Patent Description:
This disclosure is generally related to laser eye surgery, and more specifically to laser eye systems and methods for treating vitreous humor including ocular floaters.

Vitreous floaters are small particles consisting of cells, pigment, or fibrin that move in the vitreous of the eye. Patients with opaque vitreous humor floaters can suffer from blind spots and deteriorated vision. Vitreous surgery can improve visual acuity in these patients. Traditionally, vitreous surgery (vitrectomy) was performed by cutting the eye to remove the floaters with mechanical surgical tools, such as a vitreous infusion suction cutter that cut the vitreous and removed the debris from the eye by suction. Other vitrectomy methods have included using a nanosecond pulsed laser to tediously steer the laser visually to treat the floaters, thereby subjecting the retina to shock waves, mechanical distortions, as well as direct laser exposure of energy levels needed to treat the floaters effectively.

<CIT> discloses a laser eye surgery system used to treat vitreous bodies including a laser source, a ranging subsystem, an integrated optical subsystem, and a patient interface assembly.

The techniques and systems disclosed herein provide many advantages over the current standard of care.

The present invention provides a full depth ophthalmic surgical system for performing surgery on an eye of a subject, which includes: a femtosecond laser source configured to produce a femtosecond pulsed laser beam; an imaging assembly configured to emit an imaging beam; a scanning assembly including a Z scanner and an XY scanner, configured to scan a focal spot of the laser beam and the imaging beam within the eye in a depth direction and two transverse directions, respectively; an illumination light source configured to emit an illumination light; a video camera assembly; an objective lens assembly configured to focus the laser beam and the imaging beam, wherein the objective lens assembly includes four doublet lenses and a meniscus lens; a patient interface configured to be coupled to the objective lens assembly and to engage the eye, the patient interface including a lens having a predefined optical power, wherein the lens is a double-convex focusing lens; and optical components including at least one beam splitter, configured to direct the laser beam and the imaging beam output by the scanning assembly and the illumination light to the objective lens assembly, and to direct light emitted from within the eye, which has passed through the objective lens assembly, to the video camera assembly; wherein the scanning assembly, the objective lens assembly and the lens of the patient interface are configured to forma focal spot of the laser beam at any depth within a range of <NUM> to <NUM> in water beyond a distal surface of the lens of the patient interface. The video camera assembly includes a detector and a tunable lens in front of the detector, and wherein the tunable lens of the video camera assembly is configured to focus light emitted from any distance within a range of <NUM> to <NUM> in water beyond the distal surface of the lens of the patient interface. The system further includes a fixation light source configured to generate a fixation light, wherein the optical components are further configured to direct the fixation light to the objective lens assembly. TheXY scanner includes two scanning mirrors, wherein the objective lens assembly and the lens of the patient interface are configured to form respective images of the two scanning mirrors at locations <NUM> to <NUM> from a distal surface of the lens of the patient interface.

This summary and the following detailed description are merely exemplary, illustrative, and explanatory, and are not intended to limit, but to provide further explanation of the invention as claimed. Additional aspects, features, objectives and advantages of the invention will be set forth in the descriptions that follow, and in part will become apparent from the written description, taken in conjunction with the accompanying drawings, illustrating by way of example the principles of the invention, or may be learned by practice of the invention. The scope of the invention is defined by the set of claims which follow.

The techniques and systems disclosed herein provide many advantages over the current standard of care. Specifically, a removeable focal point extension assembly used with an existing laser ophthalmic surgical system provides a full depth ophthalmic surgical system capable of performing surgical procedures along the entire length of the eye from the cornea to the retina. A method for calibrating the full depth ophthalmic surgical system for procedures performed posterior to the lens, by using the focal zone of the optical coherence tomographer beam as a proxy for the focal zone of the femtosecond laser source, provides a fast and efficient calibration methods that can be done for each individual patient. A treatment method and system performed in the vitreous humor achieves liquification of the vitreous humor, which allows floater inclusions to move to a location where the central field of view of the eye is not affected by the floater, without disturbing other structures.

The present invention can be implemented by a system that projects or scans an optical beam into a patient's eye <NUM>, such as system <NUM> shown in <FIG> which includes an ultrafast (UF) light source <NUM> (e.g. a femtosecond laser, or a dual purpose system capable of emitting pulses in a lower and in a higher range of pulse energies, perhaps with different pulse durations. Using this system, a beam may be scanned in a patient's eye in three dimensions: X, Y, Z. In this embodiment, the UF wavelength can vary between <NUM> to <NUM> and the pulse width can vary from <NUM> fs to <NUM> fs. The pulse repetition frequency can also vary from <NUM> to <NUM>. Safety limits with regard to unintended damage to non-targeted tissue bound the upper limit with regard to repetition rate and pulse energy; while threshold energy, time to complete the procedure and stability bound the lower limit for pulse energy and repetition rate. The peak power of the focused spot in the eye <NUM> and specifically within the crystalline lens <NUM> and anterior capsule of the eye is sufficient to produce optical breakdown and initiate a plasma-mediated ablation process. The peak power of the focus spot is also sufficient to treat the vitreous humor, e.g. to treat floaters or liquify the vitreous humor, as described later. When treating the vitreous humor, the laser focus spot may be larger than when treating other parts of the eye, so the peak power will be higher to maintain sufficient peak fluence. Near-infrared wavelengths are preferred because linear optical absorption and scattering in biological tissue is reduced across that spectral range. As an example, laser <NUM> may be a repetitively pulsed <NUM> device that produces <NUM> fs pulses at a repetition rate of <NUM> and an individual pulse energy in the ten microjoule range. Although not illustrated, UF Light Source <NUM> may be further configured to provide higher energy pulses with the same or longer pulse durations than those exiting the system after pulse compression. That is, the un-compressed beam may be extracted from UF Light Source <NUM> in order to provide those higher energy pulses. Regardless, the following system description details means to achieve the usage of higher and/or lower energy pulses.

The laser <NUM> is controlled by control electronics <NUM>, via an input and output device <NUM>, to create optical beam <NUM>. Control electronics <NUM> may be a computer, microcontroller, etc. In this example, the entire system is controlled by the controller <NUM>, and data moved through input/output device IO <NUM>. A graphical user interface GUI <NUM> may be used to set system operating parameters, process user input (UI) <NUM> on the GUI <NUM>, and display gathered information such as images of ocular structures.

The generated UF light beam <NUM> proceeds towards the patient eye <NUM> passing through half-wave plate, <NUM>, and linear polarizer, <NUM>. The polarization state of the beam can be adjusted so that the desired amount of light passes through half-wave plate <NUM> and linear polarizer <NUM>, which together act as a variable attenuator for the UF beam <NUM>. Additionally, the orientation of linear polarizer <NUM> determines the incident polarization state incident upon beamcombiner <NUM>, thereby optimizing beamcombiner throughput.

The UF beam proceeds through a shutter <NUM>, aperture <NUM>, and a pickoff device <NUM>. The system controlled shutter <NUM> ensures on/off control of the laser for procedural and safety reasons. The aperture sets an outer useful diameter for the laser beam and the pickoff monitors the output of the useful beam. The pickoff device <NUM> includes of a partially reflecting mirror <NUM> and a detector <NUM>. Pulse energy, average power, or a combination may be measured using detector <NUM>. The information can be used for feedback to the half-wave plate <NUM> for attenuation and to verify whether the shutter <NUM> is open or closed. In addition, the shutter <NUM> may have position sensors to provide a redundant state detection.

The beam passes through a beam conditioning stage <NUM>, in which beam parameters such as beam diameter, divergence, circularity, and astigmatism can be modified. In this illustrative example, the beam conditioning stage <NUM> includes a <NUM> element beam expanding telescope comprised of spherical optics <NUM> and <NUM> in order to achieve the intended beam size and collimation. Although not illustrated here, an anamorphic or other optical system can be used to achieve the desired beam parameters. The factors used to determine these beam parameters include the output beam parameters of the laser, the overall magnification of the system, and the desired numerical aperture (NA) at the treatment location. In addition, the optical system <NUM> can be used to image aperture <NUM> to a desired location (e.g. the center location between the <NUM>-axis scanning device <NUM> described below). In this way, the amount of light that makes it through the aperture <NUM> is assured to make it through the scanning system. Pickoff device <NUM> is then a reliable measure of the usable light.

After exiting conditioning stage <NUM>, beam <NUM> reflects off of fold mirrors <NUM>, <NUM>, & <NUM>. These mirrors can be adjustable for alignment purposes. The beam <NUM> is then incident upon beam combiner <NUM>. Beamcombiner <NUM> reflects the UF beam <NUM> (and transmits both the OCT <NUM> and aim <NUM> beams described below). For efficient beamcombiner operation, the angle of incidence is preferably kept below <NUM> degrees and the polarization where possible of the beams is fixed. For the UF beam <NUM>, the orientation of linear polarizer <NUM> provides fixed polarization.

Following the beam combiner <NUM>, the beam <NUM> continues onto the z-adjust or Z scan device <NUM>. In this illustrative example the z-adjust includes a Galilean telescope with two lens groups <NUM> and <NUM> (each lens group includes one or more lenses). Lens group <NUM> moves along the z-axis about the collimation position of the telescope. In this way, the focus position of the spot in the patient's eye <NUM> moves along the z-axis as indicated. In general there is a fixed linear relationship between the motion of lens <NUM> and the motion of the focus. In this case, the z-adjust telescope has an approximate <NUM>× beam expansion ratio and a <NUM>:<NUM> relationship of the movement of lens <NUM> to the movement of the focus. Alternatively, lens group <NUM> could be moved along the z-axis to actuate the z-adjust, and scan. The z-adjust is the z-scan device for treatment in the eye <NUM>. It can be controlled automatically and dynamically by the system and selected to be independent or to interplay with the X-Y scan device described next. Mirrors <NUM> and <NUM> can be used for aligning the optical axis with the axis of z-adjust device <NUM>.

After passing through the z-adjust device <NUM>, the beam <NUM> is directed to the x-y scan device by mirrors <NUM> & <NUM>. Mirrors <NUM> & <NUM> can be adjustable for alignment purposes. X-Y scanning is achieved by the scanning device <NUM> preferably using two mirrors <NUM> & <NUM> under the control of control electronics <NUM>, which rotate in orthogonal directions using motors, galvanometers, or any other well known optic moving device. Mirrors <NUM> & <NUM> are located near the telecentric position of the objective lens <NUM> and contact lens <NUM> combination described below. Tilting these mirrors <NUM>/<NUM> causes them to deflect beam <NUM>, causing lateral displacements in the plane of UF focus located in the patient's eye <NUM>. Objective lens <NUM> may be a complex multielement lens element, as shown, and represented by lenses <NUM>, <NUM>, and <NUM>. The complexity of the lens <NUM> will be dictated by the scan field size, the focused spot size, the available working distance on both the proximal and distal sides of objective <NUM>, as well as the amount of aberration control. An f-theta lens <NUM> of focal length <NUM> generating a spot size of <NUM>, over a field of <NUM>, with an input beam size of <NUM> diameter is an example. Alternatively, X-Y scanning by scanner <NUM> may be achieved by using one or more moveable optical elements (e.g. lenses, gratings) which also may be controlled by control electronics <NUM>, via input and output device <NUM>.

The aiming and treatment scan patterns can be automatically generated by the scanner <NUM> under the control of controller <NUM>. Such patterns may be comprised of a single spot of light, multiple spots of light, a continuous pattern of light, multiple continuous patterns of light, and/or any combination of these. In addition, the aiming pattern (using aim beam <NUM> described below) need not be identical to the treatment pattern (using light beam <NUM>), but preferably at least defines its boundaries in order to assure that the treatment light is delivered only within the desired target area for patient safety. This may be done, for example, by having the aiming pattern provide an outline of the intended treatment pattern. This way the spatial extent of the treatment pattern may be made known to the user, if not the exact locations of the individual spots themselves, and the scanning thus optimized for speed, efficiency and accuracy. The aiming pattern may also be made to be perceived as blinking in order to further enhance its visibility to the user.

An optional contact lens <NUM>, which can be any suitable ophthalmic lens, can be used to help further focus the optical beam <NUM> into the patient's eye <NUM> while helping to stabilize eye position. The positioning and character of optical beam <NUM> and/or the scan pattern the beam <NUM> forms on the eye <NUM> may be further controlled by use of an input device such as a joystick, or any other appropriate user input device (e.g. GUI <NUM>) to position the patient and/or the optical system.

The UF laser <NUM> and controller <NUM> can be set to target the surfaces of the targeted structures in the eye <NUM> and ensure that the beam <NUM> will be focused where appropriate and not unintentionally damage non-targeted tissue. Imaging modalities and techniques described herein, such as for example, Optical Coherence Tomography (OCT), Purkinje imaging, Scheimpflug imaging, or ultrasound may be used to determine the location and measure the thickness of the lens and lens capsule to provide greater precision to the laser focusing methods, including 2D and 3D patterning. Laser focusing may also be accomplished using one or more methods including direct observation of an aiming beam, Optical Coherence Tomography (OCT), Purkinje imaging, Scheimpflug imaging, ultrasound, or other known ophthalmic or medical imaging modalities and/or combinations thereof. In the embodiment of <FIG>, an OCT device <NUM> is described, although other modalities are within the scope of the present invention. An OCT scan of the eye will provide information about the axial location of the anterior and posterior lens capsule, the boundaries of the cataract nucleus, as well as the depth of the anterior chamber. This information is then loaded into the control electronics <NUM>, and used to program and control the subsequent laser-assisted surgical procedure. The information may also be used to determine a wide variety of parameters related to the procedure such as, for example, the upper and lower axial limits of the focal planes used for cutting the lens capsule and segmentation of the lens cortex and nucleus, and the thickness of the lens capsule among others.

The OCT device <NUM> in <FIG> includes a broadband or a swept light source <NUM> that is split by a fiber coupler <NUM> into a reference arm <NUM> and a sample arm <NUM>. The reference arm <NUM> includes a module <NUM> containing a reference reflection along with suitable dispersion and path length compensation. The sample arm <NUM> of the OCT device <NUM> has an output connector <NUM> that serves as an interface to the rest of the UF laser system. The return signals from both the reference and sample arms <NUM>, <NUM> are then directed by coupler <NUM> to a detection device <NUM>, which employs either time domain, frequency or single point detection techniques. In <FIG>, a frequency domain technique is used with an OCT wavelength of <NUM> and bandwidth of <NUM>.

Exiting connector <NUM>, the OCT beam <NUM> is collimated using lens <NUM>. The size of the collimated beam <NUM> is determined by the focal length of lens <NUM>. The size of the beam <NUM> is dictated by the desired NA at the focus in the eye and the magnification of the beam train leading to the eye <NUM>. Generally, OCT beam <NUM> does not require as high an NA as the UF beam <NUM> in the focal plane and therefore the OCT beam <NUM> is smaller in diameter than the UF beam <NUM> at the beamcombiner <NUM> location. Following collimating lens <NUM> is aperture <NUM> which further modifies the resultant NA of the OCT beam <NUM> at the eye. The diameter of aperture <NUM> is chosen to optimize OCT light incident on the target tissue and the strength of the return signal. Polarization control element <NUM>, which may be active or dynamic, is used to compensate for polarization state changes which may be induced by individual differences in corneal birefringence, for example. Mirrors <NUM> & <NUM> are then used to direct the OCT beam <NUM> towards beamcombiners <NUM> & <NUM>. Mirrors <NUM> & <NUM> may be adjustable for alignment purposes and in particular for overlaying of OCT beam <NUM> to UF beam <NUM> subsequent to beamcombiner <NUM>. Similarly, beamcombiner <NUM> is used to combine the OCT beam <NUM> with the aim beam <NUM> described below.

Once combined with the UF beam <NUM> subsequent to beamcombiner <NUM>, OCT beam <NUM> follows the same path as UF beam <NUM> through the rest of the system. In this way, OCT beam <NUM> is indicative of the location of UF beam <NUM>. OCT beam <NUM> passes through the z-scan <NUM> and x-y scan <NUM> devices then the objective lens <NUM>, contact lens <NUM> and on into the eye <NUM>. Reflections and scatter off of structures within the eye provide return beams that retrace back through the optical system, into connector <NUM>, through coupler <NUM>, and to OCT detector <NUM>. These return back reflections provide the OCT signals that are in turn interpreted by the system as to the location in X, Y Z of UF beam <NUM> focal location.

OCT device <NUM> works on the principle of measuring differences in optical path length between its reference and sample arms. Therefore, passing the OCT through z-adjust <NUM> does not extend the z-range of OCT system <NUM> because the optical path length does not change as a function of movement of <NUM>. OCT system <NUM> has an inherent z-range that is related to the detection scheme, and in the case of frequency domain detection it is specifically related to the spectrometer and the location of the reference arm <NUM>. In the case of OCT system <NUM> used in <FIG>, the z-range is approximately <NUM>-<NUM> in an aqueous environment. Extending this range to at least <NUM> involves the adjustment of the path length of the reference arm within OCT system <NUM>. Passing the OCT beam <NUM> in the sample arm through the z-scan of z-adjust <NUM> allows for optimization of the OCT signal strength. This is accomplished by focusing the OCT beam <NUM> onto the targeted structure while accommodating the extended optical path length by commensurately increasing the path within the reference arm <NUM> of OCT system <NUM>.

Because of the fundamental differences in the OCT measurement with respect to the UF focus device due to influences such as immersion index, refraction, and aberration, both chromatic and monochromatic, care must be taken in analyzing the OCT signal with respect to the UF beam focal location. A calibration or registration procedure as a function of X, Y Z should be conducted in order to match the OCT signal information to the UF focus location and also to the relate to absolute dimensional quantities.

Observation of an aim beam may also be used to assist the user to directing the UF laser focus. Additionally, an aim beam visible to the unaided eye in lieu of the infrared OCT and UF beams can be helpful with alignment provided the aim beam accurately represents the infrared beam parameters. An aim subsystem <NUM> is employed in the configuration shown in <FIG>. The aim beam <NUM> is generated by an aim beam light source <NUM>, such as a helium-neon laser operating at a wavelength of <NUM>. Alternatively a laser diode in the <NUM>-<NUM> range could be used. The advantage of using the helium neon <NUM> beam is its long coherence length, which would enable the use of the aim path as a laser unequal path interferometer (LUPI) to measure the optical quality of the beam train, for example.

Once the aim beam light source generates aim beam <NUM>, the aim beam <NUM> is collimated using lens <NUM>. The size of the collimated beam is determined by the focal length of lens <NUM>. The size of the aim beam <NUM> is dictated by the desired NA at the focus in the eye and the magnification of the beam train leading to the eye <NUM>. Generally, aim beam <NUM> should have close to the same NA as UF beam <NUM> in the focal plane and therefore aim beam <NUM> is of similar diameter to the UF beam at the beamcombiner <NUM> location. Because the aim beam is meant to stand-in for the UF beam <NUM> during system alignment to the target tissue of the eye, much of the aim path mimics the UF path as described previously. The aim beam <NUM> proceeds through a half-wave plate <NUM> and linear polarizer <NUM>. The polarization state of the aim beam <NUM> can be adjusted so that the desired amount of light passes through polarizer <NUM>. Elements <NUM> & <NUM> therefore act as a variable attenuator for the aim beam <NUM>. Additionally, the orientation of polarizer <NUM> determines the incident polarization state incident upon beamcombiners <NUM> and <NUM>, thereby fixing the polarization state and allowing for optimization of the beamcombiners' throughput. Of course, if a semiconductor laser is used as aim beam light source <NUM>, the drive current can be varied to adjust the optical power.

The aim beam <NUM> proceeds through a shutter <NUM> and aperture <NUM>. The system controlled shutter <NUM> provides on/off control of the aim beam <NUM>. The aperture <NUM> sets an outer useful diameter for the aim beam <NUM> and can be adjusted appropriately. A calibration procedure measuring the output of the aim beam <NUM> at the eye can be used to set the attenuation of aim beam <NUM> via control of polarizer <NUM>.

A device for imaging the target tissue on or within the eye <NUM> is shown schematically in <FIG> as imaging system <NUM>. Imaging system includes a camera <NUM> and an illumination light source <NUM> for creating an image of the target tissue. The imaging system <NUM> gathers images which may be used by the system controller <NUM> for providing pattern centering about or within a predefined structure. The illumination light source <NUM> for the viewing is generally broadband and incoherent. For example, light source <NUM> can include multiple LEDs as shown. The wavelength of the viewing light source <NUM> is preferably in the range of <NUM> to <NUM>, but can be anything that is accommodated by the beamcombiner <NUM>, which combines the viewing light with the beam path for UF beam <NUM> and aim beam <NUM> (beamcombiner <NUM> reflects the viewing wavelengths while transmitting the OCT and UF wavelengths). The beamcombiner <NUM> may partially transmit the aim wavelength so that the aim beam <NUM> can be visible to the viewing camera <NUM>. Optional polarization element <NUM> in front of light source <NUM> can be a linear polarizer, a quarter wave plate, a half-wave plate or any combination, and is used to optimize signal. A false color image as generated by the near infrared wavelength is acceptable.

The illumination light from light source <NUM> is directed down towards the eye using the same objective lens <NUM> and contact lens <NUM> as the UF and aim beam <NUM>, <NUM>. The light reflected and scattered off of various structures in the eye <NUM> are collected by the same lenses <NUM> & <NUM> and directed back towards beamcombiner <NUM>. There, the return light is directed back into the viewing path via beam combiner and mirror <NUM>, and on to camera <NUM>. Camera <NUM> can be, for example but not limited to, any silicon based detector array of the appropriately sized format. Video lens <NUM> forms an image onto the camera's detector array while optical elements <NUM> & <NUM> provide polarization control and wavelength filtering respectively. Aperture or iris <NUM> provides control of imaging NA and therefore depth of focus and depth of field. A small aperture provides the advantage of large depth of field which aids in the patient docking procedure. Alternatively, the illumination and camera paths can be switched. Furthermore, aim light source <NUM> can be made to emit in the infrared which would not directly visible, but could be captured and displayed using imaging system <NUM>.

Coarse adjust registration is usually needed so that when the contact lens <NUM> comes into contact with the cornea, the targeted structures are in the capture range of the X, Y scan of the system. Therefore a docking procedure is preferred, which preferably takes in account patient motion as the system approaches the contact condition (i.e. contact between the patient's eye <NUM> and the contact lens <NUM>. The viewing system <NUM> is configured so that the depth of focus is large enough such that the patient's eye <NUM> and other salient features may be seen before the contact lens <NUM> makes contact with eye <NUM>.

An alternative beamcombining configuration is shown in the alternate embodiment of <FIG>. For example, the passive beamcombiner <NUM> in <FIG> can be replaced with an active combiner <NUM> in <FIG>. The active beamcombiner <NUM> can be a moving or dynamically controlled element such as a galvanometric scanning mirror, as shown. Active combiner <NUM> changes it angular orientation in order to direct either the UF beam <NUM> or the combined aim and OCT beams <NUM>,<NUM> towards the scanner <NUM> and eventually eye <NUM> one at a time. The advantage of the active combining technique is that it avoids the difficulty of combining beams with similar wavelength ranges or polarization states using a passive beam combiner. This ability is traded off against the ability to have simultaneous beams in time and potentially less accuracy and precision due to positional tolerances of active beam combiner <NUM>.

Another alternate embodiment is shown in <FIG> which is similar to that of <FIG> but utilizes an alternate approach to OCT <NUM>. In <FIG>, OCT <NUM> is the same as OCT <NUM> in <FIG>, except that the reference arm <NUM> has been replaced by reference arm <NUM>. This free-space OCT reference arm <NUM> is realized by including beamsplitter <NUM> after lens <NUM>. The reference beam <NUM> then proceeds through polarization controlling element <NUM> and then onto the reference return module <NUM>. The reference return module <NUM> contains the appropriate dispersion and path length adjusting and compensating elements and generates an appropriate reference signal for interference with the sample signal. The sample arm of OCT <NUM> now originates subsequent to beamsplitter <NUM>. The potential advantages of this free space configuration include separate polarization control and maintenance of the reference and sample arms. The fiber based beam splitter <NUM> of OCT <NUM> can also be replaced by a fiber based circulator. Alternately, both OCT detector <NUM> and beamsplitter <NUM> might be moved together as opposed to reference arm <NUM>.

<FIG> shows another alternative embodiment for combining OCT beam <NUM> and UF beam <NUM>. In <FIG>, OCT <NUM> (which can include either of the configurations of OCT <NUM> or <NUM>) is configured such that its OCT beam <NUM> is coupled to UF beam <NUM> after the z-scan <NUM> using beamcombiner <NUM>. In this way, OCT beam <NUM> avoids using the z-adjust. This allows the OCT <NUM> to possibly be folded into the beam more easily and shortening the path length for more stable operation. This OCT configuration is at the expense of an optimized signal return strength as discussed with respect to <FIG>. There are many possibilities for the configuration of the OCT interferometer, including time and frequency domain approaches, single and dual beam methods, swept source, etc., as described in <CIT>;<CIT>;<CIT>; <CIT>; and <CIT>.

The device according to <FIG> is suitable for scanning <NUM>-dimensional patterns within lens <NUM> and cornea of the patient's eye <NUM> as described in <CIT>. Suitable patterns may be scanned using the laser to provide convenient splitting of lens <NUM> into segments that are easy to aspirate using existing technology and devices. Phacoemulsification is particularly well suited for this. Several such aspiration devices are commercially available and well known in the art.

Full Depth Laser Ophthalmic Surgical System using a Focal Point Extension Assembly. A first aspect of the present invention is directed to a Full Depth Laser Ophthalmic Surgical System. Most prior ophthalmic surgical systems are capable of only being used for a portion of the eye. These include the system described in detail later with reference to <FIG> (referred to hereinafter as "the existing laser ophthalmic surgical system"), which, when used with a corresponding patient interface device having a lens with a predefined optical power, can focus the laser beam in the region of the eye from the cornea to about the location of the lens posterior in the patient. Traditionally, other surgical systems, or cold steel methods, would be required for surgical interventions on the eye that were posterior to the lens, such as surgical interventions in the vitreous humor or the retina. According to an embodiment of the present invention, a removeable lens assembly has been made which extend the focal point of the existing laser ophthalmic surgical system so that the focal point of the laser system can be extended to portions of the eye posterior to the lens. This "focal point extension assembly" extends the focal point of the existing laser ophthalmic surgical system so that it can reach the vitreous humor and the retina. Thus, the existing laser ophthalmic surgical system coupled with the removeable focal point extension assembly create a "full depth" ophthalmic surgical system capable of performing surgical procedures along the entire length of the eye.

<FIG> and <FIG> illustrate an example of a modified patient interface that incorporates a focal point extension assembly, which is configured to extend the focal zone of the existing laser ophthalmic surgical system so that the focal zone can extend into the portions of the eye posterior to the lens capsule. The modified patient interface constitutes a removable focal zone extension adapter. The patient interface has a structure otherwise similar to an existing patient interface, such as that described in <CIT>, but includes an optical assembly (an adapter lens) that provides a less positive or morenegative optical power (as compared to the patient interface that is used with the same existing laser ophthalmic surgical system when treating the lens of the eye) to lengthen the focus. In the embodiment shown in <FIG>, the optical assembly is a doublet, i.e. two lenses joined together, without any air gap in between. This optical element provides better optical property and correspondingly better beam quality than a single lens. In addition, the optical assembly is optimized to achieve a reasonable beam quality over a reasonable field such as substantially the entire distance between the retina and the posterior lens capsule. In the embodiment shown in <FIG>, the upper end of the modified patient interface is configured to be attached to the objective lens assembly of the laser beam delivery system, and the lower end of the patient interface has a suction ring configured to engage the patient's eye via a vacuum force. The structures of the upper and lower ends of the patient interface are similar to those in the above- cited patent.

Note that the modified patient interface having the focal point extension assembly extends the focus distances of both the treatment laser and the imaging system beam, as well as that of the aim beam. A calibration or registration procedure as a function of X, Y Z should be conducted using the modified patient interface in order to match the OCT signal information to the treatment later focus location and also to the relate to absolute dimensional quantities. As described earlier, a calibration or registration procedure is conducted for the existing patient interface. The resulting calibration parameters for both calibrations, for example, in the form of different lookup tables, are stored in the controller.

Thus, when using the existing laser ophthalmic surgical system to perform surgical procedures in the vitreous humor of the eye, the modified patient interface having the focal point extension optical assembly is used, along with the corresponding calibration parameters; when using the same existing laser ophthalmic surgical system to perform surgical procedures in the anterior portions of the eye such as the cornea, the lens capsule and the lens, such as for cataract surgery, the existing patient interface without the focal point extension optical assembly is used, along with the corresponding calibration parameters. This allows the same laser system to be used to treat the full range of locations of the eye.

Full Depth Laser Ophthalmic Surgical System optimized for treating the vitreous. In an alternative embodiment of this aspect of the invention, the optical system is configured to scan the laser beam and the imaging beam (e.g., the OCT light) in the Z (depth) direction within the full range of the eye, from the cornea to the retina; the optical parameters (such as the NA and the Strehl ratio) are optimized for scanning the laser beam and the imaging beam in the entire vitreous volume. <FIG> schematically illustrates an example of such an optical system. As shown in <FIG>, the laser beam and the imaging beam, having been directed into the same optical path by a beam combiner (not shown in <FIG>), passes through a Z-scanner, which includes a lens L1 COM that is moveable along the optical axis. In a preferred embodiment, the lens L1 COM is a negative spherical lens having an effective focal length (EFL) of -<NUM> and a back focal length (BFL) of -<NUM> (for <NUM> light). The laser beam and the imaging beam then pass through another lens L2 COM, and are aligned by two folding mirrors M3 COM and M4 COM and delivered to the XY-scanner which comprises two scanning mirrors GX and GY. The laser beam and imaging beam output from the XY-scanner are transmitted through a beam splitter BC3, and then focused by an objective lens OBJ. The reflected OCT light from the eye travels in the opposite direction through the objective lens OBJ, the beam splitter BC3, and the scanner back to the OCT assembly for detection.

In a preferred embodiment, the objective lens OBJ is a lens group including five lenses OBJ-<NUM> to OBJ-<NUM>, with lenses OBJ-<NUM> to OBJ-<NUM> being doublet lenses and the downstream-most lens OBJ-<NUM> being a meniscus lens. Each of OBJ-<NUM> and OBJ-<NUM> has a focal length of <NUM> and each of OBJ-<NUM> and OBJ-<NUM> has a focal length of <NUM>, and the meniscus lens OBJ-<NUM> has an EFL of <NUM> and a BFL of <NUM> (for <NUM> light).

The optical system is intended to be used with patient interface coupled to the downstream end of the objective lens group OBJ during ophthalmic procedures. The patient interface includes a focusing lens WCL. In a preferred embodiment, the focusing lens is plano-convex having a spherical surface at the proximate end (closer to the objective lens) and a flat surface at the distal end (closer to the eye), with an EFL of <NUM> and a BFL of <NUM> (for <NUM> light). The parameters of the meniscus lens OBJ-<NUM> described in the preceding paragraph are optimized for a patient interface having the plano-convex focusing lens WCL described above. In an alternative embodiment, the focusing lens WCL of the patient interface is a double convex lens, and has a convex surface at the distal end. The convex surface facing the eye can help reduce bubbles in the fluid between the focusing lens WCL and the surface of the eye. In such an embodiment, the parameters of the meniscus lens OBJ-<NUM> will be modified accordingly to optimize it for the double convex lens WCL. The doublet lenses and the meniscus lens of the objective lens group OBJ, together with the lens WCL of the patient interface, deliver the focal spots of the laser beam and the imaging beam to the vitreous volume with minimal aberrations.

The optical system is also configured to direct a fixation light and an illumination light to the eye and direct light reflected or otherwise emitted from the eye back to a video camera for imaging. The fixation light from a fixation light source is collimated by a lens L-FIX, then reflected by a beam splitter Fixation BS to the beam splitter BC3, which in turn reflects the fixation light to the objective lens OBJ. The illumination light, emitted by an LED illumination ring, is transmitted through the beam splitter Fixation BS and then reflected by the beam splitter BC3 to the objective lens OBJ. The objective lens OBJ focuses the fixation light and the illumination light into the eye. An important feature of the camera illumination light source placement and configuration is that it can be pupil matched to the pupil of the patient's eye in a manner similar to locating the image of the X & Y scan mirrors near the pupil (described later). That is, the objective lens group OBJ and the focusing lens of WCL of the patient interface preferably form an image of the illumination source ear the pupil of the patient's eye, and substantially centered with the fixation light, as are images of the X & Y scan mirrors. This configuration allows for illumination of a larger field in the full depth range of the eye. The light reflected or otherwise emitted from the eye travels backwards through the objective lens OBJ and is reflected by the beam splitter BC3. After transmitting through the beam splitter Fixation BS, the light from the eye passes through an aperture AS VIEW and is focused by additional lenses, including a tunable lens and three doublet lenses L1 VIEW to L3 VIEW in a preferred embodiment, onto the camera detector CAM which generates an image of the eye. The tunable lens can be either tunable by moving the lens, or by adjusting the optical power of the lens.

The Z scanner (L1 COM), the XY scanner (GX and GY), and the tunable lens for the video camera CAM are electrically coupled to the controller of the laser ophthalmic surgical system, and controlled by the controller to scan of the laser beam and the imaging beam (e.g. the OCT beam) within the eye and to image the eye with the video camera.

This optical system is optimized to focus and scan the laser beam and the imaging beam in substantially the entire depth of vitreous humor of the eye from the posterior lens capsule to the retina. In actual surgery, it is desired to maintain safe zones (i.e. zones free of laser focal spot scanning) of a depth of about <NUM> after the lens (the posterior lens capsule) and a depth of about <NUM> before the retina. <FIG> schematically illustrates the focusing of the laser beam and the imaging beam in a depth range between the two safe zones, with a viewing angle of <NUM> degrees for the X scan and <NUM> degrees for the Y scan. It should be noted that the light exiting the focusing lens WCL of the patient interface is further refracted by the optical elements of the eye, including the cornea and the lens of the eye (it should be noted, however, that due to the fluid bath in the patient interface that contacts the cornea, focusing by the cornea tends to be minimal); the optical system design takes this into consideration. Moreover, when the patient interface is one where the distal surface of the focusing lens WCL does not directly contact the cornea of the eye but uses a fluid bath located between the WCL lens and the cornea, the relative distance of the patient's eye from the WCL lens surface is a function of the design of the patient interface such as the size of the fluid chamber. Thus, the optical system is designed such that the focusing range covers the entire desirable range within the eye after taking into consideration all of the factors discussed above. In one particular example, the optical system is designed such that without the eye, the focal spot of the laser beam can be scanned within a depth range of <NUM> to <NUM> in water beyond the distal surface of the focusing lens WCL of the patient interface.

As shown in <FIG>, the optical system focuses the laser beam and imaging beam within the vitreous volume with wide viewing angles in both the X and Y scanning directions. This is achieved by the objective lens group OBJ and the focusing lens of WCL of the patient interface, which together function as relay optics to form respective images of the scanning mirrors GX and GY at locations near or at the pupil plane of the eye, as shown in <FIG>. In a conventional optical beam scanning system that is designed to scan the focal spot of the laser to treat the anterior segments of the eye including the cornea and the lens, as illustrated in <FIG>, the depth range of the focal spot is typically limited to the lens posterior capsule. When the conventional optical beam scanning system is used with a different patient interface lens that extends the depth range of the laser focal spot to the vitreous and the retina, as shown in <FIG>, due to the fact that in such conventional optical beam scanning system the pivot points of the scanning mirrors GX and GY are located far away from the pupil of the eye, the pupil diameter limits the transverse scanning range of the beam. In this situation, the volume that can be reached by the laser focal spot is approximately a tapered cylinder that has a maximum diameter less than the pupil diameter. In embodiments of the present invention, as shown in <FIG>, due to the relay optics, the scanned laser beam emanates from the locations of the images of the scanning mirrors GX and GY, which are located near or at the pupil plane of the eye. This makes the scanning angle in the eye much wider, and the beam can reach a volume in the vitreous that has a maximum diameter larger than the pupil diameter. In a preferred embodiment, the images of the scanner mirrors are located about <NUM> to <NUM> from the patient's iris plane. When the system is used with a patient interface that employs a fluid bath between the cornea and the distal surface of the focusing lens WCL of the patient interface, the optical design takes into account the approximate distance between the WCL lens surface and the cornea in order to form the images of the scanner mirrors at the above-described desired location relative to the patient's iris plane. In one particular example, the optical system is designed such that the images of the scanner mirrors are located about <NUM> to <NUM> from the distal surface of the focusing lens WCL of the patient interface. It should be noted that the same considerations apply to the design of the optical system to form an image of the illumination light source near the iris plane as discussed earlier. In operation, the depth location of the iris from the distal surface of the focusing lens of the patient interface may be measured, and the objective lens is controlled to form the image of the scanning mirrors, as well as the illumination light as mentioned earlier, at the pupil.

By using the tunable lens before the video camera, the optical system is configured to focus light emitted anywhere within the entire depth of the eye onto the video camera CAM, so as to image different planes of the eye from the iris to the retina. Again, the light emitted from the vitreous of the eye will be refracted by optical elements of the eye such as the lens (and to a much lesser extent the cornea), and the design of the optical system takes this into consideration. Without the eye, the optical system can image objects within a depth range of <NUM> to <NUM> in water beyond the distal surface of the focusing lens WCL of the patient interface. In a preferred embodiment, the tunable lens has a tunable range of -10D to +10D. As mentioned earlier, the tunable lens can be either tunable by adjusting the optical power of the lens, or by moving an imaging lens having a fixed optical power. Such a fixed power lens may be mounted on a mechanically moveable stage (moved by a motor or an actuator), and the movement of the stage is controlled by the controller to change the imaging location of the video camera. The movement range of the stage is configured to allow the video camera to image different planes of the eye from the iris to the retina.

The illumination light source is ring shaped, and located near the aperture AS VIEW of the video camera and coaxially with the optical axis of the video camera. This configuration allows the illumination light to illuminate a wide field and the entire depth range from the iris to the retina to aid camera imaging.

While the optical system of this alternative embodiment is described in considerable detail, variations of the optical design are also within the scope of the invention, so long as they can scan the laser beam and the imaging beam to the desired depth range within the eye, direct the illumination light to a wide field and the entire depth range of the eye, and focus the light emitted from the entire depth range eye to the video camera. For example, different combinations of imaging lenses to reach the same or different F numbers or magnifications may be used.

Systems and methods for calibrating the Ophthalmic Surgical Systems for procedures performed posterior to the lens. A problem associated with using laser systems, such as that according to <FIG> in the posterior portion of the eye is that it is very difficult to properly calibrate the surgical system for use in the posterior portion. This is due to the fact that there is an unknown optical element, i.e. the patient's lens (with an unknown refractive power), between the laser source and location of the focal zone in the patient's eye. As a result of this unknown optical element, there can be a significant difference between the apparent location of the focal zone and the actual location of the focal zone. This can be dangerous as the focal zone could unknowingly be in dangerous location. While using average values for eye physiology may be satisfactory for much of the population, there could be a significant risk for persons whose eyes are outside normal physiological averages. One method is to use structural data of the lens of the patient's eye, obtained using the OCT system, to determine the individual active optical power of the lens, in order to accurately place the laser focus within the vitreous volume. Alternatively, fast and efficient calibration methods can be done for each individual patient as described below.

One aspect of the present description is directed to systems and methods for calibration of laser surgical systems for use in portions of the eye that are posterior to the human lens. A system and method for the calibration is based on the use of the imaging subsystem's optical coherence tomographer. In essence, the focal zone of the optical coherence tomographer beam is used as a proxy for the focal zone of the femtosecond laser source. This method may be implemented using either the existing laser ophthalmic surgical system with the modified patient interface shown in <FIG>, or a laser ophthalmic surgical system having the optical system shown in <FIG>.

The method can be summarized by the following steps:.

More specifically, the focal zone of OCT beam is scanned in the depth direction (using the Z scan device of the laser system) in regions of the eye posterior to the posterior lens capsule, and the intensity of the returned (i.e. reflected) OCT beam is measured as a function of depth. Peaks of the intensity function are identified to determine the depth location of the retina, the posterior lens capsule, and optionally other structures of the eye. The safety limiting depths are set based on the depths of the retina and posterior lens capsule, e.g., at <NUM> anterior of the retina and <NUM> posterior of the posterior lens capsule. Note that this safety distance is a function of pulse energy. Other appropriate safety values may be determined based on experiments. The parameters of the Z scan device that correspond to these various depths are determined, which can then be used by the controller to scan the focal zone of the treatment laser beam for treating the vitreous humor, e.g. to treat floaters or liquify the vitreous humor as described herein. The scan of the treatment beam is performed within the volume bound by the safety limitation positions described above. This method allows one to safely set the location of the laser beam taking into account the individual refractive power of each person's eye and at the same time allows one to identify a safe surgical zone for the ophthalmic surgery in portions of the eye posterior to the lens.

Treatment Methods and Systems Carrying out the Method: Liquification of the Vitreous Humor. Surgical procedures sometimes require removal of the vitreous. However, this can be difficult because of the gelatin like structure of the vitreous. In addition, the vitreous may be adhered to various delicate structures of the eye (e.g. the retina) and it can be difficult to remove the vitreous without disturbing other structures. An in vitro study in rabbits performed by the instant inventors showed the liquification of the vitreous humor with the femtosecond laser surgical system described above can be performed safely. Liquification of the vitreous may be indicated in surgical procedures where removal of the vitreous is required.

To liquify the vitreous humor, either the existing laser ophthalmic surgical system with the modified patient interface shown in <FIG>, or a laser ophthalmic surgical system having the optical system shown in <FIG>, is used. The retina and the lens posterior capsule are located, for example using the process described above. The imaging system such as the OCT system may be used to image the structure of the vitreous humor to identify the outer boundaries for liquification. The outer boundaries may be determined by the human operator or by the controller based on the images. A treatment plan can then be defined with the assistance of the controller. For example, a scan pattern of the laser focal spot within the volume defined by the boundaries can be programmed using the controller. The treatment plan is then executed to scan the focal spot of the laser beam according to the scan pattern, to thereby perform a volumetric treatment of the vitreous humor to enable liquification. The following laser beam parameters may be used in preferred embodiments: Wavelength: preferably <NUM> - <NUM>, and more preferably <NUM>; pulse width: preferably <NUM> - <NUM> fs, and more preferably <NUM>-<NUM> fs; pulse energy: preferably <NUM> - <NUM>µJ; repetition rate: preferably <NUM> -<NUM>; total energy: preferably less than <NUM> J. In an animal test using rabbit eyes, the following laser parameters were used to successfully liquify the vitreous humor: Wavelength: <NUM>; pulse width: about <NUM> fs; pulse energy: about <NUM>µJ; repetition rate: <NUM>-<NUM>; total energy: <NUM>-<NUM> J.

Treatment Methods and Systems Carrying out the Method: Treatment of floaters and other vitreous treatment. In addition to vitreous liquification, either the existing laser ophthalmic surgical system with the modified patient interface shown in <FIG>, or a laser ophthalmic surgical system having the optical system shown in <FIG>, may be used to treat vitreous floaters. The OCT and/or the video camera system may be used to locate the vitreous floaters, and treatment plan is determined and executed to specifically target the floaters for treatment. For example, the laser focal spot may be scanned in a volume that includes the floaters to destroy or remove the floaters. Further, the system may be used to pre-treat the vitreous volume before vitrectomy to simplify the surgical procedure and reduce possible complications.

In this disclosure, as is customary in the field of ophthalmic surgery, the term "depth" refers to a direction of the eye that extends from cornea to the retina; it corresponds to the Z direction of the laser ophthalmic surgical system as described above. The terms "anterior" and "posterior" are relative terms and refer to locations closer to the cornea and closer to the retina, respectively.

The use of the terms "a" and "an" and "the" and similar referents in the context of describing the invention (especially in the context of the following claims) are to be construed to cover both the singular and the plural, unless otherwise indicated herein or clearly contradicted by context. The term "connected" is to be construed as partly or wholly contained within, attached to, or joined together, even if there is something intervening. The use of any and all examples, or exemplary language (e.g., "such as") provided herein, is intended merely to better illuminate embodiments of the invention and does not pose a limitation on the scope of the invention unless otherwise claimed.

Claim 1:
A full depth ophthalmic surgical system for performing surgery on an eye of a subject, comprising:
a femtosecond laser source (<NUM>) configured to produce a femtosecond pulsed laser beam;
an imaging assembly (<NUM>) configured to emit an imaging beam;
a scanning assembly (<NUM>) including a Z scanner and an XY scanner, configured to scan a focal spot of the laser beam and the imaging beam within the eye in a depth direction and two transverse directions, respectively;
an illumination light source (<NUM>) configured to emit an illumination light;
a video camera assembly;
an objective lens assembly (OBJ) configured to focus the laser beam and the imaging beam, wherein the objective lens assembly includes four doublet lenses and a meniscus lens;
a patient interface configured to be coupled to the objective lens assembly and to engage the eye, the patient interface including a lens having a predefined optical power, wherein the lens is a double-convex focusing lens; and
optical components including at least one beam splitter (BC3), configured to direct the laser beam and the imaging beam output by the scanning assembly and the illumination light to the objective lens assembly, and to direct light emitted from within the eye, which has passed through the objective lens assembly, to the video camera assembly;
wherein the scanning assembly (<NUM>) , the objective lens assembly (OBJ) and the lens of the patient interface are configured to form a focal spot of the laser beam at any depth within a range of <NUM> to <NUM> in water beyond a distal surface of the lens of the patient interface.