APPARATUS AND METHOD FOR MEASURING AN OPTICAL BREAK-THROUGH IN A TISSUE

Abstract

The invention relates to a device for measuring an optical penetration that is triggered in a tissue underneath the tissue surface by means of therapeutic laser radiation which a laser-surgical device concentrates in a treatment focus located in said tissue. The inventive device is provided with a detection beam path comprising a lens system which couples radiation emanating from the tissue underneath the tissue surface into the detection beam path. A detector device generating a detection signal which indicates the spatial dimension and/or position of the optical penetration in the tissue is arranged downstream of the detection beam path.

Claims

1.-31. (canceled)

32. A device for measuring an optical break-through which is created in a tissue, beneath a tissue surface, by treating laser radiation which a laser surgical unit focuses into a treatment focus, said focus being located in the tissue wherein said device comprises a detection beam path comprising optics, wherein the optics couple radiation emitted by the tissue from beneath the tissue surface, into the detection beam path, and a detector unit is arranged following the detection beam path, said detector unit generating a detection signal which indicates the spatial extent, position or both of the optical break-through in the tissue.

33. The device as claimed in claim 32, further comprising an illumination radiation source, which directs illumination radiation into the tissue.

34. The device as claimed in claim 33, wherein the illumination radiation source supplies the treating laser radiation.

35. The device as claimed in claim 33, wherein the illumination radiation source and the detection beam path are part of an interferometer structure.

36. The device as claimed in claim 35, wherein the interferometer structure comprises a measuring arm and an adjustable reference arm and the illumination radiation has a coherence length, in the direction of light propagation and in which the resolution at which the detection signal indicates the spatial extent depends on the coherence length, and wherein interference appears only, if the lengths of the measuring arm and of the reference arm differ by no more than the coherence length.

37. The device as claimed in claim 35, wherein the illumination source radiation focuses the illumination radiation into an illumination focus located in the tissue, wherein the position of the illumination focus is adjustable to generate the detection signal.

38. The device as claimed in claim 37, wherein the illumination radiation is coupled into a light path of the treating laser radiation, and further comprising adjustable optics by which the divergence of the illumination radiation is changeable without changing the divergence of the treating laser radiation.

39. The device as claimed in claim 32, wherein the detector unit detects the radiation emitted by the tissue by confocal imaging.

40. The device as claimed in claim 39, wherein the detector unit generates the detection signal by adjusting the focus of the confocal imaging, preferably along a ray direction of the treating laser radiation.

41. The device as claimed in claim 39, wherein the optics of the detection beam path have certain light dispersing properties, so that they comprise different focal points during the confocal imaging for different spectral regions, wherein the detector unit effects a spectrally selective detection of the radiation recorded in the confocal imaging, to generate the detection signal.

42. The device as claimed in claim 41, further comprising a multi-channel spectrometer for picking up radiation behind a pinhole.

43. The device as claimed in claim 33, wherein the source of illumination radiation comprises a plurality of partial radiation sources, which are individually operable and have different spectral properties, so that spectral selective sensing is obtained by sequentially operating said partial radiation sources.

44. The device as claimed in claim 32, wherein the detection beam path has an optical axis which is located obliquely to an optical axis of the treating laser radiation or of illumination radiation.

45. The device as claimed in claim 33, wherein the source of illumination radiation causes a slit illumination of the tissue.

46. The device as claimed in claim 44, further comprising a scanning unit by which the position of the optical axis of the detection beam path is adjustable relative to the optical axis of the treating laser radiation or of the illumination radiation.

47. The device as claimed in claim 32, wherein the detector unit determines a measure of the spatial extent, the position or both of individual scattering centers, which are generated by the break-through.

48. The device as claimed in claim 32, wherein the detection signal indicates a diameter of a plasma bubble, which was generated by an optimal break-through.

49. The device as claimed in claim 32, further comprising a scanning device for scanning the tissue.

50. A method of measuring an optical break-through which is created in a tissue, beneath a tissue surface, by treating laser radiation comprising the steps of: detecting radiation emitted by the tissue from beneath the tissue surface; and determining a spatial extent of the optical break through, a position of the optical break through or both of the foregoing from detection of the emitted radiation.

51. The method as claimed in claim 50, wherein the spatial extent, the position or both of scattering centers generated by the optical break-through is determined. Application No.

52. The method as claimed in claim 50, wherein observation radiation is directed into the tissue, and radiation emitted by the tissue in the form of back-reflection is evaluated.

53. The method as claimed in claim 51, wherein the radiation emitted by the tissue is interferometrically detected.

54. The method as claimed in claim 53, wherein a position of the radiation emitted by the tissue along an optical axis of detection is determined from occurring interference.

55. The method as claimed in claim 50, wherein the radiation emitted by the tissue is detected by confocal imaging and the spatial extent is determined by adjusting a focus of said confocal imaging.

56. The method as claimed in claim 55, wherein different spectral focal points are generated in confocal imaging by dispersive optics and radiation recorded behind a pinhole is spectrally evaluated.

57. The method as claimed in claim 52, wherein spectrally different radiation is sequentially directed toward the tissue and the radiation emitted by the tissue is sequentially recorded.

58. The method as claimed in claim 50, wherein the emitted radiation is detected along an optical axis which is oblique relative to an optical axis along which the treating laser radiation or observation radiation is directed into the tissue.

59. The method as claimed in claim 58, wherein the treatment radiation is directed into the tissue as a slit-shaped beam.

60. The method as claimed in claim 58, wherein an angle between the optical axis of detection and the optical axis of irradiation is adjusted to obtain information on the spatial extent of the interaction.

61. The method as claimed in claim 50, wherein a measure of the spatial extent of individual scattering centers of the optical break-through is generated.

62. The method as claimed in claim 61, wherein a diameter of a plasma bubble is determined.

Description

BRIEF DESCRIPTION OF THE DRAWINGS

[0061] The invention will be explained in more detail below, by way of example and with reference to the Figures, wherein:

[0062] FIG. 1 shows a perspective view of a patient during a laser surgical treatment using a laser surgical instrument;

[0063] FIG. 2 shows a schematic representation of the laser surgical instrument of FIG. 1;

[0064] FIG. 3 shows the focusing of a beam onto the eye of the patient in the instrument of FIG. 2;

[0065] FIG. 4 shows a signal course during measurement of scattering centers which are generated during the laser surgical treatment in the eye;

[0066] FIG. 5 shows a schematic representation explaining a cutting path during the laser surgical treatment;

[0067] FIG. 6 shows a light source for the laser surgical instrument of FIG. 2;

[0068] FIG. 7 shows a scanning device for the laser surgical instrument of FIG. 1;

[0069] FIG. 8 shows a further embodiment of a measurement device of the laser surgical instrument of FIG. 1;

[0070] FIG. 9 shows a schematic representation of a further embodiment of the laser surgical instrument of FIG. 1;

[0071] FIG. 10 shows a beam path in the region of the focus of treating laser radiation from the laser surgical instrument acting on the eye;

[0072] FIG. 11 shows a schematic representation of focus positions in a further embodiment of a measurement device for the laser surgical instrument of FIG. 1;

[0073] FIG. 12 shows a detailed representation of the radiation directed onto a detector of the embodiment according to FIG. 11 with a pinhole;

[0074] FIG. 13 shows a functional connection between the displacement of the focus and the wavelength in the embodiment of FIG. 11;

[0075] FIG. 14 shows a further embodiment of the laser surgical instrument of FIG. 1;

[0076] FIG. 15 shows a signal course as obtained with the embodiment of FIG. 14;

[0077] FIG. 16 shows a further embodiment similar to that of FIG. 14;

[0078] FIG. 17 shows a further embodiment similar to that of FIG. 14;

[0079] FIG. 18 shows a further embodiment similar to that of FIG. 14;

[0080] FIG. 19 shows a schematic representation of a further embodiment of the laser surgical instrument of FIG. 1;

[0081] FIG. 20 shows an example of a measurement regime for the scanning of a tissue with the laser surgical instrument of FIG. 19, and

[0082] FIGS. 21 to 22 show further examples of a measurement regime for the scanning of a tissue with the laser surgical instrument of FIG. 19.

DETAILED DESCRIPTION OF THE DRAWINGS

[0083] FIG. 1 shows a laser surgical instrument 1, which emits a treatment beam 2 being directed onto the eye 6 of a patient. The laser surgical instrument 1 is able to generate a pulsed treatment beam 2 such that the method described in U.S. Pat. No. 6,110,166 can be carried out. The pulse width is in the nanosecond or femtosecond range.

[0084] A visual deficiency in the eye 6 of the patient is remedied by means of the laser surgical instrument 1 by removing material from the cornea of the eye 6 so as to change the refractive characteristics of the cornea by a desired extent. In doing so, the material is removed from the stroma of the cornea. The stroma is located beneath the epithelium and Bowman's membrane and above Decemet's membrane and the endothelium.

[0085] The material removal is effected by separating layers of tissue in the stroma by focusing the high-energy pulsed treatment beam 2 of the laser surgical instrument 1. In this case, each individual optical break-through initiates a plasma bubble, so that the separation of tissue layers covers a larger area than the focus of the treatment beam 2. By suitable deflection of the treatment beam 2, plasma bubbles are lined up during treatment. The lined-up plasma bubbles circumscribe a partial volume of the stroma: the material to be removed.

[0086] Due to the treatment beam 2, the laser surgical instrument 1 operates in the manner of a surgical knife which, without affecting the surface of the cornea, cuts material directly inside the stroma. If the cut is guided to the surface of the cornea by generating further plasma bubbles, the material of the stroma isolated by the cutting path can be laterally pulled out of the cornea and thus removed.

[0087] The cut performed by the laser surgical instrument 1 is guided according to predetermined parameters, so that the removed partial volume of the stroma causes such a change in the optical properties of the cornea that a previously existing visual deficiency is corrected to the greatest possible extent.

[0088] In order to monitor the precision of the cutting path, the laser surgical instrument 1 is provided with a measurement device which detects the spatial extents and/or the positions of the plasma bubbles generated by the treatment beam 2 in the corneal tissue of the eye. The measurement allows optimum control of the treatment beam 2. Said control may affect both the scanning movement of the treatment beam 2 and the control of the beam parameters of the treatment beam 2 with regard to the generation of the optical break-through. The diameter of the plasma bubbles generated is of great importance in guiding the laser beam during the scanning operation, because the distance between individual plasma bubble centers should not exceed the average plasma bubble diameter. Otherwise, there would be a gap in the line of plasma bubbles, i.e. there would still remain continuous tissue between the disrupted volumes. Isolation of stroma material to be removed would then be possible only with great difficulty or even not at all. In any case, the optical result would be unsatisfactory.

[0089] However, optimizing the parameters of the treatment beam 2 also allows to keep the plasma bubbles as small as possible. The smaller the plasma bubbles are, the finer will be the cut formed by the laser surgical instrument 1. This is important, in particular, if it is taken into consideration that usually lens-shaped partial volumes are to be removed from the stroma. The precision of the cutting path and the fineness of the cut are particularly important at the edges of such lens-shaped volumes.

[0090] The measured values of the measurement device may either be transformed into a display on the laser surgical instrument 1, which allows to check the parameters of the treatment beam 2 or the guidance of the treatment beam 2 as to whether desired and pre-defined values are actually present. As an alternative, it is possible to effect closed-loop control of the parameters of the treatment beam 2 or of the deflection of the treatment beam 2 by means of an automatic process control to obtain certain values. Such laser surgical instrument 1 works fully automatically.

[0091] FIG. 2 shows a first embodiment of the laser surgical instrument 1. The measurement device is realized here as an optical coherence tomograph 3 and follows the principle described in the publication Arch. Ophtalmol., Vol. 112, p. 1584, “Micrometer-scale Resolution Imaging of the Interior Eye in vivo with Optical Coherence Tomography” by J. Izatt et al.

[0092] The optical coherence tomograph 3 is integrated into the treatment beam 2 of the laser surgical instrument 1 via a coupling-in beam splitter 4 and comprises a measuring arm 3a as well as a reference arm 3b. The coupling-in beam splitter 4 couples radiation from a measuring laser 5 into the light path of the laser surgical instrument 1 such that the optical axes of the treatment beam and of the measurement laser beam coincide. The focus of the treatment beam 2 and the focus of the radiation of the measurement laser 5 are located in approximately the same position laterally (not visible in FIG. 2). Both foci are formed in the cornea of the eye 6 by focusing optics 13 arranged following the coupling-in beam splitter 4.

[0093] FIG. 3 shows how the already combined radiation of the measurement laser 5 and of the treatment beam 2 is focused into the cornea of the eye 6 as an incident beam 12 by means of the focusing optics 13. If the energy density of the treatment beam 2 (pulse length and radiation power) is above a certain threshold value, an optical break-through appears in the focus, creating a plasma bubble 11.

[0094] The radiation of the measurement laser 5 is scattered back at the plasma bubble 11, coupled out again at the dichroic coupling-in beam splitter 4 and then passes to a measurement beam splitter 7, where it is superimposed on radiation from the measurement laser 5, which was previously split off by a measurement beam splitter 7 toward a mirror 8. The position of the mirror 8 is adjustable. Thus, the measuring arm 3a extends from the measurement beam splitter 7 through the coupling-in beam splitter 4 to the eye 6, and the reference arm 3b is formed between the measurement beam splitter 7 and the adjustable mirror 8.

[0095] The radiation returning from the reference arm 3b and from the measuring arm 3a to the measurement beam splitter 7 is brought to interference on a photo receiver 9. Since the measurement laser 5 of use is a light source which exhibits a relatively short axial coherence length (temporal coherence length) of about 10 μm in combination with a high spatial coherence, an interference pattern appears on the photodiode 9 only if the optical lengths of the measurement arm 3a and of the reference arm 3b differ by not more than the temporal coherence length of the radiation from the measurement laser 5. In order to achieve a maximum in resolution using the optical coherence tomograph 3, a measurement laser 5 having a minimal temporal coherence length, which is below 10 μm in the embodiment example, is used.

[0096] In order to prevent any influence of the treatment beam 2 of the laser surgical instrument 1, which is emitted by a treating laser 10 and is directed onto the eye 6, on the optical coherence tomograph, the wavelength region of the radiation from the measurement laser 5 differs from the wavelength region of the observation radiation 2, and the coupling-in beam splitter 4 has suitable dichroic properties. For example, a measurement laser 5 is used which emits radiation at approximately 850 nm with a bandwidth as large as possible. A large bandwidth promotes axial resolution, because the bandwidth is inversely proportional to the temporal coherence length of light. When using different wavelengths for the radiation of the measurement laser 5 and the observation beam 2, superposition and separation of the beams may be effected using a dichroic coupling-in beam splitter 4. Additionally or alternatively, suitable filters may be employed in the beam path of the optical coherence tomography.

[0097] In order to measure the location of a plasma bubble 11, the mirror 8 is displaced until an interference pattern appears on the photodiode 9. If the signal from the photodiode 9 which is read out by a computer PC indicates interference of the radiations from the measuring arm 3a and from the reference arm 3b, the measuring arm 3a and the reference arm 3b have equal length, and the distance from the focus of the radiation from the illumination laser 5 in the cornea of the eye 11 is known. The accuracy of measurement is the temporal coherence length of the radiation of the measurement laser 5, i.e. the coherence length in the propagation direction.

[0098] The back-scattering of radiation of the measurement laser 5 at the eye 6 provides an indication of the discontinuities in refractive index in the cornea of the eye 6, which occur both at the boundary of different tissues, e.g. between the stroma and the Bowman's membrane, and at the site of interaction between the treatment beam 2 and the cornea of the eye 6, both below and above a threshold at which the above-mentioned optical break-through occurs with photo disruption and/or photo ablation.

[0099] In order to allow a maximum axial area to be measured, a device for changing the divergence of the radiation of the measurement laser 5 (not shown in FIG. 2 for simplification) is provided between the measurement beam splitter 7 and the coupling-in beam splitter 4, in addition to a long stroke for the mirror 8, so that the divergence of the radiation coming from the measurement laser 5 can be adjusted before coupling into the treatment beam 2. The focusing optics 13 arranged following the coupling-in beam splitter 4 then effect focusing of the measurement beam 2 and of the radiation coming from the measurement laser 5 into different focal points, wherein the adjustment of the device affecting the divergence of the radiation coming from the measurement laser 5 allows to adjust the focus position of the radiation from the measurement laser 5, independently of the focus of the treatment beam 2. Thus, the focus of the observation radiation is adjustable along the optical axis relative to the focus of the treatment radiation 2. Thus, back-scattered radiation also becomes detectable along the optical axis of the treatment beam 2 in an area which is greater than the stroke of the mirror 8.

[0100] The device for adjusting the divergence of the measurement laser 5 further allows to design the spot size of the focused radiation of the measurement laser 5 to be approximately as large as the spot size of the focused observation beam, because scanning in an axial direction, i.e. along the axis of the observation beam 2, is then achieved by synchronous adjustment of the divergence-changing device and the mirror 8. Due to the relatively narrower focusing of the radiation of the measurement laser 5 in this variant, the coherence tomograph 3 thus provides an axial resolution which is even better than the temporal coherence length of the radiation from the measurement beam 5. The precision of measurement accordingly increases. If this is not required, the device influencing divergence can be omitted between the coupling-in beam splitter 4 and the measurement beam splitter 7. Axial displacement of the area from which back-scattered radiation may lead to interference on the photo receiver, is then effected exclusively by adjusting the mirror 8, and axial resolution is mainly determined by the coherence length of the measurement laser 5 (typically in the order of magnitude of 10 μm).

[0101] FIG. 4 shows a signal obtained with a laser surgical instrument 1 which comprises an optical coherence tomograph 3 of the construction shown in FIG. 2. The signal is plotted here as a function of the adjustment of the mirror 8, with the stroke L of the mirror 8 being indicated as a variable in FIG. 4. The signal is derived from the interference phenomenon detected by the photo receiver and has been generated by the computer PC. The signal encodes the depth of modulation of said interference and has a high level whenever the signal of the photo receiver 9 indicates an interference phenomenon.

[0102] As FIG. 4 shows, a first interference appears at a displacement L1. Said interference is caused by radiation scattered back at the epithelium of the cornea. With a further displacement of the mirror 8, a second peak of the signal appears at a displacement L2. This interference is caused by the discontinuity in the refractive index at the plasma bubble 11. Said interference lasts until a displacement L3 occurs. Although the size of the plasma bubble affects the distances L2, L3, the measurement resolution of the coherence tomograph 3 does not allow measurement of bubble diameters of approximately 30 μm or less. The use of a source of radiation having a shorter temporal coherence length enables such diameter measurement as well.

[0103] The distance between the displacement L1 assigned to the epithelium and the beginning of the interference peak at the displacement L2 indicates how far beneath the epithelium, i.e. the surface of the cornea, the plasma bubble is located. Thus, the position of the plasma bubble 11, caused by the treatment beam 2 in the cornea of the eye 6, can be detected from the signal in the form of the distance from the epithelium. As an alternative to referring to the epithelium, reference may be made, of course, to the endothelium. In this case, a new peak is to be expected, due to radiation scattered back at the endothelium, in the representation of FIG. 4, upon further displacement. Instead of the endothelium or the epithelium, a discontinuity in the refractive index at the Bowman's membrane or at the Decement's membrane may, of course, be used also as a reference.

[0104] The computer PC, which reads out the signal from the optical coherence tomograph 3, generates the signal. It further serves to control the treating laser 10, for which purpose it feeds a display which indicates the position of the plasma bubble 11 generated by the treatment beam 2. On the other hand, this axial position of the plasma bubble is evaluated to control the treatment beam 2 and is used to guarantee that the step of isolating the partial volume in the stroma is effected as desired.

[0105] In an embodiment which fully automatically monitors the observance of values pre-defined by treating surgeons, the computer controls parameters of the treatment beam 2. The surgical instrument then operates using on-line monitoring and on-line control.

[0106] The measurement laser 5 has the structure shown in FIG. 6. It consists of a laser 2 as well as a subsequently arranged beam-forming unit comprising a lens 18, a stop 19 and a further lens 20. The beam diameter thus generated is adapted to the beam diameter of the treatment beam 2. By adjusting the lenses 18 and 20, the beam diameter may be additionally reduced or enlarged in order to adjust the range of resolution or measurement. In addition, divergence may be changed by adjusting the lens 20, thus allowing the focus position of the measurement beam to be set.

[0107] In a schematic representation, FIG. 5 shows the cut executed by the surgical instrument 1 under on-line control. The incident beam 12 is focused into the cornea 14 of the eye 6 by the focusing optics 13. As already explained, the plasma bubble 11, which is generated by a pulse of the treatment beam 2, acts as a surgical knife which separates layers of tissue inside the cornea 14. By adjusting the incident beam 12 by means of a scanning unit 21, which is shown in FIG. 7 and comprises scanner mirrors 22, 23 being rotatable about orthogonal axes, a plurality of plasma bubbles are placed next to each other, and a lenslet 15 is circumscribed, which is thus cut out of the stroma of the cornea 14. A possible form of cut which may achieve this is shown in U.S. Pat. No. 6,110,166, which describes a spiral-shaped line-up of plasma bubbles 14.

[0108] After such cut around the lenslet 15, a lateral cut 16 is performed, which allows to remove the now isolated lenslet 15 from the cornea 14 along the direction 17. The cutting path has the advantage that, in the central area of vision, in which the optical correction is effected by removal of the lenslet 15, there is no cut leading through the endothelium to the outside. Instead, said lateral cut 16 is located at the optically less important periphery of the cornea 14.

[0109] The scanning unit 21 shown in FIG. 7 serves to adjust the incident beam 12. It comprises two scanning mirrors 22, 23, which are independently rotatable about two axes which are perpendicular to one another. Thus, the incident beam 12 can be guided across the cornea 14 in a two-dimensional manner. Axial adjustment of the incident beam 12 is effected by adjusting the focusing optics 13, i.e. by changing the position of the focal point in which the energy density required for an optical break-through, and thus the plasma bubble 11, is generated.

[0110] In order to precisely determine the axial position of the generated plasma bubble, the mirror 8 of the optical coherence tomograph 3 is adjusted as described with reference to the signal shown in FIG. 4. This may be effected, for example, by an oscillation which passes through the area between positions L1 and L2 each time. In the case of a measurement device having high axial resolution, such a great adjustment of the mirror 8 may, however, be mechanically very complex or time-consuming in some cases. For such cases, the construction of the optical coherence tomograph 3 schematically shown in FIG. 8 is provided, which comprises several, e.g. two, reference arms 3b and 3b′ that are substantially similar to each other, i.e. they have an adjustable mirror 8 (reference arm 3b) and 25 (reference arm 3b′), respectively. Each reference arm 3b, 3b′ is associated with a photo receiver 9, 26.

[0111] The length of the reference arm 3b can be set independently of the length of the reference arm 3b′. Thus, radiation scattered back with high resolution simultaneously at the plasma bubble and at the epithelium can be measured, for example, by adjusting the reference arm 3b′ to the reflection to a reference surface, i.e. the epithelium, and reference arm 3b detecting the boundary of the plasma bubble.

[0112] Moreover, FIG. 8 shows a possible construction of the device for changing the divergence of the radiation from the measurement laser 5, which device was already mentioned with reference to FIG. 2, but not shown therein. The coherence tomograph of FIG. 8 is provided with an adjustable zoom optics 27.

[0113] The optical coherence tomograph 3 as used in the constructions according to FIG. 2 or 8 provides for a spatial selection of the area from which back-scattered radiation is recorded. Axial selection is effected by the interferometer construction comprising the measuring arm 3a and the reference arm 3b (or additional pairs thereof). The focusing optics 13 effect a lateral selection by focusing. The spatial selection is freely selectable, i.e. independently of the position of the treatment beam focus, if the optical coherence tomograph is provided with an independent scanning unit similar to FIG. 7. Then the area around an optical break-through or a plasma bubble 11 can be scanned at higher resolution. The zoom optics 27 allow sensing of the area around the plasma bubble 11 not only laterally, i.e. transversely of the optical axis of the treatment beam 2, but also axially along said optical axis, in a large measurement range.

[0114] FIG. 9 shows a second embodiment of the laser surgical instrument 1. The basic principle here is to detect radiation having a large and adjustable depth of focus, which radiation is scattered back by suitable spatial filtering at the eye 6 and, in particular, at a plasma bubble 11 generated there, in a manner allowing measurement of at least the axial extent of a scattering structure in the cornea of the eye, for example of the plasma bubble 11 generated by the optical break-through. For this purpose, a confocal microscope 28 is provided, which operates in the manner of a laser scanning microscope.

[0115] The confocal microscope 28 is provided with a laser 29 whose radiation is adapted by an optical arrangement 30 with regard to beam parameters, such as the position of necking at the focus and the beam cross-section. The illumination radiation thus obtained is passed to a scanning unit 21 by means of a splitter 31, said scanning unit comprising two scanning mirrors 22 and 23 like the scanning unit of the previous constructions. The scanning mirrors 22 and 23 are arranged closely adjacent and in immediate proximity to a pupil of the beam path. As shown in FIG. 9, the mirrors have rotary axes that are perpendicular to each other, and they may be separately controlled. For each beam deflection which depends on the actuation of the scanning unit 21, radiation is collected by scanning optics 32 in an aperture plane 35, from where an objective 36 generates a spot image which is located in an object plane situated in the eye 6. The treatment beam 2 is also effective in the area of this object plane. As in the previous embodiments, the treatment beam 2 comes from a treating laser 10 and is combined with the beam path of the confocal microscope via a coupling-in beam splitter 4.

[0116] Radiation scattered back in the object plane, e.g. at a plasma bubble 11, passes back to the beam splitter 31 via the described optical path. The radiation of the sensed volume element and coming from the object plane, i.e. radiation scattered back at the eye 6, is detected in a detecting arm 42 after having passed through the beam splitter 31. Depending on the design of the beam splitter 31, the radiation is guided, at least partially, or, in the case of a dichroic design of the beam splitter 31, almost completely, to an interference filter 38 and to a lens 39, by which a spot image of the back-scattering object in the object plane is generated in a pinhole plane in which a pinhole aperture 40 is located. The pinhole aperture 40 is followed by a photo multiplier 41 which provides a signal that is transformed into an image signal by an evaluating unit (not shown) connected thereto, which unit considers the current position of the scanning unit 21.

[0117] The size of the pinhole aperture sets the size of the object structure to be detected in the object plane. Even more decisive is, however, that a decreasing diameter of the pinhole aperture leads to an enhanced depth discrimination in the object plane, i.e. the size of the pinhole aperture determines from which axial area around the object plane radiation can reach the photo multiplier.

[0118] The confocal method used in the confocal microscope 28 allows to direct radiation from a selected volume element in the object under examination to the photo multiplier 41 and to almost completely suppress radiation emitted by other volume elements. In particular, radiation from tissue layers located further down or higher up is blocked out.

[0119] The pinhole aperture 40 and the object plane from which radiation is detected are located in conjugated planes. Adjustment of these planes, for example by adjusting the lens 39, makes it possible to effect an axial depth scan such that the object plane from which radiation reaches the photo multiplier 41 is adjusted. Said adjustment is required, for example, in order to measure a plasma bubble 11 in an axial direction, i.e. along the axis on which the treatment beam 2 is incident.

[0120] Increased independence between the axial position of the aforementioned object plane and the focal point of the treatment beam 2 may be achieved by arranging the divergence-changing device already mentioned with respect to the constructions according to FIGS. 2 and 8, e.g. in the form of zoom optics, such that it precedes the coupling-in of the treatment beam 2 via the coupling-in beam splitter 4. Suitable adjustment of this device then allows the object plane to be adjusted virtually freely relative to the focal plane of the treatment beam 2.

[0121] In the construction shown in FIG. 9, a single scanning unit 21 is provided for structural simplification, said scanning unit 21 synchronously changing the lateral position of the spot image, from which reflected radiation is passed to the photo multiplier 41, as well as the focal point of the treatment beam 2. It is not possible to laterally displace the spot within the object plane relative to the focal point of the treatment beam 2.

[0122] Use of two independent scanning units avoids this limitation. The illumination and detection beam path of the confocal microscope 28 can then be independently adjusted relative to the treatment beam 2 at the eye 6. Such a construction is indicated in broken lines in FIG. 9. The coupling-in beam splitter 4 is then replaced by a coupling-in beam splitter 4a which is arranged following the scanning unit 21 of the confocal microscope 28 in the direction of illumination. Here, the treating laser 10 is provided with its own treatment scanner 37, which is operable independently of the scanning unit 21 of the confocal microscope 28. This more complex construction allows the confocal microscope 28 to record radiation from the eye 6 at points which may be selected completely independently of the focus position of the treatment beam 2 and thus completely independently of the generation of the plasma bubbles 11. This allows not only to determine the axial extent of a plasma bubble 11, but also to measure lateral dimensions.

[0123] In the construction of the measurement device of the laser surgical instrument 1 shown in FIG. 9, separate observation radiation is generated by means of the laser 29, the optical arrangement 30 and the coupling-in mirror 31. For some applications, this may be dispensed with, and just a confocal detection of scattered light coming from the treatment beam 2 at the eye 6 may be effected. In addition, a phase-sensitive detection, e.g. according to the Nomarski method, may be applied as described in the publication by Padawer J., “The Nomarski interference-contrast microscope. An experimental basis for image interpretation.” J. Royal Microscopial Society (1967), 88, pp. 305-349, whose disclosure is explicitly incorporated by reference herein.

[0124] In the construction of FIG. 9, particularly good sensitivity is achieved, if the coupling-in beam splitter 4 or 4a, respectively, is dichroic, i.e. if it has a beam-deflecting effect essentially only for the wavelengths of the treatment beam 2. This dichroic property is also advantageous for the variant without separate illumination radiation, because it has been shown that in plasma bubbles 11 broad-band radiation is generated also outside the spectral range of the treatment beam 2.

[0125] The confocal detection method is usually effected using optics which are chromatically corrected to the best possible extent. This ensures that the radiation from a specific volume element arrives at the detector, e.g. the photo multiplier 41, in a wavelength-independent manner.

[0126] In contrast thereto, according to a third embodiment modifying the second embodiment, the optics which project light from the objective onto the pinhole aperture 40 are deliberately designed to be dispersive. In this connection, FIG. 10 shows an enlarged view of the detecting arm 42 of the confocal microscope 28.

[0127] An incident detection beam 43 is transformed by the lens 39, which is now provided as a multi-component lens, into a focused detection beam 44. Due to dispersive, i.e. wavelength-dependent, focusing properties of the optical group forming the lens 39, the effective focal length for light in the blue spectral region is clearly shorter than for light in the red spectral region. Thus, white light, which consists of red, green and blue spectral components and is emitted by a volume element, is focused into different foci, as shown by way of example in FIG. 11. The blue radiation 48 is focused into a focal plane 45, which is located in front of a focal plane 46 for green radiation, which is in turn followed by a focal plane 47 for red radiation 50.

[0128] The dispersive properties of the lens 39 are utilized for spectrally selecting depth information or for obtaining further depth information. As shown in FIG. 12, for a given volume element, the dispersive optics of the lens 39 focus only green spectral components precisely in the pinhole aperture 40, because only for these spectral components is the focal plane 46 located in the plane of the pinhole aperture 40. The focus of the blue radiation 48 is located on the optical axis in front of the plane of the pinhole aperture 40, so that the blue radiation diverges again on the way between its focal plane 45 and the plane of the pinhole aperture 40. Thus, only a small, almost unnoticeable part of blue radiation 48 is transmitted through the pinhole aperture 40, resulting in effective suppression of the blue radiation 48 coming from the given volume element. This likewise applies for the red radiation 50, because the focal plane 47 thereof is located behind the plane of the pinhole aperture 40; the red radiation from the volume element is thus also blocked by the pinhole aperture 40. In conclusion, in the schematic representation of FIG. 12, only green radiation 49 of the volume element reaches the detector, which is provided as a spectrometer 51 here.

[0129] Referring now to a further white-light radiating volume element, which is located on the optical axis between the previously considered volume element and the optics, the green component thereof is now focused in a plane behind the plane of the pinhole aperture 40. However, the blue radiation from this further volume element has a focus in the pinhole aperture, if the axial distance of both volume elements corresponds exactly to the focus displacement, with the magnification of the optical projection having to be taken into consideration as well. The same applies to a volume element, which, as seen from the objective, is accordingly located behind the first explained volume element. From this third volume element, only red components of the radiation can pass through the pinhole aperture 40. Thus, a spectral radiation mixture passes through the pinhole aperture 40, wherein the spectral information encodes the axial position of the volume element emitting said radiation.

[0130] The spectral analysis of the radiation passing through the pinhole aperture 40 by means of the spectrometer 51, therefore, provides axial resolution, without having to adjust the confocal microscope. The spectral information carries depth information about the volume element, from which the radiation of the spectral region comes. The axial position of the volume element, from which the radiation of the spectral channel originates, is determined from the relative height of individual spectral channels. In the simplest case, an evaluation using two or three channels will suffice.

[0131] Further, the size of a radiating volume element is determined in this approach. In a three-channel spectral analysis a small radiating volume element will generate, in only one channel, a signal exceeding a certain threshold value. For a medium-sized volume, two signals may also indicate a corresponding signal, but never all three channels, as long as the volume element is smaller than the focus offset between those channels that are spectrally the furthest apart. If there is a large radiating volume element, however, the spectrometer will essentially provide approximately equal signals in all channels. The width of the spectral distribution, which is indicated by the spectrometer 51, thus bears information on the size of the volume element. The wider the spectral distribution is, the larger is the volume element.

[0132] The focus offset dF achieved by the dispersive optical group of the lens 39 is plotted in FIG. 13 as a function of the wavelength LAMBDA, with all values being indicated in μm. For a spectral region from blue to red, there will be a focus displacement in the order of magnitude of at least 1 mm. Thus, the use of the aforementioned spectrally selectively detecting confocal microscope 28 allows detection of plasma bubbles having a maximum extent of up to 1 mm. Since this range of measurement usually suffices, the third embodiment can also dispense with an axial adjustment of the object plane of the confocal microscope and still provide information on the axial extent of a plasma bubble 11 sufficient for on-line control.

[0133] Of course, the spectral distribution of the radiation coming from the examined object also plays a role here. The spectral detection of confocally recorded radiation when using dispersive optics requires that radiation having a certain broad-band spectral distribution be available from the object plane. In this case, in the simplest embodiment, the radiation of a plasma bubble 11 to be detected may be sensed directly, because the plasma emits, during the process of disruption which was initiated by the treatment beam 2, radiation having a broad spectrum. For certain applications, however, the spectral distribution of the radiation may have properties that are not sufficiently constant temporally or spectrally. For these cases, external illumination is provided. Said external illumination must have the required spectral bandwidth, e.g. a white-light LED or a light bulb can be used.

[0134] In a variant of the third embodiment, no spectrometer 41 is employed. Instead, illumination of the volume element to be detected or of the spot, respectively, is effected using spectrally controlled radiation. The radiation is spectrally adjusted in a sequential manner. In a simple construction, a blue, a green and a red LED are used for illumination, which are sequentially switched on. Information on the individual spectral channels is then obtained in a time-sequential manner, so that a suitable design of the source of illumination in this modification will allow working with an inexpensive detector which exhibits no, or otherwise only insufficient spectral discrimination properties.

[0135] FIG. 14 relates to a fourth embodiment of a laser surgical instrument 1. For the measurement device, it uses the principle of a slit lamp by irradiating illumination radiation on an optical axis which extends inclined to an optical axis along which observation takes place.

[0136] The measurement device is designed as a slit-lamp arrangement 52, which illuminates the cornea 14 of the patient's eye 6 in a slit-shaped manner. The slit optics 53 image a very narrow light bundle into the cornea, along an optical axis 60 of the slit illumination. Said light bundle may be generated, for example, using a slit or a part which serves as a controllable or programmable stop, such as a micromirror array or the like. In the construction of FIG. 14, a micromirror array is used, which is illuminated with radiation. Only one or few mirror lines reflect the light to the cornea 14. The slit illumination via the slit optics 53 may be adjusted with regard to the angle of incidence, i.e. with regard to the position of the optical axis 60, in a scanning unit 54, which is indicated by the bent double arrow in FIG. 14. This is effected by suitably controlling the micromirror array.

[0137] In the cornea, the input slit-shaped illumination is scattered at scattering centers, for example the epithelium, the endothelium, or plasma bubbles 11 generated by the laser surgical instrument 1. The scattered light is recorded by an objective 56, which also focuses the treatment beam 55 into the cornea 14. The treatment beam 55 of the laser surgical instrument 1 is coupled into the optical path via the coupling-in beam splitter 4 also in this embodiment. The coupling-in beam splitter 4 separates the scattered light recorded by the objective 56 and guides it to a photo receiver 59 by means of imaging optics 58. The objective 56, the coupling-in beam splitter 4 and the imaging optics 58 form an observation beam path.

[0138] The photo receiver 59 only records radiation which comes from the intersection point between the optical axis 60 of the slit illumination and the optical axis 61 of the selection or observation beam path. A scattered-light channel 57, from which scattered light resulting from the slit illumination is recorded, is defined at the intersection point.

[0139] By the scanning movement of the slit illumination, the scattered-light channel is displaced along the optical axis 61 of the detection beam path. A rotation of the optical axis 60 of the slit illumination to the left, in the representation of FIG. 14, displaces the scattered-light channel toward the endothelium, and a rotation to the right displaces it toward the epithelium. The rotation of the optical axis 60 of the slit-shaped illumination allows scanning of the cornea 14 along the optical axis 61, and the position of a plasma bubble 11 can be precisely determined by a scanning displacement of the scattered-light channel 57 from the endothelium to the epithelium. The bubble's position can be referenced by referencing to the reflections from the endothelium or the epithelium, respectively, or Bowman's membrane with regard to the distance therefrom, thus enabling not only determination of the diameter of the plasma bubble 11, but also (absolute) position detection relative to the endothelium and the epithelium.

[0140] The signal obtained with the slit-lamp arrangement 52 is represented in FIG. 15, which shows the signal level, i.e. the radiation intensity registered by the photo receiver 59 in a signal course 63, plotted against a scanning angle α of the optical axis 60, along which the slit-shaped illumination is irradiated.

[0141] As can be seen, a first signal peak appears at an angle α0, said angle resulting from a reflection at the back-surface 64 of the cornea 14, i.e. from the reflection at the endothelium. A plasma bubble reflection 65, whose width is a measure of the extent of the plasma bubble 11 along the optical axis 61, is detected between angles α1 and α2. Finally, at the angle α3, a front-surface reflection coming from the endothelium is registered. The known thickness of the cornea 14 gauges the distance between the angles α0 and α3 as a measure of thickness, allowing to convert the distance between α0 and α1, i.e. the height of the plasma bubble above the endothelium, the distance between α1 and α2, i.e. the thickness of the plasma bubble, as well as the distance between α2 and α3, i.e. the depth of the plasma bubble beneath the epithelium, to a measure of length.

[0142] If lateral information is to be obtained, in addition to the axial information provided by the scanning movement of the slit-shaped illumination, imaging onto an imaging detector instead of the photo receiver 59 may be effected in the detection beam path.

[0143] This principle is shown in FIG. 16, which relates to a variant of the fourth embodiment. The construction of FIG. 16 corresponds largely to that explained with reference to FIG. 14. Identical parts have the same reference numerals and are, therefore, not described again.

[0144] The slit-shaped illumination is now incident centrally, i.e. coaxially to the treatment beam 55. This provides an illumination beam path 70, whose optical axis 60 coincides with the optical axis 61 of the treatment beam path. Observation is effected along an optical axis 61 which is located obliquely to the optical axis 60 of the illumination beam path. Thus, the observation beam path 67 is located at an angle to the illumination beam path, as was the case also in the embodiment of FIG. 14. However, in the construction of FIG. 16, the scanning movement of the illumination beam path is caused by a scanning unit 54 that is provided for the treatment beam anyway. The radiation coming from the scanning unit 54 is deviated again, in the construction shown in FIG. 16, through a mirror 68 which may be optionally provided as a beam splitter and enables observation of the field of operation through a microscope.

[0145] The scanning unit 54 changes the angle of incidence of the optical axis 60 relative to the eye 6, at which angle the illumination radiation of the illumination beam path 70 is incident on the cornea 14. The observation beam path 67 as well as its optical axis 61 are not adjustable in the embodiment of FIG. 16, although this may be optionally possible, of course, in order to obtain additional information. The scattered-light channel 57 is thus displaced along the (fixed) optical axis 61 of the observation beam path. In the observation beam path 67, an image receiver 69 is arranged following the imaging optics 58, said image receiver 69 recording an image of the scattered-light channel 57 located at the point of intersection of the axes 61 and 60. Evaluation of the image from the image receiver 69 yields information on the size and the position of the plasma bubble 11, and an intensity evaluation of the image in a cut leading through the plasma bubble 11 or a corresponding projection provides a signal similar to FIG. 15.

[0146] FIG. 17 shows a further variant of the fourth embodiment of FIG. 16. Parts already shown in FIG. 16 are identified by the same reference numerals and are not explained again. Like the construction of FIG. 16, the slit-lamp arrangement 52 of FIG. 17 also provides an observation beam path 67, whose optical axis 61 extends obliquely to the optical axis 62 of the treatment beam path. In the observation beam path 67, the imaging optics 58 are arranged preceding an image receiver 69. However, no additional illumination radiation is coupled in now, but instead, treatment radiation scattered directly at a plasma bubble 11, or radiation generated in the plasma bubble 11 itself, is detected. Observation is effected from an inclined view, so as to derive depth information from the signal of the image receiver 69. In addition, visual observation by means of a microscope is possible via the beam splitter 4, which no longer serves for coupling in now.

[0147] FIG. 18 shows a further variant of the fourth embodiment. The slit-lamp arrangement 52 which serves as the measurement device in the laser surgical instrument 1, corresponds essentially to the construction of FIG. 17. Now an illumination beam path 70, which irradiates a slit-shaped illumination via the objective 56, along the optical axis 62 of the treatment radiation 55, is coupled in via the coupling-in beam splitter 4. The coupling-in beam splitter 4 combines the optical axis 60 of the illumination radiation with the optical axis 62 of the treatment radiation 55.

[0148] An illumination scanning unit 71 is arranged preceding the coupling-in beam splitter 4, said unit carrying out a deflection of the slit-shaped illumination independent of the treatment radiation 55. This allows the scattered-light channel 57, from which scattered light of the slit-shaped illumination reaches the observation beam path 67, to be adjusted also laterally relative to the focus of the treatment radiation. Thus, in addition to the information obtained by the construction of FIG. 16, the signal of the image receiver 69 allows to obtain information on the scattered-light image and, thus, on the structure of the cornea laterally of the focal point of the treatment radiation 55.

[0149] Of course, the information obtained through the measurement device of the third or fourth embodiment with regard to position or size of the plasma bubble 11 is used to control the laser surgical instrument, so that open-loop and/or closed-loop on-line control are achieved here as well.

[0150] FIG. 19 schematically shows a fifth embodiment of a laser surgical instrument, comprising a pulsed source of laser radiation 71 with a pulse energy that is sufficiently large for treatment of the tissue in question; a deflecting unit 72, which laterally deflects the laser beam coming from the source of laser radiation 71; a tunable focusing unit 73, by which the position of the focal point in the depth of the tissue is set; a positioning unit 74 for positioning the tissue 6 to be treated, as well as a control unit 75 for control of the source of laser radiation 71, the deflecting unit 72 and the focusing unit 73. The control unit 75 controls the aforementioned components such that the focus of the laser radiation is sequentially focused on real target points ZP′ having the coordinates (x″, y″, z″). To this end, information on the coordinates (x″, y″, z″) of desired target points ZP has to be available to the control unit 75. Depending on the application, the positioning unit 75 may be omitted, or may be replaced by a device ensuring coarse positioning of the tissue 76.

[0151] In the laser surgical instrument of FIG. 19, there are further provided an energy reducer 77, a detector 78 and a memory unit 79. The radiation energy of the pulsed source of laser radiation 71 is attenuated by the energy reducer switched into the beam path to such an extent that the focus of the laser radiation in the tissue 76 will not cause any irreversible changes. The laser beam 80 emitted by the source of laser radiation 71 can thus be scanned in a focused manner in the tissue 76 as a measurement beam (in a so-called regime of measurement), on the one hand, and also as a treatment beam (in a so-called regime of treatment), on the other hand. As measurement beam the laser beam 80 causes a laser-induced signal S in the real measurement point MP as a function of the properties of the tissue said signal S being received by a detector 78 via the detection beam path (which is not shown in further detail). The detected laser radiation-induced signals S are supplied from the output of the detector 78 to the input of a memory unit 79 and are stored in the memory unit 79 together with the coordinates (x′, y′, z′) of the detected associated points of measurement MP′. The laser radiation-induced signals S are compared in a comparator unit 81, which is connected to the output of the memory unit 79, with threshold values S stored therein.

[0152] This selects those points of measurement onto which the treating laser beam is to be directed as target points when the energy reducer 7 is inoperative. This defines the regime of treatment. The coordinates of these selected detected points of measurement MP′ are transmitted to the control unit 75 and are available to control the deflecting unit 72 and the focusing unit 73.

[0153] Transformation of the coordinates of the detected points of measurement MP′ into coordinates of the desired target points is not required, because they are based on the same coordinate system. The coordinate system (x′, y′, z′) of the measurement is identical, with regard to the reference point, with that of the target points (x′″, y′″, z′″) as well as with that set for the treating laser radiation. Possible deviations from this identity due to instrument-specific tolerances do not interfere strongly, because they remain constant in all cases and can thus be compensated for.

[0154] FIG. 20 shows a first regime of measurement for sensing a tissue 76. Beginning at a starting point 88, the laser focus is directed onto points of measurement arranged in a grid. Different types of tissue 80a and 80b, which contact each other at a boundary 91, lead to differently laser-induced signals S. Due to the evaluated signal S, each point of measurement can be assigned to one type of tissue. Thus, the point of measurement 82a is located in the first type of tissue 80a and the point of measurement 82b is located in the second type of tissue 80b. This allows an expected area 83 to be derived for the possible location of the boundary 91. Said area is shaded in FIG. 20 from top left to bottom right. The precision with which the actual location of the boundary 91 can be indicated by the expected area 83 depends on the lateral scanning resolution 84 and on the normal scanning resolution 85.

[0155] The regime of measurement shown in FIG. 21 differs from that of FIG. 20 by a higher depth resolution, i.e. by a smaller normal scanning resolution 85. The location of the boundary 91 can be determined considerably more precisely for a constant lateral scanning resolution 84, if the boundary 91 extends approximately parallel to the surface and does not have any major slopes. Therefore, depending on the assumed course of the boundary 91, the accuracy of measurement can be increased by providing an asymmetrical scanning resolution, for example by an increase in only one direction, or the number of points of measurement 82 and thus the duration of measurement can be reduced by reducing the scanning resolutions in the other directions. The volume sensing has the advantage that all structures present therein can be recognized. Thus, for example, inclusions 89 are detected equally well as the boundary 91.

[0156] However, if there is only an interest in knowing the location of the boundary 91, the regime of measurement can be changed so as to reduce the duration of measurement, as shown in FIG. 22, which illustrates a third variant of a regime of measurement, which senses the course of a boundary 91 using as few points of measurement as possible, but at high resolution. The sensing starts at a starting point 88, which is known to be located in the first type of tissue 80a. The focus is then displaced, by the predetermined scanning stepping or resolution, toward the assumed position of the boundary 91, and the signal S is evaluated. If the focus traverses the boundary 91, the signal S changes. The depth coordinate of the boundary 91 is thus known for this lateral position. The sensing operation is then continued at the lateral scanning resolution, starting from a new starting point 86 at the same depth. However, instead of moving the focus up or down now, alternating measurements are made at increasing distances from the new starting point 86 above and below the expected depth position of the boundary 91. This is indicated by small bent arrows in FIG. 22.

[0157] In this manner, the depth position of the boundary 91 is determined at all lateral points of interest. The better the assumed position of the boundary 91 at the new lateral position matches the actual position, the less points of measurement will be required. Thus, general knowledge about the course of the boundary 91 allows a further reduction of the number of points of measurement required. Such general knowledge exists, for example, for the course of the Bowman's membrane, which is located at an approximately constant depth beneath the outer surface of the cornea and is nearly spherical. Therefore, after several points of measurement, the assumed depth position of this membrane can be predicted rather precisely, and there are only few further points of measurement required in order to determine the exact position.

[0158] Also, many tissue structures may be presumed to have a smooth course without depth variations at a large spatial frequency, so that a lower lateral resolution of ca. 0.1 mm may be sufficient. An area having a diameter of 10 mm can then be detected by means of ca. 100,000 sensing points, at a lateral resolution of 0.1 mm and a depth resolution of 1 μm.

[0159] The above-described embodiments may be employed in a particularly advantageous manner for the above-mentioned surgical method. To this end, the cornea of the eye can be aspired onto a contact glass and measured at high depth resolution (ca. 1 μm) and low lateral resolution (e.g. 100 μm) from the epithelium to the endothelium, over the entire area in which the surgical operation is to be effected. According to the fifth embodiment, the energy reducer 77 is switched on for this purpose, and the layers of the cornea are sensed at several positions in a lateral 100 μm grid, substantially perpendicular to the surface of the cornea. In the detection beam path, multi-photon fluorescence is then detected, for example, in a spatially resolved manner, at a suitable wavelength which is sensitive to differences in the different layers and/or boundaries. Alternatively, each of the aforementioned principles of measurement may be applied.

[0160] A three-dimensional image of the layer of the cornea can be generated from a plurality of thus obtained depth profiles. In said image, the laterally resolved depth position of the Bowman's membrane can be recognized, which may be of importance depending on the cutting path. For the treatment, the energy reducer is removed from the beam path, so that the desired cut is effected below the border of the epithelium when the scanning operation is carried out again. The epithelium thus remains largely uninjured, so that the cut heals again after a few days.

[0161] However, it is possible not only to remove the epithelium along the Bowman's membrane, but also to effect a cut located deeper down in the stroma. In doing so, the thickness of the stroma remaining on the epithelium can be precisely set by the previous measurement, thus excluding damage to or loss of the epithelium.