CH713869B1 - System for eye observation or therapy and methods for eye observation or therapy preparation, in particular device for laser-assisted cataract surgery. - Google Patents

Abstract

Various aspects of a system for eye monitoring or therapy are described, which individually or jointly support a user in a workflow in ophthalmology, in particular a cataract operation and its preparation. According to the first embodiment, a two-part scanner comprising a fast, short-throw and a slow, long-throw part is combined via a 4-f optic. According to the second embodiment, the radiation between the scanners and a lens is transmitted via an articulated arm, with 4 f optics being used in the articulated arm. According to the third embodiment, a main lens is provided with specific focal length characteristics to give a long working distance and hence the user's freedom of work. According to the fourth embodiment, there is a particularly precise and rapid referencing between a pre-operative biometric image and a current image during the current observation or the subsequent intervention. According to the fifth embodiment aspect, a monitoring camera is preferably coupled, which spatially impairs the patient's eye particularly little.

Description

[0001] In eye observation and therapy, systems are increasingly being used that implement various functions, such as in particular the acquisition of structural information about the eye and the carrying out of interventions on the eye using therapeutic radiation, in particular cutting or material removal using laser radiation. The invention therefore relates to a corresponding device for treating the eyes or preparing for therapy, which guides illumination or therapy radiation onto the eye.

[0002] Eye monitoring and/or therapy systems are i. i.e. R. used in a complex workflow. An example of this is laser-assisted refractive eye surgery or laser-assisted cataract surgery. In refractive eye surgery, laser radiation is used to create cut surfaces within the cornea that isolate a volume and make it removable. The volume is sized so that removing it changes the curvature of the cornea in a way that compensates for a pre-existing refractive error. In cataract surgery, the eye's natural lens, which has become opaque, is replaced with an artificial intraocular lens (IOL). To do this, a hole is cut in the front of the capsular bag of the lens of the eye. Through this hole, the lens is removed after previous fragmentation and an artificial intraocular lens (IOL) is inserted. An incision is made in the cornea and/or sclera to access the anterior chamber. Additionally, to reduce corneal astigmatism, incisions on the cornea, e.g. B. arcuate cuts possible. In the case of a so-called "post-cataract", it may be necessary post-operatively to remove the rear capsular bag in whole or in part. The term "capsulotomy" is used here for making incisions in the capsular bag (on its front and/or back side). Cataract surgery is the most frequently performed operation on the human eye and is therefore the focus of constant improvements in terms of the quality of the surgical result, efficiency in performing the operation and risk minimization. Recent developments and advances in ophthalmic femtosecond (fs) laser technology, especially in the field of refractive eye surgery, and optical coherence tomography (OCT) as an imaging technology, are increasingly automating cataract surgeries. Here, short-pulse lasers are used to “cut” eye tissue using photodisruption. This technology is hereinafter referred to as Laser Assisted Cataract Surgery (LCS).

According to current application principles, the capsulotomy (e.g. circular cutting open of the anterior capsular bag of the eye lens), lens fragmentation (dividing the eye lens nucleus), access cuts in the cornea/sclera (main access and auxiliary cuts), and possibly incisions are performed in the context of LCS carried out on the cornea using laser radiation. This laser radiation is treatment laser radiation or therapeutic laser radiation because it changes eye tissues.

[0004] US Pat. No. 6,325,792 B1 proposes focusing pulses from a femtosecond laser into the lens of the eye in order to “liquefy” the lens of the eye—this corresponds to the lens fragmentation mentioned above—or to cut open the capsular bag. The pulse focus of the femtosecond laser is positioned using ultrasound imaging.

It is disclosed in US Pat. No. 5,246,435 to focus pulses of a short-pulse laser in a three-dimensional cutting pattern in the natural lens of the eye in order to fragment the lens into fragments and thereby liquefy it through the cuts and the subsequent formation of bubbles. US Pat. No. 6,454,761 B1 proposes using optical coherence tomography (OCT) instead of ultrasound imaging for the automatic positioning of laser pulses in ophthalmic surgery on the cornea or other transparent structures, e.g. B. when removing a cataract in the eye lens to use.

The increasing maturity of femtosecond laser technology and OCT technology meanwhile allows a combination and integration of these two technologies and the establishment of largely automated femtosecond laser systems in cataract surgery. Fixed lenses and fast mirror scanners for lateral xly deflection of the laser beam in the eye and slowly adjustable lenses for z-deflection of the focus position along an optical axis of the eye are used to deflect the femtosecond pulses. Such systems are described in US 2006/195076 A1 or US 2009/131921 A1, for example. On the other hand, systems are also known in which the objective is slowly moved laterally, with a rapidly moving lens being used for the z-deflection of the focus along the optical axis of the eye. Such a system performs what is known as a lens scan and is e.g. B. described in DE 102011085046 A1.

During the first few years of development of the LCS, some application-related problems were solved, in particular by introducing a liquid interface as a mechanical-optical contact between the laser system and the eye, see US 2012/0078241 A1 or US 6019472, the integration of the technologies into a device and less the integration of the technologies in an overall workflow or a work environment in the foreground.

In particular, the interaction between the femtosecond laser and the surgical microscope, which is still necessary in cataract operations, shows considerable deficits in the systems available on the market. Most of the currently known systems are independent of the surgical microscope and, due to their size, are often located outside of the operating room later used for the actual implantation of the intraocular lens (IOL). As a result, a time-consuming repositioning and transfer of the patient to bed is usually necessary. Only recently was this deficit recognized and corresponding improvements proposed: DE 102010022298 A1 and US 2012/316544 A1 propose coupling the femtosecond laser directly and permanently to an operating microscope during the course of the operation. However, the required components are still too large according to the current state of the art, so that such a system would be too large during the IOL implantation phase and therefore too restrictive and cumbersome for the surgeon. According to WO 2008/098388 A1, a femtosecond laser is inserted under an operating microscope, quasi between the operating microscope and the patient, and docked to the eye for cornea-refractive eye surgery. Here, the surgical microscope and the femtosecond laser work more or less sequentially and independently of each other. Most importantly, they are still separate devices.

[0009] Furthermore, a number of deficits with regard to specific components have been found in the established systems, which negatively affect the quality of the operation result, the efficiency of the operation, or risk minimization. A microlens scan, as is also described in WO 2008/098388 A1, is relatively time-efficient with regard to the z-deflection for capsulotomy cuts or for lens fragmentation. However, this solution is very time-consuming for access incisions that not only provide for a small-scale movement along the optical axis of the eye, as disclosed in US 2007/173794 A1. Furthermore, in systems with fast z-deflection for the capsulotomy, the incision is time-critical. While a closed path in a lateral x/y plane does not pose a problem for the capsulotomy in fast galvoscan systems, it is safety-critical in systems with fast z-deflection, in which the path only closes after some time, that the eye can move during this time. Even with corneal access and auxiliary cuts, the advantage of a fast z-deflection of the laser beam does not come into play, since long lateral paths have to be covered here as well.

An articulated mirror arm with 4f optics is known from DE 202004007134 U1.

The various functionalities present in a combined device for eye observation or therapy must satisfy both the anatomical conditions of the human eye and the practical requirements of the environment during the operation.

The invention is therefore based on the object of specifying a system for eye observation and/or therapy that is optimized with regard to the integrated components and functions, in particular with regard to the workflow in ophthalmology. The invention is based in particular on the object of specifying a system for eye observation or therapy of the type mentioned at the outset that has the lowest possible spherical aberration over a depth range of several millimeters without disadvantageous mediation and thus achieves focusing in a very small focus.

[0013] This object is achieved by the invention with regard to various aspects of a system or a method for eye observation and/or preparation for therapy, which are defined in the independent claims.

Overall, the aspects of the invention enable the combination with other functions for a complex device, thus supporting a user in the workflow of ophthalmology, in particular in cataract surgery. The aspects can each be realized individually or in any combination of two, three or four. It is also possible to combine all aspects in one device. The latter complete combination will be explained below in exemplary embodiments with reference to the drawings. However, the invention is not restricted to realizing all aspects simultaneously, since the realization of any subsets and subgroups is advantageous and achieves the success according to the invention of supporting an operator in workflows in ophthalmology, in particular in cataract surgery.

[0015] As far as eye observation or therapy is mentioned in this description, the corresponding system should of course also cover variants that are able to optionally carry out eye observation and eye therapy or also to couple eye observation and eye therapy with one another. The same applies to the procedure for eye observation or preparation for therapy. The "or" is therefore not to be understood as an exclusive "or".

[0016] Various aspects of a system for eye monitoring or therapy are described, which individually or jointly support a user in a workflow in ophthalmology, in particular cataract surgery and its preparation.

According to a first embodiment, a two-part scanner comprising a fast, short-excursion and a slow, long-excursion part is combined via a 4-f optic. According to a second embodiment, the radiation between the scanner and a lens is transmitted via an articulated arm, with 4 f optics being used in the articulated arm. According to a third embodiment, a lens is provided with certain focal length characteristics to give a long working distance and thus the user freedom in his workflow. According to a fourth embodiment, there is a particularly precise and rapid referencing between a preoperative biometric image and a current image during the current observation or the subsequent intervention. According to a fifth embodiment, an observation camera is preferably coupled, which spatially impairs the patient's eye particularly little.

A preferred system for cataract surgery has the following features in particular: a short pulse laser source; a divergence-varying focusing unit for slow depth variation of the laser focus in the eye with a focal range of approx. 10 to 15 mm; a divergence-varying focusing unit for rapid depth variation of the laser focus in the eye with a focal range of approx. 1 mm; a fast x-y scanning system for lateral focus variation with a scan field between 1 mm and 6 mm; a compact applicator with a laterally movable lens for laser treatment on a field of approx. 13 mm, which can be oriented in different azimuths to the patient's head; an articulated arm with several joints for the free position of this applicator in space; an interface to the surgical microscope for the docking process, whereby the surgical microscope must enable observation through the applicator; a video (preferably infrared and/or green and/or green-yellow) co-monitoring of the laser treatment; an optical coherence tomography (OCT) unit to plan laser treatment with the least possible back reflections; a fixed console containing as many of the optical components as possible, but especially all the rapidly vibrating parts.

An optical device that fulfills all of the above properties must overcome a number of technical hurdles. These are exacerbated by physical-technological or biological limits. In particular, a corresponding optical system respect the anatomical contours of the eye socket of the human head, d. H. run within a conical outer contour until a minimum distance from the eye is reached; ensure a consistently good quality of the laser focus and, in particular, a high, in particular unchangeable, numerical aperture in the entire 3D treatment volume; ensure the spectral and/or geometric and/or other separation of the individual optical functionalities; design all moving parts with as little mass as possible in order not to jeopardize the treatment due to the recoil during acceleration and braking processes and to minimize the treatment time; Prevent intermediate foci on optical components, which could otherwise be damaged by the short-pulse radiation; Intermediate foci also in air, i. H. apart from optical components, produce only with a numerical aperture <0.10, preferably <0.05, in order to avoid air breakthroughs by the short-pulse radiation.

A first aspect of the invention provides for the system to be configured in a way that combines a large focus range and fast focus adjustment. These properties are usually in opposite directions, since mechanisms that allow the focal position to be adjusted over a large distance along the optical axis are usually not fast. According to the first aspect, a system for eye observation or therapy is therefore provided, which has a radiation source that provides illumination or therapy radiation, and a focusing device that bundles the radiation into a focus in an observation or therapy volume, the Focusing device has at least one focusing objective and a variable, divergence-varying optical element arranged upstream of it, which adjusts a z-position of the focus, the divergence-varying optical element having a first divergence-varying optical module with a first z-position adjustment speed and a first z-position adjustment path and a second divergence-varying optics module having a slower second z-position adjustment speed and a larger second z-position adjustment path, each divergence-varying optics module producing a plane of constant beam cross-section with variable back focus and a 4-f optics the Eb ene of a divergence-varying optical module images in an input plane of the other divergence-varying optical module and thus transmits the back focus variation between the modules.

[0021] The system must cover a large axial therapy/observation area, particularly when used for cataract surgery. If the optical focusing were carried out, for example, by changing the focal length with a constant beam diameter on the lens of the objective, then the numerical aperture would change during focusing, which would severely disturb the optical properties in the focus, e.g. the threshold energy for laser treatment. Therefore this is not permitted. According to the invention, the beam cross section in the rear focal point of the objective remains constant during the z adjustment in the first aspect. The numerical aperture in the focus is then identical for different eye-side optical focal lengths.

According to the invention, in the first aspect, (i) the divergence-varying element produces a plane of constant beam cross-section with variable optical focal lengths relative to this plane and (ii) the optical system creates an image of this plane in the rear focal point of the lens.

Examples of a divergence-varying element are a shape-changing mirror (membrane mirror, MEMS mirror, etc.), a specially shaped pair of free-form surfaces (Alvarez element) and a variable lens (liquid lens). However, a preferred implementation is a telescope, e.g. B. Galilean type, the negative lens is moved axially. Then the rear focal point of the other lens of the telescope is the plane with a constant beam cross-section and at the same time a variable optical focal length. For this purpose, the light must be collimated when it hits the negative lens. Another implementation would be a Kepler-type telescope. However, the real intermediate focus in such a telescope is unfavorable for some short-pulse applications, since if the numerical aperture is too large in the intermediate focus, breakthroughs, i. H. ionization of the air can occur. This problem is all the greater in telescopes that require high sensitivity for focusing and therefore a large numerical aperture in the intermediate image. The Galilean type is therefore preferred.

In order to achieve a fast and at the same time long-stroke z-adjustment, a combination of a module for fast focusing with a small focus range and a module for slow focusing with a large focus range is provided. The modules can be combined in any order, whereby the beam cross-section at the module location must remain the same during focusing. Separate modules must be used one after the other in the optical beam path with a non-zero distance. Then, without further measures, the concept of the constant beam cross-section would be violated. According to the invention, this problem is solved in the first aspect by a so-called 4-f system, which maps the two planes onto one another and transmits the back focus variation.

Preferably, the module with the small focus stroke (rapid focusing) is arranged first in the light direction, and then the module with the large focus stroke in the beam path.

The distance between the two lenses of the 4 f system is always the sum of the two lens focal lengths. The planes of constant beam cross-section are each located at the outer focal points of the two lenses of the 4-f system. In this configuration, it is achieved according to the invention in particular that the numerical aperture in the real intermediate image inside the telescope remains unchanged during focusing, ie in particular it does not increase. In this way, the above-mentioned ionization can be specifically prevented over the entire focus area.

Of course, this last condition does not have to be met exactly, e.g. H. the imaging of the plane of constant beam cross-section does not take place from focus to focus. However, the detuning must be so small that the changes in the numerical aperture that occur in the intermediate image during focusing are still permissible for the application.

An eye monitoring or therapy system must fit into the workflow as performed by the user. In particular, changing beds or changing the position of a patient should be avoided as far as possible. It is therefore provided in the system to transmit the radiation through an articulated arm so that the coupling of the radiation for eye observation or eye therapy is locally variable. Accordingly, according to the second embodiment, the invention provides the system for eye observation or therapy, which has an articulated arm with at least two rigid members, which are connected to one another in an articulated manner by joints that can be adjusted into different joint positions, a transmission beam path, which has an optical system which transmits radiation as a free beam with a maximum beam diameter along the articulated arm, with the optics system having deflection mirrors in the joints which deflect radiation according to the current joint position, with the optics system generating several successive levels of the same beam cross-section in the transmission beam path and each level having a 4-f Optics in the following plane maps.

For z-focusing, a divergence variation is usually performed in front of a lens by a divergence-varying optical element, as can be provided, for example, according to the first aspect. Corresponding mechanical movements that are performed to control the radiation for eye observation or therapy must be decoupled from the end of the articulated arm in front of which the patient is located. The divergence-varying optical element, in particular fast-moving components, are therefore preferably located in a fixed console. However, the change in back focus must then be transmitted along the articulated arm, i.e. over a long distance (> 1 m).

The radiation is preferably transmitted via one or more 4-f systems. For application reasons, the transmission takes place via the movable articulated arm and the rigid links. There are optical mirrors on the joints, which deflect the light according to the position of the articulated arm. However, these mirrors, as well as all other optical elements required for the transfer, must not be in any focal position in contact with the laser focus in intermediate images of the 4-f systems. However, due to the large focal range required for the application, the intermediate foci within the 4 f systems also move over large axial distances. No optical elements, i. H. neither articulated mirrors nor lenses of the 4-f system are located. This is achieved in embodiments by (i) a series connection of 4-f optics is used for the transmission, which are adapted to the articulated arm geometry required for the application. In sequential switching, the rear focal point of the second lens of the front 4 f system is located at the location of the front focal point of the first lens of the subsequent 4 f system. The 4-f systems are not necessarily 1:1 systems, but can increase or decrease the beam diameter in the gaps compared to the initial diameter, (ii) the articulated mirrors are preferably located in the areas between the 4-f systems (i.e. between the lenses mentioned in point (i)), (iii) the focal lengths and beam diameters of the 4 f systems are chosen in such a way that the following applies to the numerical aperture in the intermediate image of the 4 f system:

The upper limit of 0.05 is a preferred variant. Values up to 0.10 are generally possible. Here, STMAX means the maximum deviation of the focus position in the eye from the zero position, measured in depths of field, i. H. from the focus position with a collimated beam in front of the movable lens. The diameter of the collimated beam on either side of the 4 f system is D. The focal length of the lens on that side is f'.

For realistic therapy scenarios, the maximum focus position adjustment relative to the zero position can be ±7.5 mm, for example. With a numerical aperture for eye therapy of 0.2, for example, the depth of focus in an aqueous solution is approximately 0.034 mm. Consequently, STMAX=220 applies. For light with a wavelength of approx. 1.0 μm, the permissible range for the numerical aperture for beam diameters up to 16 mm is 0.1 (preferably 0.05)≧NA′≧20.03 .

The upper limit for the numerical aperture serves to avoid ionization of the air. For the minimum possible beam diameter for the transmission chain it therefore follows that D ≥ (2.2/0.1) *<>λ * STMAX and preferably D ≥ (2.2/0.05) *<>λ<>* STMAX.

For realistic scenarios, approximately D ≥ 5 mm follows, preferably D ≥ 10 mm. Because NA' = D/2f, the smallest permissible partial focal length of a 4-f system is f' ≥ D (2 * 0.1) ≥ 50 mm, preferably f' ≥ D (2 * 0.05) ≥ 100 mm. Since there should preferably be no articulated mirrors inside the 4-f systems, the joints must have a minimum distance of 100 mm, preferably 200 mm and more.

Since different maximum focus deviations can occur for deviations from the zero position in depth (into the eye) or in height (away from the eye), the above-mentioned condition applies to the respective lens and the respective beam diameter on which the intermediate focus is located at the feed considered.

It should also be noted that the planes of constant beam cross section do not necessarily have to be located exactly at the front or rear focal points of the lenses of the 4 f systems. Then the numerical aperture changes slightly during the focusing movement. This is permissible as long as the above Limits of the numerical aperture are not yet exceeded and no intermediate foci migrate across optical surfaces.

According to a second aspect of the invention, the system for eye observation or therapy has a radiation source that provides illumination or therapy radiation. A focusing device bundles the radiation into a focus in an observation or therapy volume. The focus covers a focus volume that has a lateral extent and an axial extent. The focus preferably has a lateral and/or axial extension of no more than 50 μm. An xy scanner deflects the focus laterally in the observation or therapy volume. The focusing device has focusing optics arranged downstream of the xy scanner and a z scanner arranged upstream of the xy scanner. This adjusts the (axial) depth of the focus. A control device controls the z-scanner for setting the depth. The focusing optics are aberration-corrected in relation to a specific setting of the z-scanner and thus a specific depth of focus. This depth represents a zero level. If the z-scanner is adjusted relative to the zero level, the penetration of the focusing optics changes and this then causes spherical aberration. In preferred embodiments of the focusing optics, this change is linear to the z adjustment. Next, the system includes an adjustable, z. B. corrective optics controlled by the control device. An adjustment of the correction optics changes the spherical aberration in the observation or therapy volume. The focusing optics are corrected in such a way that there is little or no spherical aberration in the zero plane. This is a specific setting of the adjustable corrective optics that represents a zero setting. If the z-scanner is in a position corresponding to the zero plane, the spherical aberration is minimized. This state is referred to below as "free from spherical aberrations". When adjusting the z-scanner and thus adjusting the depth of the focus from the zero plane, the correction optics is given a setting that deviates from the zero setting, and it is such that the correction optics compensate for the changes in spherical aberration caused by the focusing optics. compensated.

In a preferred embodiment of the second aspect, the focusing optics change the spherical aberration, i.e. the spherical aberration, linearly, i.e. proportionally to the deviation of the depth of focus from the zero plane, when the depth of focus deviates from the zero plane. At the same time, the correction optics are designed in such a way that their adjustment also causes a proportional change in the spherical aberration in the observation or therapy volume. The control device then only has to take into account the corresponding proportionality factors, ie linearity increases, and can easily set the z scanner and correction optics in opposite directions.

Since the z-scanner includes a mechanically moved component that performs a comparatively large stroke, it is preferred for ophthalmological applications to arrange such components as far away from the patient as possible. This avoids vibrations and noise that could irritate the patient. It is therefore preferred that the z-scanner includes a divergence-varying optical element which is arranged in front of the xy-scanner and which adjustably changes the divergence of the illumination or therapy radiation. An embodiment in which the z-scanner is designed as a telescope with a fixed converging lens optic and a movable lens optic is particularly preferred, with the correction optic being integrated into the converging lens optic and/or the movable lens optic. This integration is then carried out in such a way that the adjustment of the z-scanner from the zero plane and the setting of the correction optics, which deviates from the zero setting, are automatically in opposite directions. This configuration is particularly preferred if the radiation source provides short-pulse therapy radiation. Then the telescope will be expediently designed as a Galilean telescope.

Alternatively, it is also possible to design the correction optics separately from the z scanner. It is preferably arranged in a pupil plane so that it carries out the correction independently of the deflection angle of the xy scanner.

It can have at least one of the following elements: a shape-changing mirror, a pair of free surfaces, an Alvarez element, a variable lens, a liquid lens. The invention departs from the previous approach of correcting the optics of the system for all depths in such a way that the spherical aberration is tolerable. Instead, only a correction with regard to spherical aberrations is now carried out, which applies to a plane such as the zero plane. An opening error occurs for all other levels. This is corrected with the correspondingly controlled correction element, so that overall the focusing of the illumination or therapy radiation is carried out as free as possible from spherical aberrations.

[0042] Femtosecond laser-assisted systems are generally scanning systems. They cover the treatment field laterally with an element that scans in xy (tilting mirror, rotating prisms, etc.), the xy scanner. A divergence-varying element is usually used in front of the scanners for focusing. Examples of such divergence-varying elements include: a shape-changing mirror (membrane mirror, MEMS mirror, etc.) a specially formed pair of free-form surfaces (Alvarez element) a variable lens (liquid lens) a Galileo-type telescope (moving negative lens) a Keppler type telescope (moving positive lens)

The above problem is solved in embodiments in that (i) the focusing optics corrects all other optical aberrations in a conventional manner according to the prior art, (ii) the spherical aberration for a specific focal depth, the zero plane, with the degrees of freedom of scanning optics is corrected (the zero plane does not necessarily have to be part of the addressed focus area), (iii) the divergence-varying element is corrected for the input divergence associated with (ii) with regard to spherical aberration and (iv) for the other input divergence positions the linear change in the spherical aberration completely or at least partially with the opposite sign, i.e. compensated in total.

A completely corrected optics is therefore not necessary both for the divergence-varying element and for the focusing optics. The spherical aberration when focusing is usually corrected in the middle of the focus area. A residual error is accepted and compensated at the edges. This is advantageous when changing to shorter wavelengths and/or to a higher NA.

With the above elements, a linear change in the spherical aberration can be achieved, for example, by: a shape-changing mirror (via targeted additional variable deformation with Fringe-Zernike surface type Z9) a specially formed pair of free-form surfaces (Alvarez element) with a targeted design of the pairs of surfaces, a variable lens (liquid lens) through targeted modification of the lens surface with a fringe-Zernike surface type Z9, a telescope of the Galilean type (moving negative lens) or Keppler type (moving positive lens) through the targeted design of the spherical lenses, in particular the non-moving lenses (here the variable beam height at the stationary lens is used to generate variable contributions to the spherical aberration; possibly by using aspheres and/or additional spherical lenses).

[0046] Every system for eye observation or therapy requires a lens. This is also referred to below as the main objective. It is in front of the patient's eye. Behind the main lens is i. i.e. R. space required, e.g. B. for moving mirrors for field shifting and for the coupling of other functionalities, e.g. B. OCT The main objective converts the collimated laser light from the device into a focused laser beam with a numerical aperture sufficient for therapy. However, since the patient's eye is treated or observed at different depths, a certain axial focus adjustment range should be accessible on the eye side, i. H. the above-mentioned collimated zero position can be left. This focus adjustment range is approximately 10 to 15 mm axially. For this it is necessary that the laser light coming from the device is already pre-focused and hits the main lens. However, this naturally creates an intermediate focus in the applicator. In the case of a short focal length main lens, i. H. in the event that its focal length is comparable to the length of the focus area in the eye, the intermediate focus behind the lens moves very close to it. However, such intermediate foci must not coincide with optical surfaces, otherwise they may damage them. In the above In the case of a short focal length, the space then available after the main lens would be too limited. Rather, the focal length of the main lens must be as large as possible in order to still be able to place all elements after the main lens. The long focal length also has a positive effect on the achievable working distance, which should also be as large as possible in order to allow a sufficient distance from the patient's head. However, if the focal length of the main lens is large, even the smallest beam angle deviations in the device cause a large scattering circle of the therapy laser on the eye. In addition, the beam diameter in the zero position becomes very large, i. H. the moving optical elements (mirrors, main lens) are correspondingly heavy.

There is therefore provided according to the third aspect of the invention, a system for eye observation or therapy, comprising: a radiation source that provides illumination or therapy radiation, and a focusing device that focuses the radiation in an observation - or therapy volume bundles, wherein the focusing device has at least one focusing lens and a variable, divergence-varying optical element arranged in front of it, which adjusts a z-position of the focus, wherein the focusing device has a numerical aperture below 0.1 in the observation or therapy volume, preferably 0.05, the variable, divergence-varying optical element is designed to adjust the z-position of the focus over a range between 10 and 15 mm and the focal length of the lens is between 20 and 40 mm, preferably between 25 and 35 mm .

According to the third aspect, the focal length of the main objective is in the range between 20 mm and 40 mm, preferably between 25 mm and 35 mm. In order to nevertheless generate the working distance required for the above-mentioned joint observation, the main objective is preferably designed as a combination of a positive lens and a spaced-apart negative lens. In this case, the positive lens and/or the negative lens have one or more aspherical surfaces in order to minimize the aberrations for the therapy focus and to reduce the mass of the moving main objective. If the aspherical surface(s) were to be dispensed with, several spherical lenses with a higher mass could be used instead.

In order to minimize the pulse broadening of the short-pulse laser, the positive lens is preferably made from a crown glass with an Abbe number>50. In order to minimize the aberrations overall, the positive lens is also preferably made from a material with a refractive index >1.6.

To further minimize the pulse broadening, the negative lens is preferably designed as a cemented element or non-cemented group, i. H. a combination of a positive and a negative lens with an overall negative refractive power. The negative lens of the cemented element is preferably made from a high-index flint glass with an Abbe number <40 and a refractive index >1.7. The positive lens of the cemented element is preferably made from a high-index crown glass with an Abbe number>50 and a refractive index>1.6.

The negative lens of the main objective is preferably designed to be laterally adjustable in order to compensate for the axial coma that remains as a result of the prism tilting.

The main objective is preferably movably coupled to two mirrors in order to carry out a two-axis image field shift. It is then a movable lens within the meaning of this description.

A fourth embodiment of the invention supports the user with regard to the interaction of diagnosis or biometrics and further observation and/or therapy of an eye. In cataract surgery in particular, it is necessary, in the case of astigmatic patient eyes, to precisely align the orientation of a toric intraocular lens to be implanted or the position of corrective cuts on the cornea, which are intended to correct astigmatism, with the eye. For this reason, astigmatism axes are determined during the preoperative measurement of the eye, in so-called biometry. At the same time, a reference image is generated and the position of the axes is saved together with the reference image. The position of the axes must then be determined again on the basis of the reference image during a later observation of the eye or during a surgical intervention. It is therefore expedient to map reference structures in the reference image that allow a later current image to be assigned accordingly in order to determine the position of the determined eye structures (e.g Astigmatism axis) to be able to find again easily. In particular for an operative intervention or its preparation, it is expedient to find the main axes of an astigmatic eye and to fade them into the current image.

Therefore, in the fourth embodiment of the invention, a system for eye observation or therapy is provided, which has: a biometric device that generates at least one reference image of the eye, which contains at least one reference structure of the eye, at least a structural parameter of the eye, preferably an astigmatism axis, is determined and its position relative to the reference structure is determined, an observation or therapy device that has an imaging device for generating a current image of the eye, which also contains the reference structure of the eye, and an image processing device for identifying the reference structure of the Eye and determining its current position and determining the current position of the structural parameter based on the current image and the reference image, the biometric device generating the reference image of the eye in a spectral channel that uses a spectral range in which an absorption dye in blood vessels ate the sclera has an absorption maximum, or generated in a spectral channel in which a fluorescent dye fluoresces in blood vessels of the sclera, and the imaging device generates the current image of the eye in the same spectral channel.

The fourth embodiment provides for a special spectral channel to be used both for biometry and for subsequent registration in the current image, in which an absorbing dye that is present in the blood absorbs (e.g. with at least 30% absorptivity ) and is therefore dark and particularly high in contrast to the environment or a fluorescent dye fluoresces. The hemoglobin present in the blood, for example, can be used as an absorbing dye. It absorbs particularly well in the green spectral range, i. H. can be depicted in high contrast. The approach according to the invention makes it particularly easy and reliable to find the reference structures in the current image, which is also referred to as image registration, since the blood vessels used appear particularly rich in contrast in the reference image and the current image due to the spectral filtering.

According to the invention, according to the fourth embodiment, a reference image from the biometric measurement of the eye is used, which is filtered in a specific spectral channel in which the current image is also filtered. This eliminates the need for complicated algorithms to register the images with one another or to locate the position of the eye structures to be determined (e.g. astigmatism axes) in the current image.

A particularly preferred variant of the fourth embodiment is based on the hemoglobin in the blood. Since the hemoglobin in the blood absorbs light particularly well in the green-yellow spectral range, particularly high-contrast reference image recordings of vein structures are obtained with green or yellow or green-yellow, in particular red-free illumination. In order to be able to use these images as input data for the laser therapy (e.g. laser therapy devices for cataract operations), the therapy system also uses such an illumination. Ideally, there is a spectrally matched lighting (maximum lighting should be between 520 nm and 580 nm wavelength) directly at the device-eye interface. In a special version, the green, yellow or green-yellow light is generated by LEDs and guided to the patient's eye via a patient interface with built-in light guides.

Since both the reference image of the biometric system and the live image of the therapy system use the same lighting, very similar images of the patient's eye can be generated, which enable robust image registration.

The spectral properties of the "green light" are to be coordinated in such a way that the maximum contrast of the reference structures is achieved - the maximum illumination should be between 520 nm and 580 nm, since hemoglobin is the essential dye of the reference structures.

[0060] This “reference structure-optimized illumination” can be used both in the biometric device and in the observation/therapy device.

[0061] Optionally, the registration between the biometric, non-dilated reference image and the dilated eye sucked in with the patient interface takes place in one step. The following aspects are essential: 1. The image field that can be recorded by the patient interface is large enough so that enough reference structures (blood vessels, etc.) are visible around the limbus so that registration is possible. Image data of the iris are optionally hidden during this process so that they do not interfere with the registration. This makes the registration process more robust and is equally possible with non-dilated iris and dilated iris. 2. The lighting when both images were taken, i. H. the biometric and aspirated image is spectrally largely the same, it is e.g. B. essentially green, more yellow or green-yellow, so that vessels are very clearly visible.

In addition to the simpler handling, the advantage is also the higher precision due to the elimination of any intermediate steps.

[0063] In systems for ophthalmology, the ability to observe the eye is particularly important. This is especially true when the system applies ophthalmic therapeutic radiation to the eye. A system for eye observation or therapy is therefore provided according to a fifth embodiment of the invention, which has: a first beam path leading to the eye for first treatment or observation radiation, which runs along a main optical axis to the eye, a second beam path for second observation radiation, which runs along an optical secondary axis, and a prism splitter which, viewed towards the eye, couples the second beam path into the first beam path and, viewed away from the eye, has an entry surface and a first and a second exit surface, the prism splitter cutting the first beam path along the optical main axis between the entrance surface and the first exit surface, and the minor optical axis extending away from the second exit face, wherein the minor optical axis is ± 20° parallel to the major optical axis and the prism splitter is a combination of a Leman-Pr isma and an additional prism cemented to the Leman prism, wherein the Leman prism, viewed away from the eye, decouples the second beam path from the first beam path at a first deflection surface following the entrance surface, deflects it again at at least two deflection surfaces and leads to the main optical axis in the Substantially (± 20°) parallel or offset completely parallel to the second exit surface, the additional prism is cemented to the first deflection surface and has a surface parallel to the entry surface, which forms the first exit surface, and between the additional prism and the first deflection surface a dichroic or intensity -Partition layer is formed.

A co-observation through the second beam path is thus coupled in through a specially shaped prism.

The prism arrangement according to the invention required for this uses in its basic form an arrangement similar to that of a Bauernfeind prism, uses an additional total reflection surface to increase the co-observation beam path up, i. H. away from the patient's head. (Sprenger-Leman prism without roof edge), also includes a right angle prism cemented to the above prism. This creates the possibility of dichroic separation of the beam paths on the cemented surface. The opposing exit and entry surfaces are parallel to each other for the light passing through, Thanks to its geometry, the angle of incidence on the cemented dichroic splitter layer can be adjusted to < 30°. This allows for a simpler, production-ready design for the divider layer. The dichroic splitter preferentially directs part of the light (e.g. λ = 750...950 nm and/or yellow-green light) to the side, while the rest, e.g. B. Parts of the visible light (e.g. λ = 400 ... 700 nm) and the therapy laser light (e.g. λ = 1000 ... 1100 nm) pass through completely unhindered, is defined tilted by an angle in the range between 0.5° and 3° in order to suppress back reflections of the OCT beam path on the flat surfaces, is designed in such a way that the observation light is directed upwards on the side facing away from the patient's eye and hits the associated exit surface perpendicularly, increases the effective working distance for the therapy optics due to the material with a higher refractive index compared to air, which is beneficial for the integration of all functionalities, contains a glass path which is also tilted by the same angle in the range between 0.5° and 3° in an azimuth rotated by 90° with respect to the aforementioned prism. This compensates for the astigmatism caused by the tilting of the upper prism and deflects the back reflections of the OCT light to the same extent. The additional glass path also increases the working distance by a further amount. In addition, it can easily be integrated into the permissible cone contour. is also paired with the associated, e.g. movable lens of the therapy optics. This has a provision that can compensate for a fixed amount of on-axis coma. This precaution can be, for example, a laterally displaceable lens or a specially shaped free-form element in the pupil.

In embodiments, according to all aspects, the system can comprise a short-pulse laser system which contains a short-pulse laser source, a beam path and an applicator head for conducting short-pulse laser radiation from the short-pulse laser source onto the eye to be operated on. A short-pulse laser source is a laser source that does not emit light continuously but in pulsed form. This means that the light is emitted in time-limited portions. The pulse rates of such a short-pulse laser are usually in the femtosecond or picosecond range. However, pulse rates in the attosecond range are also possible. Due to the pulsed light emission, very high intensities can be realized, which are necessary for laser-tissue interactions via multiphoton absorption, such as e.g. B. photodisruption or plasma-induced photoablation are required. This is the case for all applications in which material is not only removed from the surface, but interactions are achieved in all three dimensions.

The beam path ensures that the short-pulse laser radiation emitted by the short-pulse laser source is guided to an exit point. It can be realized, for example, by an optical fiber or by a mirror system. The applicator head, which is connected to the end of the beam path opposite the short-pulse laser source, forms the exit point of the short-pulse laser radiation. There is the lens with multiple optical elements according to the third aspect mentioned. It is advantageous if the short-pulse laser system also has an x/y deflection system, also referred to as an x/y scanning system, and a deflection system or scanning system for the z-direction and/or a lens system that varies the divergence. The possibility of deflecting the focus of the short-pulse laser radiation in the x-direction and y-direction as well as in the z-direction in a volume that follows the exit point of the short-pulse laser radiation can also be realized by several deflection devices for one direction each, for example one scanner for slow movement over a larger area and one for very fast movement over a small area, as provided in said first embodiment.

The system optionally includes an operating microscope with a stand and a microscope head. The microscope head contains the optics and the object illumination of the surgical microscope. With such a surgical microscope, it is possible to get an optical overview of the status of the treatment at any time. The surgical microscope also contributes to the fact that an eye to be treated can be aligned with the system under optimal lighting.

The system optionally also includes a control unit which is set up to control the implementation of short-pulse laser eye surgery. The control unit can be designed in one piece or in several pieces. The components of the device are advantageously connected to the control unit via communication paths. If the control unit consists of several parts, all the components of the control unit are also advantageously connected to one another via communication paths. Such communication paths can be implemented using appropriate cables and/or also wirelessly.

The system also optionally includes a housing that encloses at least one short-pulse laser source as the radiation source, and two (in the second embodiment, therefore, an additional) articulated arms that are arranged on the housing or on an extension of the housing. Each articulated arm comprises a plurality of rigid members which are articulated to one another such that any two rigid members are connected by at least one joint.

The microscope head (if present) is arranged on an articulated arm. This articulated arm forms z. B. together with the housing a tripod of the surgical microscope. The applicator head is arranged on the second articulated arm (which in the second embodiment has the optical system), again advantageously at the end of the articulated arm facing away from the housing. The length of the second articulated arm is then such that the entire working area of the microscope head of the surgical microscope, which is arranged on the first articulated arm, can be used. The two articulated arms can therefore follow each other in all movements. In this embodiment, an interface is provided between the applicator head and the microscope head, with which the applicator head and the microscope head can be mechanically and optically connected to one another and released again. The interface is optionally characterized by a first structure on the first articulated arm and/or on the microscope head and a second structure on the second articulated arm and/or on the applicator head, which are either matched to one another according to the key-lock principle or via a Spacer can be connected to each other. Connecting the applicator head and the microscope head to each other mechanically and optically means, in addition to the mechanical connection and thus the creation of a fixed relationship between the applicator head and the microscope head, to also connect both optically to one another, so that an imaging beam path of the surgical microscope runs through the applicator head. Then there is an optical path through the applicator head for the structures of the eye to be observed with the surgical microscope.

The beam path, in particular for the short-pulse laser radiation, then passes through the second articulated arm according to the second embodiment. It is designed in such a way that it can follow all movements of the second articulated arm and in every position of the second articulated arm the radiation, e.g. B. can lead to their exit point on the applicator head in the same quality.

[0073] Furthermore, the applicator head and the microscope head can be moved three-dimensionally both independently of one another and when connected to one another. This mobility of the applicator head and the microscope head is also given when the applicator head and the microscope head are connected to one another. This requires corresponding additional degrees of freedom in the first and second articulated arm. Due to the mobility of the applicator head alone, but above all in connection with the microscope head, the exit point or short-pulse laser radiation can also be moved in three-dimensional space - in a preferred variant also with regard to its beam direction at the exit point To treat patients in a non-lying position, or in a lying position, but with an employed lying position. The volume that can be reached in space is limited by the stops of the joints.

The system can be designed in particular for short-pulse laser eye surgery, with which not only the cutting of tissue by means of plasma-induced ablation and/or photodisruption is possible, but also the gluing of tissue by means of coagulation and removal of tissue by ablation Effects of short-pulse laser radiation. Only the properties of the short-pulse laser radiation have to be set according to the application goals.

In a preferred embodiment, the system further comprises an optical coherence tomography (OCT) module containing an OCT light source, an interferometer and a detector. The OCT module can also be enclosed by the housing. It is particularly preferred to configure the OCT module in such a way that it is set up for coupling radiation emitted by the OCT light source either into the microscope head or into the applicator head. This can be done, for example, with the help of one or more optical switching points that are provided in the beam path of the radiation emitted by the OCT light source as well as the measuring light returning from the eye.

The coupling of the radiation from the OCT light source via the applicator head has the advantage that it can be overlaid with therapeutic short-pulse laser radiation in a simple and mechanically stable manner. In this way, both beam paths can be calibrated to one another. This variant is therefore used in practice for the planning and control of short-pulse laser treatment. In contrast, coupling the radiation from the OCT light source via the microscope head enables the surgeon to take tomographic images of the patient's eye during and/or after the manual operation phase. For example, with the help of this technology, intraocular lenses can be precisely aligned or free particles in the aqueous humor can be identified and removed. For coupling, it is technically advantageous to integrate a ring mirror for combining the short-pulse laser radiation and the radiation emitted by the OCT light source into the system for short-pulse laser eye surgery when the wavelengths of the short-pulse laser and OCT illumination are close together. The short-pulse laser radiation is z. B. reflected at the ring mirror, while the radiation emitted by the OCT light source of the OCT module propagates through a hole in the ring mirror in the direction of the eye and the OCT detector detects the reflected radiation of the OCT light source from the eye through the hole in the ring mirror . The ring mirror can be movable. A 90° position for coupling the radiation emitted by the OCT light source into the beam path of the short-pulse laser radiation is preferred, with the ring mirror being arranged in a 45° position. If the wavelengths of the short-pulse laser radiation and the OCT light source can be separated spectrally or with regard to polarization, then the laser and the OCT beam path can also be combined via dichroic and/or polarization splitters or combiners.

In a preferred system for short-pulse laser eye surgery, both the first articulated arm and/or the second articulated arm have at least three joints. With three joints, at least two joints, ideally all three joints, must fulfill the function of a ball joint, i. H. not only offer a possibility of rotation around a single axis. Rather, such a joint must make it possible for a rigid limb to describe any angle in space to the adjacent limb, both of which are articulated, whereby the radius of action may be restricted to a partial area of space by other structural obstacles, however not to movement within a plane. In a special embodiment, one of the three joints can only have a single axis of rotation. However, with only three joints, all three joints preferably fulfill the function of a ball joint. In this way, the optimum mobility of the first and the second articulated arm, both of which are attached to the housing or to an extension of the housing, is ensured both in the connected state and independently of one another in three-dimensional space. If, on the other hand, joints are used that each offer only one possibility of rotation about one axis, comparable mobility is achieved with at least five joints per articulated arm, which have different axes of rotation. Of these, three joints should allow rotation about vertical axes and two joints should allow rotation about horizontal axes, i. H. Represent tilting axes that lead to a tilting of the following rigid member after the joint. In this variant - i.e. when using joints each with the possibility of rotation about an axis - an articulated arm that has six joints, each with one axis of rotation per joint, is preferred. In this case, three joints should allow rotation around vertical axes and a further three joints should allow rotation allow around horizontal axes. Here the tilting of the rigid member following the joint or an end piece such as the applicator head or the microscope head is possible. Basically, the joints of each articulated arm realize at least six degrees of freedom, which are given by three vertical and three horizontal axes of rotation, whereby vertical and horizontal axes of rotation can alternate along an articulated arm. In particular, a pair of a vertical rotation axis joint and a horizontal rotation axis joint placed in close proximity to each other offers the same function as a ball joint.

[0078] Among the short-pulse laser sources, the femtosecond (fs) laser sources are by far the most frequently used laser sources in eye surgery. They have proven to be particularly suitable and easy to control for these applications. It is therefore advantageous if the system is designed for short-pulse laser eye surgery and has a femtosecond laser source.

Optionally, the system also includes a confocal detector in addition to an OCT. By recording an A-scan - i.e. a one-dimensional depth profile along the optical axis - and/or a B-scan - a two-dimensional scan along the optical axis and perpendicular to it - of two structures of an eye using the OCT and an intensity profile using the confocal detector at the An offset and a scaling factor between the OCT signals and the intensity profile can be determined by moving through a z-focus position. As a result, this allows the focus position of therapy radiation, e.g. B. the short-pulse laser radiation, using OCT signals, especially OCT images, to control particularly precisely.

Optionally, the coherence length or measurement length of the OCT light source in air is more than 45 mm, particularly preferably more than 60 mm. As a result, the entire anterior chamber section of an eye is captured within an A-scan without the optical path length of the reference beam path having to be adjusted, even if the optical path to the eye changes due to a lateral lens movement.

The displaceable image field of the system is, in particular in a short-pulse laser system for eye surgery, preferably larger than 1.0 mm but smaller than 6.0 mm in diameter, particularly preferably larger than 1.5 mm but smaller than 3. 0mm The image field is located in an image field plane in which it can be moved in the x and/or y direction by moving the lens. According to the first embodiment, the image field plane itself can be displaced along the optical axis by a scanning movement in the z-direction. The cross-section of the moveable lens depends in particular on the scanning area of the x/y scanning system. This means that the focus of the radiation can be placed at any point in the three-dimensional scan volume by superimposing the beam deflections from the movable lens and from the mirror scanners.

The optics, which are arranged in the beam path up to the articulated arm, as well as the modules that vary the divergence of the radiation, are preferably attached to an optics bench. The optics bench itself is optionally fastened with three points on, on or within a housing on which the articulated arm is preferably also arranged. All deformations of the fastening surface in the housing therefore have no influence on the adjustment status of the optics on the optics bench, but on the position of the optics bench for entry into the articulated arm with its beam guiding means. Changes in this position can be compensated for with beam stabilization.

The invention is explained in more detail below, for example with reference to the accompanying drawings, which also disclose features that are essential to the invention. 1 shows a first system for short-pulse laser eye surgery; 2 shows a second system for short-pulse laser eye surgery; Fig. 3 shows a device for independent counterbalancing of an articulated arm; 4 shows a short-pulse laser system for eye surgery (beam generation and optics); 5 shows a structure for combining short-pulse laser radiation from the short-pulse laser source and OCT radiation from the OCT light source; 6 shows two illustrations to explain how the movement of the focus of the short-pulse laser radiation affects a laterally scanning objective of a short-pulse laser system; Fig. 7 is a schematic representation through a beam splitter prism used in the beam path of the first or second system; 8 shows two illustrations for the designs of a main objective of the first or second system; FIG. 9 shows two illustrations to clarify the z-scanning technique used in the first or second system; FIG. 10 and 11 schematic drawings of a divergence-varying optical element for z-position adjustment of the focus of the first or second system; 12 shows a schematic representation relating to a two-stage z focus adjustment in the first or second system; 13 an optical beam path in which radiation is transmitted along an articulated arm in the first or second system, FIGS. 14 and 15 schematic representations of the coupling of the radiation onto the eye relating to spectrally selective illumination/detection for registering the position of an eye, FIG. 16 a schematic representation an optical system of a device for laser-assisted eye surgery, the beam path from a z-scanner to the eye being shown schematically, FIG. 17 a z-scanner of the device from FIG. 1, FIG. 18 details of the adjustment of the focus position in a cornea of the eye with the device FIG. 1 and FIG. 19 show different dependencies of the spherical aberration on the depth of the focus.

The invention is described below with reference to eye surgery, which is intended to be merely exemplary of various tasks in eye observation or therapy for which the various aspects of the invention can be used.

In the following embodiments, a short-pulse laser beam source with a femtosecond laser or fs-laser is used as the short-pulse laser, which is the short-pulse laser most frequently used in the field of eye surgery using lasers—and is therefore also the best-studied . Nevertheless, all systems described here can also be implemented with other short-pulse lasers. Unless the pulse length is explicitly discussed as a differentiating feature, fs lasers are synonymous with short pulse lasers.

As far as OCT, optical coherence tomography, is mentioned below, OCT is a synonym for all methods that can measure distances in the eye using optical short coherence or capture images of the eye or its components, such as time-domain optics Coherence tomography (TD-OCT), spectrometer-based spectral domain OCT (SD-OCT) or wavelength tuning-based swept source OCT (SS-OCT).

The system described here, in which the various aspects of the invention are implemented in combination purely by way of example, is used for laser-assisted cataract surgery. Using the short-pulse laser beam source, incisions are made, for example an access incision to the anterior chamber of the eye through the cornea, a capsulotomy incision, incisions to crush the lens nucleus of the eye or incisions on the front of the cornea to correct visual defects.

In order to improve the integration of the various components used with respect to a workflow optimized for the operator, i.e. preferably a doctor, in particular an eye surgeon, and an optimized working environment, FIGS. 1 and 2 show a first and a second System 100 for short-pulse laser eye surgery is shown, each with an fs laser system as a short-pulse laser system 200 with a short-pulse laser source 210, i.e. here an fs laser source, a beam path and an applicator head 220 for guiding the fs laser radiation on the eye to be operated on 900 included.

The structure of the first and second system 100 for short-pulse laser eye surgery also includes a surgical microscope with a surgical microscope head 320. The entire surgical microscope and its optics are in the microscope head 320 arranged.

The first system 100 for short pulse laser eye surgery of Figure 1 further comprises an OCT module 400 containing an OCT light source 405, an interferometer and a detector. In principle, the second system in FIG. 2 can also contain such an OCT module. However, the presence of an OCT module is not absolutely necessary for the interaction of the system components shown in FIGS. 1 and 2 .

The first as well as the second system 100 are controlled by a control unit, ie a control unit 500, which is either arranged centrally as here or is distributed over the system in several sub-units. For this purpose, communication paths between the control unit and individual components of the system or between sub-units of the control unit can be used. The systems 100 of FIGS. 1 and 2 further include a housing 110, which may also be referred to as a console. This housing 110 encloses the fs laser source 210 and the control unit as the central control unit 500; in the case of the first system in Fig. 1, the housing 110 also encloses the OCT module 400.

The microscope head 320 is attached to a first articulated arm 120 and the applicator head 220 to a second, separate articulated arm 130 through which the applicator head 220 the light from the fs laser source 210 is supplied. For this purpose, a beam path runs through the second articulated arm 130. The first articulated arm 120 and the second articulated arm 130 are attached to the housing 110 or an extension of the housing 110.

Two parts of an interface 150 are provided on the applicator head 220 and microscope head 320, through which the applicator head 220 and the microscope head 320 can be mechanically and optically connected to one another. In order to detach or assemble the microscope head 320 and the applicator head 220, the interface 150 has a mechanism to be switched by the doctor or automatically.

The second articulated arm 130 has the same degree of freedom of movement as the first articulated arm 120, the z. B. at the same time forms the tripod of the surgical microscope 300 . A corresponding number, arrangement and design of joints 140 of the articulated arms 120 and 130 generate the required degrees of freedom, through which the applicator head 220 and the microscope head 320 can be moved three-dimensionally in a volume both independently of one another and when connected to one another. In the case of the first system for short-pulse laser eye surgery 100 in FIG. 1, this is achieved by three joints 140 with a ball-and-socket function. In the second system 100 of Fig. 2, equivalent degrees of freedom such as that with three joints 140 with ball joint function are exemplified by three joints for rotation about perpendicular axes 140-O1, 140-O2, 140-O5 and 140-L1, 140-L2, 140 -L5 and a parallel support arm 145, which represents a rigid member of the first and second articulated arm 120, 130, with horizontal pivot joints 140-O3, 140-O4 and 140-L3, 140-L4 for the up and down movement, So a tilting movement provided.

The lengths of the rigid members of the second articulated arm 130 are designed so that the entire working area of the surgical microscope head 320 in a semicircle of 180° in front of the device, i.e. in front of the system for short-pulse laser eye surgery 100, is utilized can be.

The applicator head 220 serves to radiate short-pulse laser radiation into the eye. So it emits short-pulse laser radiation. It is fed through the articulated arm 130 from the housing 110 in which the fs laser source 210 is seated. In principle, one could think of guiding the short-pulse laser radiation in an optical fiber that runs through the articulated arm 130 . However, optical fibers are disadvantageous in the case of short-pulse laser radiation that is used for material processing, ie also in the present example of cataract surgery, with regard to the high beam intensity in the short pulses. It is therefore preferred to guide the radiation in the articulated arm 130 with free-beam optics. The free diameter of the optics is dimensioned in such a way that there is no vignetting of the laser beam when the adjustment range is used. The requirements for rigidity of the bearings and the parts of the second articulated arm 130 are reduced by automatic beam tracking. The second articulated arm 130 also offers options for passing through electrical cables, an OCT optical fiber 410 and the vacuum hoses for suction of a patient interface 600 to the patient's eye 900 and for suction of the patient interface to the applicator head 220 At the transition of the joints 140-L2/140-L3 and 140-L4/140-L5, all cables are routed outside of the joints 140 in order to avoid excessive stress on the cables against torsion. At joint 140-L1, the cables are routed concentrically to the optics through joint 140.

Depending on the embodiment variant, a parking arm 190 with a storage area for the applicator head 220 is provided on the housing 110 and/or a storage structure 190 matched to the geometry of the applicator head 220 is attached.

As Fig. 3 shows, spring elements are advantageously provided on one or both articulated arms 120, 130, which are matched to one another in such a way that the respectively assigned applicator head 220 or microscope head 320 moves within a predetermined spatial area around the Housing 110 and the operating field without external forces each holds itself in space. The applicator head 220 weighs approximately 5 kg and cannot be carried by the surgical microscope 300 or microscope head 320. The spring compensation of the first articulated arm 120, on which the microscope head 320 is arranged, is designed for up to 1 kg with viewing, eyepieces and possibly a monitor. The second articulated arm 130, on which the applicator head 220 is arranged, therefore contains a device for independent counterbalancing, as illustrated in FIG.

The weight compensation for all masses to be compensated takes place with respect to the joint 140-L3 (140-A in FIG. 3). The part of the second articulated arm 130 between the joints 140-L3 and 140-L4 is designed as a parallel support arm 145. The parallel support arm 145 essentially consists of four joints 140-A, 140-B, 140-C, 140-D and four rigid members: the first rotary head 141-1, the second rotary head 141-2, the spring arm 145-1 and the Strut 145-2. The weight compensation is realized with a compression spring 147 in the lower spring arm 145-1. The compression spring 147 pulls on a toothed belt 148, which is deflected via two toothed belt wheels 149-1 and 149-2 into the strut 145-2. There the toothed belt 148 is hung on a fastening 146-2. The compression spring 147 creates a moment about the joint 140-A which counteracts the moment created by the weight G about point A and compensates for it. The lever arm of the compensating torque is generated by the vertical distance between the toothed belt 148 and the joint 140-A. This lever arm depends on the angular position of the spring arm 145-1. The spring constant of the compression spring 147 is dimensioned in such a way that the position-dependent change in both moments is compensated. This ensures that the weight compensation is within a specified tolerance range over the entire pivoting range. The balanced weight force G is independent of the pivoting position of the articulated arm 130 for the applicator head 220. Although the pivoting of the applicator head 220 changes the distance between the center of gravity and the pivot point 140-A, this has no effect on the weight compensation. The resulting changing moment is supported by strut 145-2 which is suspended at pivot points 140-C and 140-D.

[0100] In one embodiment variant, a video recording unit and a lighting unit are provided. This can be coupled into the beam path to or from the eye 900 via the applicator head 220, as will be explained below with reference to FIGS.

In a special embodiment variant, a photonic crystal fiber with a hollow core runs through the second articulated arm 130, on which the applicator head 220 is arranged, which transmits the short-pulse laser radiation (within the hollow core and by means of periodic photonic structures analogous to a Bragg mirror) is directed. In this way - similar to free radiation - there is only a slight pulse broadening due to dispersion. The second articulated arm 130 is then only used to mechanically hold the applicator head 220, i.e. it no longer influences the beam guidance through its structure itself.

The structure of a system for short-pulse laser eye surgery 100 described here supports the positioning of the applicator head and the microscope head on the patient's eye. If the applicator head 220 and the microscope head 320 are separate, they are brought together by the operator, for example the doctor. To do this, the operator places the microscope head 320 on the applicator head 220 at the interface 150 and actuates a lock; or a mechanism automatically leads to locking when a desired position is reached. The operator guides and positions the microscope head 320 over the eye 900 to be operated on. The applicator head 220 is thus also positioned over the eye 900 . The operator looks through the eyepiece of the microscope head 320 and lowers the microscope head 320 and thus the applicator head 220 as well, if necessary with further lateral alignment of the microscope head 320 onto the eye 900, until the applicator head 220 stands in a predefined position over the eye 900 or a patient interface 600 which is detachably attached to the applicator head and which contains a contact glass 610 is in contact with the eye 900. The operator processes an eye tissue 910, ie the lens and/or the capsular bag and/or the cornea, using an fs laser. The operator lifts the microscope head 320 and thus the applicator head 220 as well. The operator brings the applicator head 220 into the parking position, and in one embodiment places the applicator head 320 on the storage surface or storage structure 190 on the housing 110 . The operator releases the microscope head 320 from the applicator head 220 using the locking mechanism, or release occurs automatically when the application head 220 is correctly positioned on the support structure 190. This separates the microscope head 320 from the applicator head 220 The operator positions the microscope head over the patient's eye 900 . The operator performs the further incisions of phacoemulsification and/or aspiration of the liquefied lens and insertion of the intraocular lens. The operator positions the microscope head 320 in a parking position away from the operating field. In one embodiment variant, the operator places the microscope head on the applicator head, which is located on the storage surface 190 on the device 100, and locks the locking mechanism, or else the locking mechanism is locked automatically when the connection is reached.

In one embodiment variant, the control unit 500 calculates the necessary data with the aid of obtained OCT images and/or video images, control commands for adjustable elements on the articulated arms 120, 130 or the applicator head 220 and/or the microscope head 320 , so that in particular steps (c) and/or (e) and, where appropriate, all further steps, with the exception of step (i), are automatically controlled by control unit 500.

In terms of construction, the housing 110, in particular the interior of the housing, is preferably designed in such a way that those components of the short-pulse laser system 200 that are enclosed by the housing, i.e. the short-pulse laser source 210 (here an fs laser source) and optical components as part of the Beam path, when assembled as a whole in and on a container can be pushed laterally over the column 310 of the surgical microscope 300. The column 310, as an extension of the housing 110, represents a support structure for the first articulated arm 120 on which the microscope head 320 is arranged. The components of the short-pulse laser system 200 enclosed by the housing 110 are therefore mounted on the base plate of the Surgical Microscope 300 deposed and attached in four places. In the second system for the short-pulse laser eye surgery 100 of FIG. 2, this takes place as close as possible to the wheels, which are fastened under the footplate as transport device 180, as a rigid fastening with a distance of about 6 mm above the footplate.

In order to guarantee the stability of the optics adjustment for the components of the short-pulse laser system 200 in the housing 110 as well as in the second articulated arm 130, various measures are possible. Elastic deformations of the supporting parts of the housing 110 due to changes in the position of the first and/or second articulated arm 120, 130 must not affect the adjustment status of the optics between the fs laser source 210 and entry into the second articulated arm 130 on which the applicator head 220 is arranged , impact. These elastic deformations are not insignificant, especially when you consider that both the first articulated arm 120 with the microscope head 320 and the second articulated arm 130 with the applicator head 220, including the device for independent counterbalancing in the form of a parallel support arm 145 and whose superstructures each have a weight in the order of 50 kg. When pivoting, the center of gravity shifts, which can lead to deformations in the range of several tenths of a millimeter. Elastic deformations of the second articulated arm 130, on which the applicator head 220 is arranged, or its rigid members, are compensated for by their own beam stabilization. On the other hand, deformations of the optics of the short-pulse laser system 200 in the housing 110, ie before entry into the second articulated arm, cannot be compensated for. However, the accuracy requirements of the console optics, ie the optics that are arranged in the housing 110 behind the short-pulse laser source 210 and in front of the second articulated arm 130, are in the micrometer range.

4 schematically shows the beam path from the short-pulse laser source to the eye, i. H. among other things by the second articulated arm 130 and the applicator head 220. FIG. 4 contains various options which are described below.

With regard to the stability of the optical structure, it is preferred that the entire optics of the short-pulse laser system 200, which is located in the housing 110 before entering the second articulated arm 130 in the beam path of the short-pulse laser radiation, including the output of the fs - Laser source 210 to be arranged on an optics bench or screwed to it. The optics bench itself is attached to the housing 110 with three points. All deformations of this fastening surface of the housing then have no influence on the alignment status of the parts on the optics bench, but on the position of the optics bench for entry into the second articulated arm 130. Changes in this position can be compensated for by stabilizing the beam path using active mirrors. A first active mirror is located directly on the optics bench. Another active mirror is located in the second articulated arm 130. A laser diode 281 in the applicator head 220 sends a laser beam through all mirrors of the second articulated arm 130 including the active mirrors to two quadrant receivers 282 in the housing 110, which are fixed to the optics bench. Deviations due to deformations when moving the second articulated arm 130 or due to the movement of the optics bench are thus recognized and can be compensated for by countermeasures on the active mirrors.

The second articulated arm 130 and the electronics or the control unit 500 are suspended from the housing. Alternating forces are caused by pivoting the first articulated arm 120, on which the microscope head 320 is arranged, or the second articulated arm 130, on which the applicator head 220 is placed are transmitted directly to the wheels 180 and the floor. The device 100 must be idle during a laser treatment. Changes in the force ratios due to unevenness in the floor have a direct effect on the adjustment status of the laser optics. In stationary operation, this influence is compensated once before each operation by the described beam stabilization.

FIG. 4 shows the fs laser system 200 for ophthalmology, in particular for cataract surgery, which contains an fs laser light source 210 . The light pulses of the pulsed laser radiation are focused into the eye 900 by a lens 225 . A controlled z-displacement of the focus of the pulsed laser radiation takes place via a varying module 212 and a second module 212, which also varies the divergence. An x/y mirror scanner 240 comprising an x-mirror scanner and a y-mirror scanner, or alternatively via a gimballed mirror scanner, or again alternatively via a post-connected x-mirror scanner for rotational rotation about the optical axis, via the second articulated arm 130 containing the mirror, the lens 225 that can be moved in x/y and a patient interface 600 containing a contact glass 610, the radiation reaches a focus on or in the eye 900

By the divergence varying modules along the optical axis - which corresponds to the z-axis - are adjusted via an adjustment mechanism controlled by the control unit 500 in the position (of its lenses to each other and on the optical axis), the divergence of the pulsed laser radiation is influenced, so that the focal position of the pulsed laser radiation along the optical axis, ie in the z-direction, is changed in the eye 900 via further fixed optical elements such as relay optics 213 . This will be explained below with reference to FIGS. 10 to 12.

The x/y movable lens 225 adjusts the lateral focal position of the pulsed laser radiation perpendicularly to the optical axis of the device, ie in the x and y direction. With a given position of the x/y mirror scanner 240, the femtosecond laser pulses are focused on a spot in the eye 900 that is approximately 5 μm wide. The position of the spot can be adjusted within the field of view of the objective 225 in the eye 900 by field shifting (by moving the objective 225) and/or by scanning using the x/y mirror scanner 240 . During the simultaneous scanning by means of the x/y mirror scanner 240 and the movement of the movable lens 225, a superimposition movement occurs. In an embodiment variant of the short-pulse laser system 200 for eye surgery, the image field of the objective 225, which is swept over by the x/y mirror scanner 240, is larger than 1 mm in cross section but smaller than 6 mm. In a preferred variant it is larger than 1.5 mm but smaller than 3 mm.

An image field that is too small means that e.g. B. with laterally smaller cuts in the eye 900, the rapid movement of the x / y scanner 240 alone is not sufficient to make a complete cut. The consequence of this is that the slow movement of the lens 225 that is then necessary means that the generation of the complete cut takes considerably longer. The field size of the lens 225 is therefore optionally selected such that access incisions in the cornea 910 of an eye 900 have a length of approximately 1.5 mm in the x-direction and a projected y-width when cutting into the depth of the cornea tissue 910 of 2 mm do not require any movement of the lens 225, only scanning with the x/y mirror scanner 240. However, the image field should not be too large either, because otherwise the lens 225 would be too heavy and therefore too sluggish and slow for large-scale movements such as B. in the capsulotomy, is.

When the microscope head 320 and applicator head 220 are coupled by means of an interface 150, the beam path for the light to be received by the microscope head 320 runs through the applicator head 220. There are alternative solutions to ensure this: In a first solution, laser optics in the applicator head 220 can be designed in such a way that the mirror 224, whose task is to reflect the laser radiation coming from the fs laser source 210 onto the lens 225 im Deflect applicator head 220, has a partial transparency - especially in the visible light range, which is required for observing the eye 900 with the microscope head 320, while the short-pulse laser radiation is almost completely reflected. A further lens 335 for adaptation to the radiation coming from the laser optics can be movably arranged in front of the objective 330 of the microscope head 320 in the beam path of the surgical microscope 300 .

In an alternative second solution, the laser optics, which then contains a fully reflecting mirror 224, can be moved into the applicator head 220 by means of a carriage. In order to use the microscope head 320 to observe the eye 900, the laser optics are removed from the beam path of the surgical microscope 300, which leads through the applicator head 220. The surgical microscope 300 cannot be used to observe the eye 900 while the short-pulse laser radiation is being used. In order to optionally create a possibility of observation, the eye 900 is illuminated with light for which the camera is sensitive, e.g. B. IR light and / or yellow-green light observed, as will be explained below.

[0115] In order to define the processing pattern in the eye 900, structures of the eye 900, in particular structures of the anterior chamber of the eye 900, are measured using optical coherence tomography (OCT). In OCT imaging, the light from a short-coherence light source is projected across the eye 900 laterally, i. H. scanned perpendicular to the optical axis of the eye 900 Light reflected or scattered by the eye 900 is brought to interference with the light of a reference beam path. The interference signal measured by a detector is analyzed. The axial distances of structures in the eye 900 can then be reconstructed from this. Structures in the eye 900 can thus be recorded three-dimensionally in connection with the lateral scanning.

4 shows the (optical) integration of an OCT module 400 into the structure of a short-pulse laser system 100 or 200. in a variant of the structure, the same OCT light source 405 is selectively coupled into the surgical microscope head 320 and into the applicator head 220. Correspondingly, the light of the OCT light source 405 reflected back by the eye 900 is superimposed on the reference light via the same interferometer and detected with the same detector. This is illustrated in FIG. 4, which includes one or more switches 420 controlled by controller 500--not shown in FIG. The switching points 420 direct the light emitted by the OCT light source 405 and the light coming back from the eye 900 from the OCT light source 405 in a first state only via the applicator head 220 and in a second state only via the microscope head 320. This enables, for example, the use of the OCT module together with the microscope head 320 when the applicator head is not needed, e.g. B. for inserting the intraocular lens (IOL). Then the applicator head 220 rests in a park position. The illumination and detection beam path of the OCT module 400 corresponds to the beam path of the fs laser radiation for making the cuts using the focus of the fs laser radiation, as a result of which alignment errors can be avoided. This is possible with the switching point or switching points 420 without having to integrate a further OCT module.

In order to further improve the integration of the OCT module 400 and to offer alternatives, another solution is also outlined in FIG. 4: The short-pulse laser system 200 shown here has the fs laser source 210 and the OCT module 400, which contains the short-coherence light source 405 and the interferometer, the fs laser radiation being fed to the applicator head 220 via the x/y mirror scanner 240 for lateral deflection and then via the second articulated arm 130 containing the mirror, the radiation of the OCT However, the short-coherence light source here runs via an optical fiber 410 (shown in dashed lines) to the applicator head 220 without being guided via the x/y mirror scanner 240. This solution has the advantage that the OCT beam path does not pass through is exposed by the second articulated arm 130 disturbing reflections.

For an integration of the OCT module 400 with the OCT short-coherence light source 405 and the interferometer, FIG from the OCT short-coherence light source 405 of the OCT module 400 on a common optical axis 215 and to realize a common optical beam path 250 to and from the eye 900 . For this purpose, the fs laser radiation 4000 coming from the fs laser source 210 hits a ring mirror 430 after fs laser beam shaping optics 211 and is reflected by it in the direction of the eye 900 . The radiation from the OCT short-coherence light source 405 of the OCT module 400, on the other hand, runs through a hole arranged centrally in the ring mirror 430 to the eye 900 and thus on the same path as the fs laser radiation. Light from the eye is also detected through the hole in the ring mirror 430 by an OCT detector arranged in the OCT module 400 . This has the advantage that primarily the high aperture areas are used for the shaping of the fs laser radiation by the fs laser beam shaping optics 211 . On the one hand, this improves the focusing. On the other hand, when the fs laser radiation is focused into the lens of an eye 900 as it continues through the eye 900 in the area of the retina, only the peripheral areas are illuminated, which reduces the risk for the patient of being damaged by the radiation in the central macular area. The ring mirror division also has the advantage that the radiation from the OCT short-coherence light source 405, i.e. the OCT measurement and detection beam, is guided onto the optical axis 215 of the short-pulse laser system 200 without a surface that is optically disruptive due to its reflections . This would not be the case with coupling in by means of a dichroic filter or (given that the radiation from the OCT short-coherence light source 405 and the fs laser radiation have almost the same wavelength) by means of a color-neutral splitter. A color-neutral division would also lead to additional intensity losses both for the radiation from the OCT short-coherence light source 405 and for the fs laser radiation.

4 shows an optional confocal detector 260 whose focal aperture is conjugate to the focal position of the fs laser radiation. This confocal detector 260 also makes it possible to measure structures of the eye when scanning the focus of the fs laser radiation in all spatial directions.

In the description it was explained that the lens 225 is movable. In one embodiment, this can take place in that a conventional lens comprising at least one main lens is moved laterally. Upstream deflection mirrors can be used, the distance between which is adjusted. This is shown schematically in FIGS. 6a and 6b, which also show the use of the annular mirror 430 purely by way of example. This ring mirror 430 is one way of superimposing the f1 laser radiation 4000 on the OCT radiation 406 . It is essential that the short-pulse laser radiation 4000 and the OCT beam path with the OCT radiation 406 on the common optical axis 215 impinge on the movable lens 225 .

6a shows the two end points of the displacement of the radiation through the movable lens 225 in a left and right representation. FIG. 6b shows the structure schematically in three dimensions. The movable lens 225 has two deflection mirrors 4010 and 4012, between which the beam path runs transversely to the axis onto the image field 4002. A displacement of the mirror 4012 can thus be used to set where the radiation strikes a lens body 4014 of the movable objective 225 along an optical axis 4000 . This lens body 4014 is shown in FIG. 6 in the y/x plane, ie perpendicular to the z plane, unlike the preceding elements 4012, 4010, 430. An image field 4002, in which the radiation can be adjusted in the x/y plane by the movable lens 225, is also drawn in perpendicularly to the z-plane. 6b shows that the lens body 4014 of the objective 225 is fixedly coupled to the mirror 4012, ie moves with it. This has advantages with regard to the imaging quality, since the lens body 4014 of the objective 225 is always irradiated correctly with respect to its optical axis. The three-dimensional representation of FIG. 6b clearly shows the adjusting effect of the mirrors 4010 and 4012 with the entrainment of the respective downstream elements (lens body 4014 of the objective 225 together with the mirror 4012 or mirror 4012 and lens body 4014 when the mirror 4010 moves).

The illustration on the left of FIG. H. can occur when the mirror 4012 is fully extended. The illustration on the right accordingly shows the shortest beam path when the mirror 4012 is fully retracted. Due to the sectional view, Fig. 6a only shows the adjustment of the mirror 4012 in the form of a shortening of the beam path between the mirrors 4010 and 4012. Of course, this shortening would only result in a shift in the image field 4002 realized along an axis. The second axis of displacement is realized either by rotating the mirror 4010 and pivoting the mirror 4012 and the lens body 4014, or by shortening the distance between the mirror 4013 and the mirror 430 while simultaneously moving the mirror 4012 and the lens body 4014.

Due to the large coherence length of the OCT light source 405 in air greater than 45 mm, particularly preferably greater than 60 mm, it is possible for the entire anterior chamber section to be captured within an A-scan given by tuning the swept-source source even if the optical path to the eye 900 is lengthened or changed as a result of the lateral lens movement, without the optical path length of the reference beam path z. B. must be adjusted by moving a reference mirror. In order to compensate for the influence of the movement of the lens 225 on the OCT signal, the path length differences—typically up to 6 mm for different lens positions—are preferably taken into account when calculating the A-scans from the OCT signals.

[0124] As can be seen in FIG. 4, a second beam path for observation, for example onto a camera 360, is coupled via a prism 350. FIG. A Bauernfeind prism is used in the representation of FIG. 4 . However, the prism splitter 1000 of FIG. 7 is preferred for the system. It couples a first beam path 1001 and a second beam path 1002. The first beam path 1001 falls through the (main) objective 225 onto the eye 900. The second beam path 1002 also runs to the eye 900, but is divided by the prism plate 1000. The prism splitter 1000 consists of a Leman prism 1003. This type of prism is also referred to as a Sprenger-Leman prism or Leman-Sprenger prism. It is, for example, from the publication Lexikon der Physik, Spektrum academic publisher Heidelberg, 1998, or the publication H. Haferkorn, "Optics: physical-technical principles and applications", Vieweg Verlag, there in Fig. 5.131, or S. Flügge , "Optischeinstrumente / Optical Instruments", Springer Verlag, there page 218, Fig. 21, known. This prism, referred to below as the Leman prism, is supplemented in the prism splitter 1000 by an additional prism 1004 which is cemented to the deflection surface 1005 of the Leman prism 1003 which is closest to the eye. The additional prism 1004 ensures that the optical axis of the first beam path 1001 is not deflected when passing through the prism splitter 1000 . In contrast to the Bauernfeind prism of FIG. 4 , the second beam path 1004 is emitted at its exit surface 1006 parallel to the first beam path 1001 after it has exited at the exit surface 1007 . The first and second beam paths propagate coaxially at the entry surface 1008 .

The prism splitter 1000 is slightly tilted relative to the optical axis between the lens 225 and the eye 900, for example between 0.5 and 3 degrees, in order not to reflect illumination radiation incident through the lens 225 back into the beam path. This would be disruptive, for example, in the case of an OCT. In order to compensate for the tilting, a plane-parallel compensation prism 1009 is also provided between the entry point 1008 and the eye 900, which lies at the same tilting angle as the prism splitter 1000 in the superimposed first and second beam path, but with an azimuth rotated by 90 degrees. An astigmatism produced by the tilting of the prism splitter 1000 is thereby compensated. In addition, this advantageously increases the working distance.

The main objective 225 in the applicator head 320 is movable, e.g. B. by means of the structure explained with reference to FIG. After the main objective 225, as FIG. 4 shows, there must be space for coupling in by means of further beam paths, for example the beam splitter 350/1000. In order to avoid an intermediate focus after the main lens, the main lens must have a certain focal length. From this point of view, the focal length of the main objective should therefore be selected to be large in order to be able to place further elements, for example the prism splitter 1000, after the main objective 225. A long focal length also has a positive effect on the working distance, so that a sufficient distance to the patient's head is possible. From another point of view, however, the focal length of the main lens should be small. With large focal lengths, even the smallest beam angle deviations in the device cause a large scattering circle of the laser radiation on the eye, which is very annoying, especially with short-pulse laser radiation, and must be avoided. In addition, the beam diameter in the zero position of the main lens becomes very large. A main objective with a focal length between 20 mm and 40 mm, preferably between 25 mm and 35 mm, achieves a particularly good balance. Furthermore, the main objective is designed as a combination of a positive lens and a negative lens spaced therefrom, as shown schematically in FIG. 8, which contains two different variants for the main objective shown side by side. In both figures, the main objective consists of a positive lens 2001 and a negative lens 2002, which together produce a focus 2000, which is seen from the last lens element, the positive lens 2001, at a working distance d. The negative lens 2002 widens the beam path, the positive lens 2001 focuses it into the focus 2000 with a large working distance. The positive lens is preferably made of crown glass with an Abbe number of at least 50. In order to minimize aberrations overall, it is preferably made of a material with a refractive index of at least 1.6. In one embodiment, which is contained in the illustration on the right in FIG. The order of these lens elements is not relevant. Preferably, the positive lens member has the same material properties as the positive lens 2001 and the negative lens member is made of flint glass having an Abbe number of not more than 40 and a refractive index of not less than 1.7. The two lens elements can be cemented together. In this configuration of a three-part main objective 225, a large working distance is advantageously combined with aberration correction.

[0127] In cataract surgery in particular, a large axial depth range must be covered in which the laser radiation can produce an optical breakthrough. The numerical aperture should not change. In the conditions shown on the left in Fig. 9, in which a focal length change with a constant beam diameter is carried out on the main objective 225 (in the case of the multi-part construction of Fig. 8, for example on the positive lens 2001 or on a plane lying between the positive lens 2001 and the negative lens 2002 ), the numerical aperture changes depending on the position of the focus 2000. It is therefore intended to design the beam path in such a way that the beam cross section does not remain constant in the main plane 2005 of the main objective 225, but in the rear focal plane 2006 of the main objective 225. Then even with different z-positions of the focus, i. H. with different focal lengths on the eye side, the numerical aperture in the focus 2000 remains unchanged. In the system 200 this is achieved in that the divergence-varying element is designed as a Galilean telescope 2010 (cf. FIG. 10), which is composed of a negative lens 2011 and a positive lens 2012. The negative lens 2011 is shifted, which is illustrated by an arrow in FIG. As an alternative to a Galileo-type telescope, the Kepler-type telescope shown in FIG. 11, which is constructed from two positive lenses 2012 and 2013, is also possible. However, it has a real intermediate focus 2014, which can be disadvantageous depending on the application, particularly in the case of short-pulse laser radiation for material processing, since there, in cases of large numerical apertures, optical breakthroughs, i. H. ionization of the air. It is therefore preferable for the first and second systems 200 to perform the divergence variation using a Galilean type telescope 2010 .

In order to be able to carry out a z-variation that is as fast as possible and at the same time long-throw, two divergence-varying elements, for example Kepler-type telescopes, are connected in series, with a divergence-varying element having a short adjustment path and accordingly enabling rapid adjustment, whereas the other divergence-varying element performs a long-stroke, comparatively slower adjustment in the z-direction. To maintain a constant beam cross-section, the exit plane of one divergence-varying module 2010 is mapped into an entry plane 2017 of the next divergence-varying module. This is shown in FIG. 12. The imaging takes place using a 4 f system comprising two positive lenses 2018 and 2019. The distance between the two positive lenses of the 4 f system is always the sum of the two lens focal lengths. The level of constant beam cross-section 2016 and 2017 is located at the outer focal points of the two lenses 2018 and 2019. The numerical aperture in the real intermediate image 2020 inside the 4-f system remains unchanged during focusing, i.e. it does not increase in particular . It can then be designed in such a way that, with short-pulse laser radiation, there is no optical breakthrough in the real intermediate image, i. H. Intermediate focus 2020 is created. In this way, unwanted air ionization can be specifically prevented over the entire focus adjustment range.

In FIG. 10, the rear focal plane 2006 of the main objective 225 was entered as the plane of constant beam cross section. This is only an example to explain the function of the divergence varying module. In order to have as few moving parts as possible near the patient, in particular not the moving negative lens 2011, the plane of constant beam cross section that results on the output side of the divergence-varying module 2010 is transferred to the rear focal plane of the lens 225 via a transfer image. In this way, the mechanical focusing movements, ie the mechanical vibrations from the applicator head 320, which is in contact with the patient's eye, can be decoupled and the divergence-varying modules can be arranged in the housing 110 and in particular on the optical system mentioned. This transmission is carried out by a beam path through the articulated arm 130 by means of at least a 4-f system. 13 shows this transmitted beam path in two different positions of the divergence-varying element 2010, only a Galilean telescope being drawn in for the sake of simplicity. The transmitted beam path transmits planes 2016 of constant beam cross-section into the rear focal plane 2006 of the main objective 225, which forms the focus in the eye 900. The position of this focus in the eye 900 depends on the setting of the divergence-varying element 2010. A zero position results in a zero plane 2030 around which the focus can be adjusted into a maximum front plane 2031 or a maximum rear plane 2032. The maximum distance from the null plane 2030 represents a maximum depth of field STMAX. FIG. 13 shows the beam path for the two extreme positions of the divergence-varying optical element 2010, which in the specific embodiment, as explained with reference to the previous figures, is realized by two modules that are not shown in FIG. The plane 2013 of constant beam cross-section, which is present at the exit of the divergence-varying element 2010, is transferred into the rear focal plane 2003 of the main objective 225 by at least one 4-f system, two are shown in FIG. A first 4 f system 2022 has two positive lenses 2023 and 2024 of equal focal length, which are in the known 4 f configuration so that an intermediate focus 2025 is formed between them. The same applies to a second 4-f system 2026, which has two positive lenses 2027 and 2028 and an intermediate focus 2029. The planes of constant beam cross-section 2016 and the rear focal plane 2006 each lie in the focal plane of the associated 4-f system or the corresponding positive lens. Because of the large focus adjustment that the divergence-varying element 2010 generates, the intermediate foci 2025 and 2029 move over comparatively long axial distances. The corresponding deflection mirrors along the articulated arm 130 are located between the last lens of the divergence-changing element 2010 and the first lens 2023 of the first 4-f system or the last lens 2024 of the first 4-f element 2022 and the first lens of the following 4-f system. f element. Generally speaking, the deflection mirrors are each outside of the respective 4-f systems. The distances between the 4-f systems, i. H. between the lenses 2023 and 2024 or 2027 and 2028 are sections of the transmitted beam path which lie within a rigid member of the articulated arm. In this way, it is prevented that the intermediate foci 2025 and 2029, which migrate over a wide area, lie on optical elements of the transmitted beam path. The equations and focal length information given in the general part of the description apply to the focal lengths and beam diameters of the 4-f systems. The same applies to the dimensions of the rigid links and the distances between the joints.

In order to make the workflow as simple as possible for the operator, the patient interface 600, which is necessary for optical reasons and contains a contact glass 610, is the structure shown in FIGS. 4 and 15 for processing the eye 900 using a system for the Short Pulse Laser Eye Surgery 100 shown. The structure shown here includes a patient interface 600, an applicator head 220 of the system for short-pulse laser eye surgery 100, the patient interface 600 in Fig. 14 both on the patient's eye 900 and on the applicator head 220 of the system for short-pulse laser eye surgery 100 is fixed and thus the position of the eye 900 relative to the system for short-pulse laser eye surgery 100 and consequently also to the beam path of the short-pulse laser radiation is fixed.

The patient interface 600 contains a contact glass 610, which is designed as a liquid interface in the exemplary embodiment shown here. The contact glass 610 is in one piece, made of a preferably uniform, transparent material and contains a suction ring 612, a jacket 611 and an optical element 620 on the top of the jacket 611. It also includes two openings 613,614 to which two leads are connected via fixing aids or allow the connection of two feed lines, one feed line each being or being connected to one of the openings 613, 614.

A one-piece contact glass 610, in which all functional elements are integrated, allows easier handling than multi-component contact glasses 610, which are first assembled on the patient's eye 900. Such multi-component contact glasses 610 are z. B. in the documents US 7955324 B2, US 8500723 B2, US 2013/053837 A1, WO 2012/041347 A1 described.

The two supply lines are used on the one hand to apply negative pressure, here via the lower opening 613, and on the other hand to supply or drain liquid into the contact glass 610, when the contact glass 610 is docked on the eye 900, via the upper one Opening 614.

In a preferred variant, an overflow outlet 615 is also provided in the upper jacket region of the contact glass 610, remote from the eye 900, via which excess liquid or air can escape from the contact glass 610 when it is being filled.

The patient interface 600 preferably contains a mechanically detachable coupling element 651 for mechanically fixing the contact glass 610 to the applicator head 220. Alternatively, it is possible that the patient interface 600 instead of a mechanical interface with a mechanically detachable coupling element 651, a contact glass Contains 610 with a further suction structure, which is made of the same material as the contact glass 610. This further suction structure holds the contact glass 610 on the applicator head 220 when negative pressure is applied. Since this is an alternative solution, this is not shown in FIG.

[0136] A patient interface 600, which additionally contains an applicator head protection 650, which preferably has a recess in the middle, is particularly advantageous for sterility. This applicator head protection 650 can be placed over the side of the applicator head 220 facing the eye 900 and fixed at the same time, as shown in FIG. 14 . This applicator head protection prevents the applicator head 220 from being contaminated by liquids, for example, during the operation. The recess allows the patient interface 600 to be attached directly to the applicator head 220 with the contact glass 610, so that the applicator head protection 650 is not an obstacle in the beam path of the short-pulse laser radiation between the system for short-pulse laser eye surgery 100 and the optical element 620 of the contact glass 610 represents. If the recess is realized in the center of the applicator head protection 650, spatially uniform protection of the applicator head 220 is achieved.

The applicator head protection 650 is advantageously connected to the applicator head 220 by a mechanically detachable coupling element 651 .

In order to support the docking and in particular the lateral alignment of the application head 220, an illumination system that is particularly suitable for short-pulse laser eye surgery is disclosed in FIGS 635 embedded. In turn, a light source 630-1, which emits visible light and/or a light source 630-2, which emits light of a specific spectral composition, is integrated into the applicator head 220 of one of the system for short-pulse laser eye surgery 100. In particular during the surgical intervention using short-pulse laser radiation in the eye 900, in which the short-pulse laser radiation is guided into the eye 900 via optical elements of the applicator head 220 and thereby blocks the optical path into a microscope head 320 located above or is affected, the eye 900 can be illuminated with light 630-2, for example, and the light reflected from the eye 900 via a beam splitter prism 350, which z. B. selectively reflected infrared and green or yellow light, in a camera 360, which can detect this light, are directed. The prism 350 preferably reflects wavelengths that are used by the short-pulse laser source 210 or by the OCT light source 405 . Light of these wavelengths not reflected by prism 350 passes through prism 350 without interference.

This structure has the advantage that, compared to the alternative solution of illumination by an illumination present in the surgical microscope 300, no reflections are added by the additional optical elements of the applicator head 220 located in the illumination beam path and impair the image.

It is also advantageous if a force sensor 655 is integrated in the applicator head 220, which is in contact with the contact glass 610 when the patient interface 600 is docked. The force sensor 655 and both the visible light-emitting light source 630-1 and the light source 630-2 emitting the light of the specific spectral composition are advantageously connected to a control unit 500, which also controls the system for short-pulse laser eye surgery 100, or alternatively with an additional control unit 500', which is in contact with the control unit of the system for short-pulse laser eye surgery 100 via communication paths.

In order to transmit data measured preoperatively, e.g. B. the axis position of the preoperatively measured astigmatism of the eye 900 or the cornea 910 or the target position of access incisions or relaxation incisions in relation to the preoperatively measured astigmatism axes of the eye 900 or the cornea 910 must also be correctly aligned with the eye during the operation can, in the prior art, the preoperative data or desired target positions are defined or referenced relative to reference marks obtained preoperatively. Of course existing marks such as vascular structures in the sclera or iris structures or simply an overall image of the eye 900 with its existing structures are used as reference marks.

[0142] In order to recognize these vessel structures in the sclera particularly well, it is provided to provide green or yellow illumination sources 3000, which illuminate the eye in a spectral range in which the hemoglobin of the blood absorbs particularly well. In this way, a particularly high-contrast image is obtained, which is recorded on the camera 360, for example. Corresponding examples for possible locations of the illumination source 3000 are entered in FIGS. The green illumination 3000 is configured to illuminate the limbus 3001 and portions of the sclera of the eye 900. In this way, a high-contrast image of the vascular structure in the sclera is obtained.

[0143] This vessel structure is used for referencing to a previously generated reference image that was obtained in the biometry of the eye before the eye observation currently taking place or the therapeutic intervention to be carried out, for example cataract surgery. This reference image can have been generated with the same system 100 or 200, for example. In addition to the reproduction of the reference structures, for example the blood vessels in the sclera, it also indicates the related position of structural features of the eye, for example an astigmatism axis. The eye is also illuminated in green during the biometric measurement, which is carried out preoperatively. Thus, high-contrast images of the same reference structures are available both in the reference image and in the current image, so that the current position of the eye can be referenced to the reference image and thus the correct specification of the eye structures, for example an astigmatism axis, in the current image is possible without any problems.

[0144] Alternatives to using green or yellow illumination are, of course, illumination with white light and corresponding spectral filtering of the reference image or the current image, or illumination with white light that is spectrally filtered along its path, so that a single illumination source can be used for multiple purposes.

However, the use of green illumination has the advantage that the high-contrast image can be generated with a camera 360 that is already present. For this purpose, it is then preferred that the splitter layer in the beam splitter 350 has a suitable dichroic design in order to couple the green illumination onto the limbus and the sclera and to use a camera 360 that is sensitive in the green or yellow range.

16 schematically shows the beam path of an fs laser system for ophthalmology, in particular for cataract surgery. This can in particular be one of the systems described above. Light pulses of pulsed laser radiation 5002 are focused into the eye 900 by focusing optics 5008 . A controlled z-displacement of the focus of the pulsed laser radiation 5002 takes place via a divergence-varying module, which implements a z-scanner 5004. An xy-scanner 5006, which z. B. comprises an x-mirror scanner and a y-mirror scanner or alternatively via a gimballed mirror scanner or alternatively via an x-mirror scanner with a downstream element for rotation around the optical axis, the radiation reaches a focus 5018 on or in the eye 900. A control unit 500 controls the scanners 5004, 5006.

The divergence of the pulsed laser radiation 5002 is influenced by the z-scanner 5004, so that the focus position of the pulsed laser radiation 5002 along the optical axis, ie in the z-direction, in the eye 900 is changed via the focusing optics 5008.

[0148] The xy scanner 5006 sets the lateral focal position of the pulsed laser radiation 5002 perpendicular to the optical axis of the device, ie in the x and y directions. The femtosecond laser pulses are focused on a spot in the eye 900 that is approximately 5 μm laterally extended. The position of the spot can be set laterally within the image field of the focusing optics 5008 in the eye 900 by scanning using the xy scanner 5006 . The z-scanner 5004 generates the depth setting. It is shown in more detail in FIG. The z scanner 5004 is designed as a Galilean telescope and includes a movable negative lens 5010, which is longitudinally adjustable in a guide 5012 along the optical axis OA. It works together with a fixed positive lens 5014 and adjusts the divergence of the laser radiation 5002.

18 schematically shows the adjustment of the laser radiation 5002 in the cornea 5016 of the eye 900. The xy scanner 5006 adjusts the position of the optical axis OA laterally. In the event of a displacement Δxy from the rest position, a depth adjustment Δz must be carried out at the same time due to the curvature of the cornea 5016 in order to keep the focus 5018 on a desired path within the cornea 5016. In this case, the curvature of the cornea 5016 is adjusted to a known extent by a contact glass 600 . The contact glass 600 fixes the eye 900.

The z-scanner 5006 can e.g. B. be designed as a Galilean telescope 2010, which, as shown in Fig. 10, is composed of the negative lens 2011 and the positive lens 2012. The negative lens 2011 is shifted, which is illustrated by an arrow in Figs. 10 and 17, to perform the divergence variation and as a result shift the z-position of the focus 5018. As an alternative to a Galileo-type telescope, a Kepler-type telescope is also possible, which is made up of two positive lenses. In order to avoid a real intermediate focus, which can be disadvantageous depending on the application, especially in the case of material-processing short-pulse laser radiation, since optical breakdowns, i. H. ionization of the air, the 2010 Galileo-type telescope is preferred.

[0151] In order to generate the focus with as little spherical aberration as possible over a wide z-adjustment range in high quality, the focusing optics 5008 are designed in such a way that they are corrected for a specific position of the z-scanner 5004 with regard to spherical aberration. This position of z-scanner 5004 corresponds to a null plane. It is preferably in the middle of the z area to be covered. When the z scanner is moved out of this zero plane, the focusing optics 5008 cause an aperture error (also known as spherical aberration), which increases linearly with the distance from the zero plane. This relationship is shown in FIG. 19, which shows a curve 5016 for the spherical aberration F of the focusing optics 5008 as a function of the adjustment of the depth of the focus 5018. With curve 5016 alone, ophthalmic device 5000 would be unusable. However, the z-scanner fixed lens group 5014 is designed to produce a spherical aberration satisfying curve 5020 as lens group 5010 is displaced. It thus acts as a correction lens that automatically compensates for the aperture error of the 5008 focusing lens. Thus, when the control unit 500 is used to set the z position of the focus 5018 over a wide z range, the spherical aberration is compensated. In contrast, the correction in the focusing optics 5008 only has to be carried out for the zero plane, so that the correction effort in the optics elements is drastically reduced.

The curve of Figure 19 provides a particularly low-control mechanism. However, for this purpose the focusing optics 5008 and the compensating element, in this embodiment the optics element 5014, must be matched to one another as precisely as possible. This adaptation can be reduced at the expense of greater control effort if correction optics that are independent of the z-scanner and cause an adjustable, known spherical aberration are used. The control unit 500 then controls this correction optics in such a way that the aperture error, which is generated by the focusing optics 5008 as a function of the z-position, is compensated for. In this embodiment, too, the focusing optics 5008 have no spherical aberration in a zero plane. This zero level corresponds to a zero setting of the correction optics, in which these also do not produce any spherical aberration. In the embodiment of FIG. 1, this zero setting would conveniently be the center position of the z-scanner.

With separate correction optics, it is also possible to compensate for a non-linear progression, such as that shown in dotted lines as curve 5022 . Also, the course of the curve for the correction optics and the focusing optics 5008 does not necessarily have to be opposite. All that is required is that the corresponding curves of the spherical aberration are known and stored in control unit 500 as a function of the z-position.

Claims (15)

1. Eye observation or therapy system comprising: - a radiation source (L) providing illumination or therapy radiation, and - A focusing device that bundles the radiation into a focus in an observation or therapy volume, wherein the focusing device has at least one focusing lens (225) and a variable, divergence-varying optical element arranged upstream of it, which adjusts a z-position of the focus, wherein - the divergence-varying optical element has a first divergence-varying optical module (212) with a first z-position adjustment speed and a first z-position adjustment path and a second divergence-varying optical module (212) with a slower second z-position adjustment speed and a larger second z-position adjustment path, characterized in that - Each divergence-varying optical module (212) generates a plane with a constant beam cross-section with a variable optical focal length and - A 4-f optics images the plane of one divergence-varying optical module (212) in an input plane of the other divergence-varying optical module (212) and thus transmits the back focus variation between the modules (212). 2. System according to claim 1, characterized in that both divergence-varying optical modules (212) are designed as telescopes with a fixed converging lens and an upstream additional movable lens, with the radiation source (L) in particular providing short-pulse therapy radiation and the telescope with a moving negative lens. 3. The system of claim 1, comprising: - an articulated arm (120, 130) with at least two rigid members which are articulated to one another by joints which can be adjusted to different joint positions, - a transmission beam path, which has an optics system, which guides radiation as a free beam with a maximum beam diameter along the articulated arm, the optics system having deflection mirrors in the joints, which deflect radiation according to the current joint position, - Wherein the optics system in the transmission beam path produces several successive levels of the same beam cross-section and images each level by a 4-f optics in the following level. 4. System according to claim 3, characterized in that the following applies to a numerical aperture NA in the intermediate image of each 4-f optic: 0.1>=NA>=2.2*(λ*STMAX)/D, where λ is the wavelength or mean wavelength of the radiation, D is the beam diameter and STMAX is a maximum adjustment path of the z-position of the focus by the divergence-varying optical element, and/or that a numerical aperture NA in the intermediate image of each 4-f optic is between 0.1 and 0.03, with the wavelength or mean wavelength of the radiation being 1 µm ± 0.1 µm and the beam diameter being less than 20 mm and greater than 5 mm is greater than 10 mm, with the rigid members preferably being at least 100 mm, preferably at least 200 mm, long and a smallest partial focal length of the 4-f optics being over 50 mm, preferably over 100 mm. 5. System for eye observation or therapy, in particular as a device (5000) for laser-assisted eye surgery, the system having: - A radiation source (L), the illumination or therapy radiation (5002), - A focusing device that bundles the radiation (5002) into a lateral and/or axial focus (5018) with a maximum extension of 50 µm in an observation or therapy volume, - an xy scanner (5006) for the lateral deflection of the focus (5018) in the observation or therapy volume, - wherein the focusing device (5008) has a focusing optics (5008) arranged downstream of the xy scanner (5006) and a z scanner arranged upstream of the xy scanner (5006), which adjusts a depth position of the focus (5018), - a control device (500) which controls the z-scanner for adjusting the depth position of the focus, characterized in that - The system has adjustable correction optics (5014) controlled by the control device (500), the adjustment causing a change in the spherical aberration in the observation or therapy volume, - The focusing optics (5008) is corrected for aberrations in such a way that the spherical aberration is corrected for a specific setting of the adjustable correction optics (5014) and a specific setting of the z-scanner and thus the depth of the focus (5018), the specific setting of the adjustable correction optics (5014) represents a zero setting and the determined setting of the z-scanner represents a zero plane and wherein the focusing optics are designed in such a way that they cause a change in the spherical aberration in the observation or therapy volume when the z-scanner is set outside the zero plane, - The control device for setting the depth of the focus (5018) sets the z-scanner out of the zero plane and the correction optics (5014) compensates for the change in spherical aberration caused by the focusing optics (5008). 6. System according to claim 5, characterized in that the focusing optics (5008) are designed such that when the z-scanner is adjusted from the zero plane, it causes a change in the spherical aberration in the observation or therapy volume that is proportional to this adjustment, and/or that the z-scanner comprises a divergence-varying optical element (5010) which is arranged upstream of the xy-scanner and which adjustably changes the divergence of the illumination or therapy radiation. 7. System according to claim 6, characterized in that the z-scanner is designed as a telescope with a fixed lens optic (5014) and a movable lens optic (5010), the correction optics (5014) being contained or realized in the fixed lens optic (5014). so that the adjustment of the z-scanner from the zero plane and the setting of the correction optics (5014) that deviates from the zero setting are automatically coupled, with the radiation source (L) in particular providing short-pulse therapy radiation and the telescope having a moving negative lens . 8. System according to any one of claims 5 to 7, characterized in that the correction optics (5014) is arranged in a pupil, in particular in front of the xy scanner, and that the correction optics (5014) are controlled by the control device (500). 9. Eye monitoring or therapy system comprising: - a radiation source (L) providing illumination or therapy radiation, and - A focusing device that bundles the radiation into a focus in an observation or therapy volume, wherein the focusing device has at least one focusing lens (225) and a variable, divergence-varying optical element arranged in front of it, which adjusts a z-position of the focus, thereby marked that - the focusing device in the observation or therapy volume implements a numerical aperture of less than 0.1, - the variable, divergence-varying optical element is designed to adjust the z-position of the focus over a range between 10 mm and up to 15 mm and - the focal length of the lens is between 20 mm and 40 mm, preferably between 25 and 35 mm. 10. System for eye monitoring or therapy according to any one of claims 1, 3, 5 and 9, which comprises: - a biometric device which is designed to generate at least one reference image of the eye, which contains at least one reference structure of the eye, to determine at least one structural parameter of the eye, preferably an astigmatism axis, and to determine its position relative to the reference structure, - An observation or therapy device, which has an imaging device for generating a current image of the eye, which also contains the reference structure of the eye, and an image processing device for identifying the reference structure of the eye and determining its current position and determining the current position of the structure parameter based on the current one image and the reference image, - wherein the biometric device is designed to generate the reference image of the eye in a spectral channel in which an absorption dye in blood vessels of the sclera has an absorption maximum, or in a spectral channel in which a fluorescent dye fluoresces in blood vessels of the sclera, and - wherein the imaging device is designed to generate the current image of the eye in the same spectral channel. 11. System for eye monitoring or therapy according to any one of claims 1, 3, 5, 9 and 10, which comprises: - a first beam path (1001) leading to the eye for first therapy or observation radiation, which runs along a main optical axis to the eye, - A second beam path (1002) for second observation radiation, which runs along an optical secondary axis, and a prism splitter (1000) which, viewed towards the eye, couples the second beam path (1002) into the first beam path (1001) and, viewed away from the eye, has an entry surface (1008) and a first and a second exit surface, the prism splitter separating the first beam path (1001) along the main optical axis between the entry surface and the first exit surface, and the minor optical axis runs away from the second exit surface, - the minor optical axis being parallel to the major optical axis by ± 20°, and - Wherein the prism splitter (1000) is designed as a combination of a Leman prism (1003) and an additional prism (1004) cemented to the Leman prism (1003), wherein -- the Leman prism (1003), viewed away from the eye, decouples the second beam path (1002) from the first beam path (1001) at a first deflection surface following the entry surface, deflects it again at at least two deflection surfaces and offset parallel to the main optical axis to the second exit surface leads, -- the additional prism (1004) is cemented to the first deflection surface and has a surface parallel to the entry surface (1008) which forms the first exit surface, and - A dichroic or intensity splitter layer is formed between the additional prism (1004) and the first deflection surface. 12. System according to claim 11, characterized in that the second beam path (1002) is incident on the splitter layer at an angle of at most 30 degrees to the normal and/or that the splitter layer is dichroic and reflects radiation in the wavelength range from 750 to 950 nm or 600 to 500 nm into the second beam path and transmits radiation in the visible wavelength range and in the wavelength range from 1000 to 1100 nm for the first beam path. 13. System according to one of claims 11 or 12, characterized in that the Leman prism (1003) has no roof surface and consequently the first deflection surface is flat. 14. System according to any one of claims 11 to 13, characterized in that the prism splitter (1000) in the first beam path (1001) is tilted by an angle of 0.5 to 3 degrees, so that the entry surface and the first exit surface by the angle deviating from perpendicularity to the main optical axis, in particular A plane-parallel glass block is arranged upstream or downstream of the prism splitter (1000) in the first beam path (1001), which is also tilted by an angle of 0.5 to 3 degrees in the first beam path, with an azimuth of 90 degrees compared to the tilting of the prism splitter (1000 ) is twisted, the glass block preferably being arranged downstream of the prism splitter (1000) as seen towards the eye, and optionally in the first beam path (1001) between the prism splitter (1000) and the eye there is a lens for bundling the first treatment or observation radiation into or onto the eye, which has an optical coma compensation element to compensate for axis coma caused by the tilting of the prism splitter (1000), has, in particular a laterally displaceable lens or a free-form element in a pupil. 15. A method for eye observation or therapy preparation, comprising the steps of: - providing a system according to any one of claims 10 to 14; - generating a reference image of the eye containing a reference structure of the eye, - Determining at least one structural parameter of the eye, preferably an astigmatism axis, and determining a relative position of the at least one structural parameter of the eye to the reference structure, - Generation of a current image of the eye, which also contains the reference structure of the eye, - identifying the reference structure of the eye and determining its current position and determining the current position of the structure parameter using the current image and the reference image, characterized in that - the reference image of the eye is generated in a spectral channel in which an absorption dye in blood vessels of the sclera has an absorption maximum, or is generated in a spectral channel in which a fluorescent dye in blood vessels of the sclera fluoresces, and - the current image of the eye is generated in the same spectral channel.