodiment of the oct , use is made of the fact that the illumination optical system 10 allows placement of the focus of the illumination radiation b at z position different from that of the focus which is predetermined by the detector optical system 14 for the collection of the measurement radiation m . because of multiple scatterings in the retina , measurement radiation m from the retina can have the path length suitable for interference , but can propagate in another direction , which would limit the lateral resolution at depth . this effect can be compensated through different depth planes for illumination ; and detection and the resolution at depth is optimized . for image reconstruction from the detector signals according to the fd - oct principle , it is necessary to know the current wavelength . this wavelength or the corresponding wave number k can be derived from controlling the light source 4 . alternatively , it is possible to couple out a beam portion and to record it in terms of the wavelength in order to know the current wavelength or the course of a wavelength sweep better . perpendicular to the scanning direction , detector channels can be binned in order to reduce speckle . this is particularly advantageous if only z sections through the retina are desired . for a roughly resolved image , e . g . as preview , it is possible to add all or several detector channels . this is done after the corrections ( e . g . aberration , z - position , total image generation ) were made . the resolution of known oct systems is obtained but with a higher signal - to - noise ratio and improved speckle behaviour , precisely because the combining is done after one or more of the corrections and goes beyond a normal pixel binning , thus . if a detector is used which spatially resolves in only one direction , aberrations can also be corrected only in this direction . for particular applications this may be sufficient . in one example embodiment a multilens array is arranged in front of the 2d detector 19 or in front of each of the 2d detectors 19 a , 19 b to improve a filling factor of an illumination of the pixels . in one example embodiment , an iris camera is provided , which supports the user during the setting of the device to the eye position . so far as method steps and / or signal corrections were described above , these are carried out in the oct 1 by the control device c , which is connected to the detector / the detectors . the device reads the measurement signals of the detector / detectors and receives further data about the operation of the scanner 13 and the wavelength sweep and / or actuates such components correspondingly .
an optical coherence tomograph includes a wavelength tunable illuminating device , an illumination and measurement beam path with a dividing element and a scanner and a front optical unit and a reference beam path , a detection beam path and a flat panel detector . a beam splitter conducts the separated measurement radiation to the detection beam path and an optical element acts only on the illumination radiation . the optical element sets the numerical aperture of the illumination of the illumination field in the eye . an optical element acts only on the measurement radiation and sets the numerical aperture with which measurement radiation is collected in the eye . an aperture is arranged in front of the flat panel detector in an intermediate image plane and defines the size of an object field . the flat panel detector has a spatial resolution of 4 to 100 pixels in a direction .