WO1999034759A1

WO1999034759A1 - Method and apparatus for removing tissue with mid-infrared laser

Info

Publication number: WO1999034759A1
Application number: PCT/US1998/027682
Authority: WO
Inventors: William B. Telfair; Hannah J. Hoffman; Paul R. Yoder, Jr.
Original assignee: Irvision, Inc.
Priority date: 1997-12-31
Filing date: 1998-12-28
Publication date: 1999-07-15
Also published as: AU2096499A

Abstract

A medical apparatus and method for removing corneal tissue from an eye of a patient is disclosed. As described in one aspect of the disclosure, the apparatus comprises a laser source that produces short pulses of mid-infrared radiation, wherein the infrared radiation has a wavelength approximately corresponding to a corneal absorption peak. The apparatus also includes a scanner-deflection means to direct the pulsed radiation across an area of the corneal tissue in a predefined pattern to remove portions of the corneal tissue primarily by photospallation so that collateral thermal damage on the tissue is less than 1 micron thick. The apparatus and method may include an eye tracker to stabilize the tissue removal pattern during the procedure.

Description

METHOD AND APPARATUS FOR REMOVING TISSUE WITH MID-INFRARED LASER

REFERENCE TO RELATED APPLICATION

This application is a continuation-in-part of United States Patent Application Serial No. 08/549,385, which was filed on October 27, 1995.

BACKGROUND OF THE INVENTION

The present invention relates to laser surgical techniques for modifying tissue such as the corneal surface of the eye, and more particularly, to laser surgical techniques, such as laser-assisted in-situ Keratomiluesis (LASIK) and photorefractive keratectomy (PRK), which direct reshaping of the cornea by means of selective volumetric removal of corneal tissue.

In recent years, numerous corneal sculpting techniques and related apparatus have been disclosed for correcting visual deficiencies such as near- sightedness, far-sightedness, and astigmatism. In addition, corneal sculpting techniques have also been utilized for therapeutic interventions in a number of pathologic conditions involving the cornea. For example, U.S. Patent Nos. 4,665,913, 4,732,148, and 4,669,466 to L'Esperance, and U.S. Patent No. 5,108,388 to Trokel, describe methods for achieving optical correction through reshaping of the anterior corneal surface. In addition, a number of instruments for affecting refractive surgery have recently become commercially available, such as the STAR model from Visx of Santa Clara, CA and the Omni-Med model from Summit of Watertown, MA.

These commercial devices, as well as most corneal sculpting methods and devices which have been disclosed and manufactured to date, utilize ultraviolet (UN) radiation with a wavelength which is preferably less than 200 nanometers. For example, many of these devices utilize an Argon Fluoride excimer laser operating at 193 nm. Generally, radiation at such short ultraviolet wavelengths is characterized by high photon energy, namely, greater than 6 eV, which, upon impact with tissue, causes molecular decomposition, i.e., the direct breaking of intramolecular bonds. Such a technique is a photochemical one. It has the advantage of producing minimal collateral thermal damage in cells adjacent to the surgical site, since the broken molecules generally leave behind only small volatile fragments which evaporate without heating the underlying substrate. Furthermore, the depth of decomposition for each laser pulse is typically very small, i.e., less than 1 micron, thus achieving accurate tissue removal with minimal risk of damage to the underlying structures from UV radiation.

In view of this small depth of penetration, coupled with the need to remove sufficient depth of tissue while minimizing the overall time for the surgical procedure, the majority of corneal sculpting techniques utilizing the excimer laser employ "wide area ablation". Generally, wide area ablation utilizes a laser beam with a relatively large spot size to successively remove thin layers of corneal tissue. The spot size is generally of a size sufficient to cover the entire optical zone of the cornea, namely, a region of 5 to 7 millimeters in diameter. Consequently, to assure required flux densities on the cornea, relatively high energy output UV lasers are typically required. Since the delivery systems of these systems usually transmit only 15 to 30 % of the laser output energy to the eye, the energy required from the laser to deliver around 150 mJ/cm^ to an area of 5 to 7 mm on the eye is 100 to 400 mJ per pulse. Such lasers, however, tend to be prohibitively large and expensive systems.

Furthermore, efficacious wide area ablation requires that the projected beam be spatially homogenous and uniform to achieve the desired smooth corneal profiles. Accordingly, additional beam shaping devices, such as rotating prisms, mirrors, or spatial integrators, must be employed within the excimer beam delivery systems. For a more detailed discussion of beam shaping and delivery systems, see, for example, U.S. Patent No. 4, 911, 711 to Telfair et al. Of course, such a multiplicity of optical elements contributes to overall transmission loss, while adding substantial optical complexity, cost, and maintenance requirements to the system.

Alternative techniques based on utilization of a scanning UV laser beam have been proposed to achieve controlled and localized ablation of selected corneal regions of the cornea. In the scanning approach, a relatively small laser spot is scanned rapidly across the cornea in a predefined pattern to accumulatively shape the surface into the desired geometry. For a more detailed discussion of laser scanning techniques employing excimer lasers, see U.S. Patent No. 4,665,913 to L'Esperance; Lin, J.T., "Mini-Excimer Laser Corneal Reshaping Using a Scanning Device," SPIE Proceedings, Vol. 2131, Medical Lasers & Systems III (1994); and U.S. Patent No. 5,520,679 issued to Lin. A scanning approach may offer a number of advantages, including lower power and energy requirements, added flexibility for refractive corrections and smooth ablation profiles, without the need for spatially uniform output beam profiles. For example, a laser scanning technique allows a tapered optical treatment zone to be achieved, which may have advantages for the correction of high myopia, for performing therapeutic tissue removal and for treating areas up to 9 millimeters in diameter which may be required for the correction of hyperopia.

While laser surgical techniques based on the excimer laser have proved beneficial for many applications, such techniques suffer from a number of limitations, which, if overcome, could significantly advance the utility of optical laser surgery. For example, techniques based on excimer lasers utilize toxic gases as the laser medium, suffer from persistent reliability problems, require lossy optics in the delivery systems, and suffer from the possibility that the UV radiation is potentially mutagenic through secondary fluorescence, which may cause undesirable long term side effects to the unexposed tissues of the eye.

Accordingly, alternatives to the excimer laser have been suggested in recent years which involve frequency-shifted radiation from a solid state laser. Current limitations of nonlinear elements used as frequency-shifting devices, however, place a lower limit of approximately 205 nm on the available wavelengths of such lasers, which may be too close to the mutagenic range, which exhibits a peak at 250 nm. In addition, multiply-shifted laser devices also face certain difficulties in providing the requisite energy outputs and are fairly complex and cumbersome, leading again to potential laser reliability problems, as well as added cost and maintenance.

More recently, ablation at mid-infrared wavelengths using, in particular, radiation around 2.94 μm, has been suggested as an alternative to the excimer laser for performing corneal refractive surgery. With an absorption coefficient of approximately 13,000 cm ' , radiation at this wavelength corresponds to the peak of the absorption curve of water, the main constituent of the cornea. Consequently, 2.94 micron radiation can produce very small regions of impact and has the potential to ablate tissue selectively with minimal collateral thermal damage, similar to what is produced with the excimer - i.e., less than one micron. The promise of such an alternative system is that it would capitalize on the fact that infrared radiation can be produced with all- solid-state technology which would provide easier handling, be cheaper, more compact and have better reliability features, while eliminating the potential of any safety concerns due to toxic gases or mutagenic side effects associated with deep UV wavelengths.

Contrary to the photoablation mechanism associated with the excimer laser, i.e., photochemical decomposition, which is due to energy absorption in molecular bonds, ablation in the infrared wavelength range is generally attributed to photothermal evaporation of water molecules contained in the tissue. This process has inherently a larger effect than photodecomposition, allowing for removal of up to several microns of tissue per pulse, thereby resulting in faster surgical operations, but also with a larger thermal damage zone. A system for performing PRK based on a photovaporization process has been suggested, for example, by T. Seiler and J. Wollensak, in "Fundamental Mode Photoablation of the Cornea for Myopic Correction", Lasers and Light in Ophthalmology, 5, 4, 199-203 (1993). Another system has been described by Cozean et al. in PCT Application No. 93/14817, which relies on a sculpting filter to control the amount of tissue removal using a pulsed 2.94 μm wavelength Er:YAG laser. Both of these systems utilize free-running lasers where pulse durations extend to hundreds of microseconds. While tissue is rapidly removed, larger thermal damages zones, of up to 50 microns, are associated with these lasers due to the high energies involved (on the order of a joule). Such extensive damage zones are considered detrimental in corneal surgery because they are typically accompanied by haze, regression of refraction correction, loss of visual acuity and other deleterious healing side effects.

Recently, it has been recognized that mid-infrared lasers emitting shorter pulses (Q-switched lasers) cause less thermal damage than lasers with longer pulsewidths (see, for example, Q. Ren, R. A. Hill & M. W. Berns in "Laser refractive surgery: a review and current status", Opt. Eng., Vol. 34, No. 3, pp. 642-660, 1995). However, even with pulses on the order of hundreds of nanoseconds (as compared to hundreds of microseconds for previous studies) the collateral damage zone still extends to over 20 μm (see, for example, Jing-cai Lian & Kang-sun Wang in SPIE., 2393, pp. 160-166, 1995). While this amount of thermal damage is less than that caused by free running lasers, such an extent of damage is still considered too much for PRK or LASD , due to potential regression and unfavorable healing characteristics This puts all prior art mid-infrared lasers at a disadvantage when compared with excimer lasers for corneal ablation.

Two other prior studies used mid-infrared lasers with even shorter pulsewidths, though still relying on a photothermal evaporation technique to achieve ablation. More specifically, Seiler et. al., in "The potential of an infrared hydrogen fluoride (HF) laser (3.0 microns) for corneal surgery", Lasers in Ophthalmology, vol. 1, no. 1, pp. 49-60 (1986), reported use of the HF laser operating near 3 microns with pulsewidths of 50 nanoseconds and damage layers of 5 to 20 microns. However, it is likely that the multi-wavelength character of the HF laser increased thermal damage by effectively lowering the overall absorption coefficient of the tissue. In the second study, Stern et. al., in "Infrared Laser Surgery of the Cornea", Ophthalmology, vol. 95, no. 10, pp. 1434-1441 (1988), reported use of a Raman-shifted Nd:YAG at 2.92 microns with a pulsewidth of 8 nanoseconds. They found the ablation threshold to be about 250 mJ/cm^ with collateral thermal damage layers between 2 and 4 microns wide on the average.

Thus, while the prior literature refers to the use of shorter pulses in the mid-infrared region, such literature invariably teaches ablation using a photothermal technique. Such photothermal techniques cause damage zones generally greater than 2 microns, which result in tissue effects that do not approach those achieved with the excimer laser. Moreover, typically, the only criteria cited to lessen thermal damage is that the pulse duration be shorter than the characteristic thermal relaxation time, which is about 2 microseconds in corneal tissue. The underlying assumption is that at pulse durations well below this limit, the thermal damage is less extensive because the tissue evaporation process is faster and less explosive. Such guidance is, however, very vague in that it covers an overly large range of laser pulse parameters, resulting in tissue effects that are not very predictable and in no case are shown to approach the submicron effects routinely obtained with excimer lasers.

As an example, ablation thresholds ranging from 200 to 600 mJ/cm² were recited by J. T. Lin in US Patent No. 5,520,679 as necessary for corneal sculpting application at mid-infrared wavelengths, including at 2.94 μm. This patent discloses that one suitable source for such sculpting is an Er:YAG laser with parameters including output energy of over 50 mJ (and up to 500 mJ) and with pulse durations that are between 50 and 400 ns. Lin also discloses use of an Optical Parametric Oscillator (OPO) as an alternative mid-infrared source with a short pulse width (1 to 40 nanoseconds), a wide wavelength range (2.6 - 3.2 μm), and energy outputs up to 10 mJ to perform the same procedure as the Er: NAG laser.

It is, therefore, clearly seen that Lin's patent covers an overly wide range of parameters. These parameters appear to be the parameters over which the lasers will operate, but pay no attention as to whether the system can perform the appropriate surgery on the tissue. The only parameters directly relating to the impact on the tissue are the ablation threshold limits disclosed as 200 to 600 mJ/cm², which is consistent with a purely photothermal ablation process.

In particular, for the case of the ErrYAG laser, the range of parameters taught in the '659 patent lead to a minimum fluence of 1600 mJ/cm² (using the largest spot size with the lowest energy) and a maximum fluence of almost 1,600,000 mJ/cm^ (using the smallest spot size with the largest energy). Even assuming substantial losses in the delivery system, fluences on the eye are likely to exceed 600 mJ/cm², corresponding to levels which are known to involve photothermal processes with damage zones well in excess of 4-5 microns based on the prior art literature. Lin also does not teach the possibility of a process with an ablation threshold lower than 200 mJ/cm² or recognize any benefits of operating at such lower fluences upon the surgical procedures contemplated.

In the case of the OPO laser disclosed in the '659 patent, the fluence range from the stated laser parameters is a minimum of 3.2 mJ/cm² (which is far below the ablation threshold and would therefore have no impact on the tissue) and a maximum of almost 130,000 mJ/cm² (which is significantly above fluences used by previous researchers of laser tissue ablation and would cause significant collateral thermal damage. Clearly, the '659 patent only contemplates configuring a laser to ablate tissue using a photothermal evaporation process.

Thus, as described above, all prior art techniques for delivering and controlling a mid-infrared laser beam each are subject to one shortcoming in particular, namely, the potential for excess thermal damage to unablated regions of the eye. As is apparent from the above discussion, a need exists for an improved method and apparatus for surgically treating corneal tissue based on the controlled removal of tissue in which there is no excess of such thermal damage.

SUMMARY OF THE INVENTION

A first aspect of the present invention is directed to a method for removing corneal tissue from an eye of a patient. The method comprises the steps of generating a pulsed beam of laser radiation for ablating the corneal tissue. The beam comprises mid-infrared radiation at a wavelength approximately corresponding to a corneal absorption peak and has a pulse duration of at least about 1 nanosecond and a clinical fluence in the range of 20-200 mJ/cm^. The method also includes the step of scanning the beam across an area of the corneal tissue in a predefined pattern to remove portions of the corneal tissue primarily by photospallation.

A second aspect of this invention is directed to a method for removing corneal tissue from an eye of a patient. The method comprises the steps of generating a pulsed beam of laser radiation for ablating the corneal tissue. The beam comprises mid-infrared radiation at a wavelength approximately corresponding to a corneal absorption peak and has a pulse duration of at least about 1 nanosecond and an ablation threshold in the range of 10-75 mJ/cm^. The method also includes the step of scanning the beam across an area of the corneal tissue in a predefined pattern to remove portions of the corneal tissue primarily by photospallation.

A third aspect of the present invention is directed to a medical apparatus for removing corneal tissue from an eye of a patient. The apparatus comprises a laser source that produces pulses of mid-infrared radiation. The infrared radiation has a wavelength approximately corresponding to a corneal absorption peak. The pulses have a duration of at least about 1 nanosecond and a clinical fluence in the range of 20- 200 mJ/cm^. The apparatus also includes a scanner-deflection means to direct the pulsed radiation across an area of the corneal tissue in a predefined pattern to remove portions of the corneal tissue primarily by photospallation.

A fourth aspect of this invention is directed to a medical apparatus for removing corneal tissue from an eye of a patient. The apparatus comprises a laser source that produces pulses of mid-infrared radiation, wherein the infrared radiation has a wavelength approximately corresponding to a corneal absorption peak. The pulses have a duration of at least about 1 nanosecond resulting in ablation thresholds for the tissue in the range of 10-75 mJ/cm^. The apparatus also includes a scanner- deflection means to direct the pulsed radiation across an area of the corneal tissue in a predefined pattern to remove portions of the corneal tissue primarily by photospallation.

A fifth aspect of the present invention is directed to a medical apparatus for removing corneal tissue from an eye of a patient. The apparatus comprises a laser source that produces pulses of mid-infrared radiation, wherein the infrared radiation has a wavelength approximately corresponding to a corneal absorption peak. The pulses have a duration of at least about 1 nanosecond, an ablation threshold in the range of 10-75 mJ/cm^, a clinical fluence in the range of 20-200 mJ/cm^, and a repetition rate of at least 10 Hz. The apparatus also includes a scanner-deflection means to direct the pulsed radiation across an area of the corneal tissue in a predefined pattern to remove portions of the corneal tissue primarily by photospallation so that collateral thermal damage on the tissue is less than 1 micron thick.

A sixth aspect of this invention is directed to a method for removing corneal tissue from an eye of a patient. The method comprises the steps of generating a pulsed beam of laser radiation for ablating the corneal tissue, wherein the beam comprises mid-infrared radiation at a wavelength approximately corresponding to a corneal absorption peak. The beam has a pulse duration of at least about 1 nanosecond, an ablation threshold in the range of 10-75 mJ/cm^, a clinical fluence in the range of 20-200 mJ/cm^, and a repetition rate of at least 10 Hz. The method also includes the step of scanning the beam across an area of the corneal tissue in a predefined pattern to remove portions of the corneal tissue primarily by photospallation so that collateral thermal damage on the tissue is less than 1 micron thick.

A seventh aspect of the present invention is directed to a method for removing corneal tissue from an eye of a patient. The method comprises the steps of: generating a pulsed beam of laser radiation for ablating the corneal tissue, wherein the beam comprises mid-infrared radiation at a wavelength approximately corresponding to a corneal absorption peak. The method also includes the step of scanning the beam across an area of the corneal tissue in a predefined pattern to remove portions of the corneal tissue primarily by photospallation so that collateral thermal damage on the tissue is less than 1 micron thick.

An eighth aspect of this invention is directed to a medical apparatus for removing corneal tissue from an eye of a patient. The apparatus comprises a laser source that produces pulses of mid-infrared radiation, wherein the infrared radiation has a wavelength approximately corresponding to a corneal absorption peak. The apparatus also includes a scanner-deflection means to direct the pulsed radiation across an area of the corneal tissue in a predefined pattern to remove portions of the corneal tissue primarily by photospallation so that collateral thermal damage on the tissue is less than 1 micron thick. BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a block diagram illustrating the functional relationship of optical, mechanical, and electrical components of apparatus incorporating features of the present invention;

FIG. 2 is an expanded schematic diagram of the optical components of FIG. 1;

FIGS. 3(a) and 3(b) illustrate scanning patterns for the laser beam passing over the cornea;

FIGS. 4(a) and 4(b) illustrate intensity profiles as a function of the diameter of the focused laser beam, measured at the cornea; and

FIGS. 5(a) and 5(b) illustrate mechanisms for transferring the laser beam from the laser system to the surgical apparatus.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

As shown in FIGS. 1 and 2, a surgical apparatus 200 includes an infrared laser source 20 and an optical assembly, including, in sequence, beam transfer optics 30, discussed below in conjunction with FIG. 5, a safety shutter 40, and partially-transmitting mirrors 50 and 60, which cooperate to focus an output beam 10 upon the cornea of a patient's eye 70, for correcting curvature of the cornea or for affecting therapeutic interventions. The laser source 20 is preferably a mid-infrared laser generating short laser pulses, to yield a tissue removal mechanism based on photospallation, discussed below. The laser beam 10 is preferably scanned over a specific central region of the surface of the cornea in a predefined manner, as discussed below in conjunction with FIGS. 3(a) and 3(b), so as to selectively remove tissue at various points within the cornea and thereby cause the curvature of the cornea to change in a predictable and controlled fashion. According to one feature of the invention, the laser source 20 is preferably a solid state laser, which emits pulsed radiation in the mid-infrared spectral region with an energy density capable of causing ablative decomposition of corneal tissue. As used herein, the term "solid state laser" includes a diode laser. Preferably, the laser emits radiation at a corneal absorption peak, i.e., at a wavelength of approximately 2.94 microns, such as 2.7 to 3.1 microns, corresponding to the maximal absorption coefficient of water, the main constituent of the corneal tissue. It has been found that absorption of laser energy by the corneal tissue of the eye 70 at such a wavelength results in complete absorption within 1 to 2 microns of tissue depth. As discussed further below, it has been found that the combination of shallow absorption depths, short radiation pulses, and low fluences reduces the undesirable thermal damage of surrounding tissue to insignificant levels.

In one embodiment, the laser source 20 consists of the output from an optical parametric oscillator (OPO) which converts the frequency from a solid state Q- switched 1 micron laser to the desired mid-IR range using methods described in our co- pending patent application, US serial No. 08/816,097, which is incorporated by reference herein. Alternately, the mid-IR laser source may comprise the direct output of a solid state laser such as a short pulse Er: YAG, as described in our co-pending patent application US serial No. 08/854,565, which is also incorporated herein by reference.

PHOTOSPALLATION

As previously indicated, according to a feature of the present invention, the surgical technique disclosed herein, whereby corneal tissue is irradiated with short pulses of a scanned mid-infrared laser beam, is based on photospallation. Generally, photospallation is a photomechanical ablation mechanism which results from strong absorption of incident radiation by the corneal tissue. When the corneal tissue absorbs the infrared radiation, a bipolar oscillating shock wave is created, which alternately compresses and stretches the corneal tissue, ejecting the tissue fragments torn apart during the stretching phase. For a more detailed discussion of the photospallation process, see Dingus, et al. , "Griineisen-Stress Induced Ablation of Biological Tissue," SPIE, Vol. 1427, Laser-Tissue Instruction II (1991) and Jacques, S.L., "Laser-Tissue Interactions: Photochemical, Photothermal, and Photomechanical," Lasers in General Surgery, 72(3), 531-558 (1992), incorporated by reference herein.

As is described in detail below, because photospallation is a mechanical ablation process, there is very little heat generated in the adjacent tissue left behind after the ablation. The process is therefore readily distinguished from a photothermal mechanism where tissue is removed upon being heated to the point of vaporization. High fluences are often required in this thermal process in order to first raise the tissue temperature to the evaporation point and then to overcome the latent head of vaporization. With spallation, on the other hand, no such phase change is required since tissue can "break" away (spall) before it vaporizes. Consequently, ablation by way of spallation can occur at much lower fluences than photo vaporization.

The laser parameters that are used in one embodiment to induce photospallation so that collateral thermal damage levels are minimized is now described. In this embodiment, the laser is intended to ablate corneal tissue. Therefore, the laser wavelength is selected to ensure that the tissue absorption coefficient is greater than 10,000 cm^~l. This corresponds to an absorption depth of less than 1.0 micron, which confines most of the absorbed energy to a 1.0 micron thick surface layer. If the absorption coefficient is weaker than this, the absorbed energy spreads further into the adjacent tissue causing a thermal effect that is more extensive than levels desired. Preferably, the absorption coefficient should be strong enough to ensure that the ablated layer is sufficiently thin, e.g., less than 0.5 microns. This allows a tissue removal process that is precise enough to generate the correct new shape on the anterior surface of the cornea to accurately produce desired corrections in vision.

In this embodiment, the pulsewidth should also be short enough so that the energy absorbed does not escape from the layer of absorbing tissue (e.g., 1 micron thick) before the bi-polar stress wave can cause the tissue to spall, and be thereby removed along with any generated heat. In this embodiment in which the layer of tissue is corneal tissue, the pulsewidth is less than 50 ns, and preferably about 10 ns.

On the other hand, the pulsewidth should also be long enough to avoid the formation of plasma and the attendant shock waves due to plasma expansion. The phenomenon of plasma formation and subsequent photodisruption of tissue due to intense laser radiation has been extensively studied in the past decade. See, for example, the recent paper by A. Vogel et al. in IEEE Journal of Selected Topics in Quantum Electron 2, 847 (1996), and the many references therein. Most of the more recent work has focused on establishing thresholds in tissue transparent to the radiation, where high power densities on the order of 10^0 W/cm^ as well as high fluences (10's of J/cm^) are typically necessary to cause tissue breakdown by photodisruption. There are, however, indications that far lower fluences may be sufficient to produce plasma in absorbing tissue, as was shown, for example, by A.A. Oroevsky et al. in SPIE, Vol. 2391, 423 (1995), where experiments indicated that breakdown could be induced with low thresholds in highly absorbing collagen gels. This, however, is an issue only when high power densities are present as well, as would be the case if picosecond lasers are used. Therefore, in this embodiment, the pulsewidth thus is longer than about 1 ns, thereby assuring that peak power densities are kept under 10^ W/cm^, a value regarded as a lower limit for plasma formation and subsequent tissue ablation by photodisruption.

It is preferred that the plasma mediated ablation processes be avoided, as these are generally accompanied by significant stresses which might cause undesirable tissue damage due to shock wave disruption even at locations substantially remote from the surgical site. It is noted that for pulses shorter than 1 ns — which was taught by the prior art Dingus article - peak powers well in excess of 10^ W/cm² are readily reached with the spot sizes utilized in the present invention. Therefore, the use of such subnanosecond pulses could result in the plasma mediated ablation processes that are intended to be avoided. Nanosecond lasers are also preferred from a practical standpoint because they are considerably simpler, cheaper and easier to construct and maintain than subnanosecond lasers, such as picosecond lasers.

In this embodiment, the ablation fluence threshold is 10-75 mJ/cm². This range of fluence thresholds is consistent with values derived from the simple analysis of photospallation as given by Dingus et al., but tailored to our preferred laser parameters and effects desired in the procedures contemplated. In particular, We have observed that fluence thresholds lower than 30 mJ/cm² are possible for pulses as long as 7-10 ns assuming tissue spall strengths on the order of 20 bars, which are reasonable values for the cornea. With such low thresholds, there are compelling reasons to use the nanosecond pulses preferred in this invention.

In particular, even the clinical fluence, which is typically two to three times the ablation threshold value, is still only about 20-225 mJ/cm² at the most. However, because a clinical fluence of greater than about 200 mJ/cm² may result in contributions from photothermal processes that can increase the amount of collateral thermal damage, we prefer that the upper range for the clinical fluence be limited to 200 mJ/cm².

We also prefer that the pulses are emitted at repetition rates of at least 10-100 Hz. The higher repetition rates are desired both to minimize the duration of the procedure and to allow utilization of smaller spot sizes with better overlap parameters as needed in a scanning system for improved surface smoothness of the surgical site.

This combination of lower fluence on the target tissue, high absorption, and temporally shorter pulses results in a much lower percentage of the energy being converted into thermal energy resulting in a dramatically lower thermal damage to the residual tissue at the surface of the ablation. This is because most of the energy converts into the spallation effect which is mechanical and happens on a time scale too short for thermal energy to transport into adjacent tissue.

Using the above parameters, we have achieved collateral thermal damage zones that are as low as 0.1 micron, but no larger than 1 micron. This, of course, is dependent on the specific laser pulse duration and trade-offs between energy available from the laser, losses in the beam delivery system, spot sizes, spot overlap and available repetition rate.

The small damage zone of our short pulse laser is an important factor in producing such highly localized ablations. These localized ablations with low fluence also produce a small amount of tissue removal per pulse - typically less than 0.5 microns - allowing a smooth reshaping of the surface. Thus, the mid-IR laser would produce clinical indications that are similar to those produced by the UV radiation from the ArF excimer lasers, where submicron collateral damage has been routinely demonstrated.

For example, assuming a 1.5 mm spot size and a clinical fluence of 100 mJ/cm², the energy per pulse at the eye surface would be about 1.8 mJ. If the scanning delivery system loses 70% of the original energy (%T = 30%), then the laser would need to generate 6.0 mJ per pulse. With such low energy requirements, the promise of a mid-IR laser system for photorefractive surgery is realized for the first time, since both laser and delivery system based on these principles would be advantageous in terms of price and simplicity as compared with excimer laser systems. In this regard, it should be appreciated that even lower energies (of just over 1 mJ) could be sufficient if more efficient delivery systems and/or smaller spot sizes are used. This would result in a highly compact laser and a system far more economical than any of systems currently in use or previously proposed in the prior art.

LINE-OF-SIGHT

To correlate the eye's reference frame to that of the surgical instrument 200, as shown in FIGS. 1 and 2, it is necessary that the line-of-sight of eye 70 be substantially coincident with the propagation axis of the incident laser beam 14. As used herein, in accordance with customary definition, the term "line-of-sight" or "principal line of vision" refers to the chief ray of the bundle of rays passing through the pupil and reaching the fovea, thus connecting the fovea with the fixation point through the center of the entrance pupil. It will therefore be appreciated that the line- of-sight constitutes an eye metric defined directly by the patient, rather than through some external measurement of the eye and further, that the line-of-sight can be defined without ambiguity for a given eye and is the only axis amenable to objective measurement using cooperative patient fixation.

Since critical vision is, by definition, centered on the line-of-sight of the eye, irrespective of the direction in which the mechanical axis of symmetry of the eye is pointed, it is generally acknowledged that for best optical performance, the point marking the intersection of the line-of-sight with the cornea establishes the desired center for the optical zone of refractive procedures seeking to restore visual acuity. It is noted that the orientation of the line-of-sight of the eye 70, as shown in FIGS. 1 and 2, may be vertical, horizontal, or intermediate to those extremes as befitting comfortable positioning of the patient for surgery without affecting the validity of the invention.

During preparation for laser surgery on the cornea, the line-of-sight of the eye 70 must be aligned to coincide with the laser beam axis by two-axis lateral- translational adjustments, in a known manner, as directed by the surgeon 55. The surgeon 55 observes the eye 70 through a surgical microscope 80 and judges the degree of centration of the frontal image of the eye 70 with respect to a crosshair or other fixed reference mark indicating, as a result of prior calibration, the location of the axis of beam 14, in a known manner. The axial location of the eye 70 can be judged by the surgeon 55 by virtue of the observed degree of focus of the image of the eye 70 relative to the previously calibrated and fixed object plane of best focus for microscope 80. Directions from the surgeon 55 allow adjustment of the axial position of the cornea of eye 70 to coincide with the plane of best focus.

The required angular orientation of the line-of-sight of eye 70 is preferably established by directing the patient to observe and focus attention, i.e., fixate, on two coaxial illuminated targets (not shown) projected into the eye 70 by a fixation target device 90, which is preferably integrated into the microscope 80. The two targets appear to be located at different axial distances from the eye 70 of the patient and will have been previously calibrated in a known manner. For a description of a suitable calibration technique, see PCT application No. WO 94/07908 to Knopp and Yoder. In this manner, when the two targets (not shown) appear superimposed, the axis of the observing eye 70 will be substantially coincident angularly with the axis of the microscope 80 and also with the axis of laser beam 14.

In a preferred embodiment, small lateral motions of the patient's eye 70, i.e., less than 5 mm in either direction, that occur after the initial alignment performed in the manner described above, and throughout surgical treatment, are rendered inconsequential by virtue of the function of a two-dimensional eye tracker 100, discussed further below in conjunction with FIGS. 6 and 7. The eye tracker 100 senses the motion of the eye 70 and provides signals that are proportional to the errors in lateral alignment of the eye 70 relative to the axis of the laser beam 14. The signals generated by eye tracker 100 are converted into commands for small angular tilts of partially-reflecting mirror 60 that compensate for errors in the location of the eye 70. The small angular tilts serve to redirect beam 14 so as to make it coincide with the instantaneous position of the eye 70. The compensation commands are sent from electronics, discussed below in conjunction with FIG. 7, within the eye tracker 100 to mirror 60 by means of one or more data connections, collectively designated 102.

Illumination of the eye 70 to facilitate tracking by the eye tracker 100 is preferably accomplished by means of a coaxial illuminator 120, preferably integrated with the microscope 80, that projects a beam of light 17 at a small angle, on the order of 8°, with respect to the line-of-sight of the microscope 80. According to a feature of the invention, the nature, i.e., the wavelength and temporal modulation frequency, of the tracking beam 17 generated by illuminator 120 is preferably selected to maximize discrimination by the detectors and related electronic circuitry within eye tracker 100 of the tracking beam 17, from ambient room illumination and radiation from laser 20. In this manner, the ambient illumination and laser beam 14 will not possess the same temporal modulation nor spectral characteristics as the tracking beam 17, and will thus be virtually invisible to the tracking detectors.

In addition, as shown in FIG. 1 , the surgical system 200 preferably includes a safety shutter 40 which closes automatically if the laser beam 14 fails to follow a prescribed path, if pulse energy-monitoring means provided within laser 20 indicates a malfunction of said laser or if the eye tracker 100 cannot follow the eye motion.

As shown in FIG. 1 and discussed further below, the surgical apparatus 200 preferably includes a video camera 140 that displays a real-time image of the patient's eye on a monitor 150 during pre-operation alignment and during surgical treatment and records the video image on a video recorder 160 for postoperative examination and documentation of the surgical procedure.

As shown in FIG. 1, the computer 110 includes multiple storage and control capabilities. Specifically, the computer 110 communicates and thereby controls the laser source 20 by means of a connection 101. In addition, the computer 110 drives the scanning mirror 50 by means of a connection 103, in accordance with stored scanning patterns and commands input to the computer 110 by the surgeon 55 or an assistant. A connection 104 between the computer 110 and the safety shutter 40 affects maximum safety of the patient, the surgeon, and attending personnel. The computer 110 monitors the operation and status of the eye tracker system 100 by means of a connection 105. Alternately, as shown in FIG 1, computer 110 can be connected to the eye tracker 100 by means of connection 106 and a separate connection 107 can be provided from computer 110 to mirror 60 so that the computer 110 could directly control the position of the mirror 60. A further alternate configuration would allow the computer 110 to combine the scanning and eye tracking functions together onto a single mirror, such as the mirror 60, thereby removing the need for connection 103.

As discussed further below, the surgical apparatus 200 preferably includes a corneal topography device 180 or a spatially resolved refractometer 190, as shown in FIG. 1. A corneal topography device 180 may be used for evaluating the shape of the corneal tissue to assist in pre-op and post-op measurements of the eyes' shape or curvature. An alternate embodiment would include a spatially resolved refractometer (SRR) 190 to evaluate the refraction of the corneal tissue.

OPTICAL MIRRORS

It may be noted from examination of FIGS. 1 and 2 that the partially- reflecting natures of mirrors 50 and 60 play important roles in the proper function of the invention. In the case of mirror 50, laser radiation in beam 12 is reflected while radiation from eye tracker 100 is transmitted. This can be accomplished, for example, through use of what is commonly called a "hot mirror" coating on the surface of mirror 50. This coating is dichroic, in other words, the coating has different reflection and transmission characteristics for light of differing wavelengths. The radiation from laser 20 has a wavelength of approximately 2.9 microns and the mirror 50 should have a high reflectance at that wavelength. The radiation to eye tracker 100 preferably has a wavelength between 0.8 and 1.0 microns for which the coating of mirror 50 should have a high transmittance.

Similarly, the dichroic coating on mirror 60 is preferably selected to have high reflectance at the wavelength of laser 20 and approximately equal transmittance and reflectance at the visible wavelengths used by the surgeon's eye in observing the alignment of the eye with respect to the surgical apparatus and progress of the surgery, at the wavelength of the fixation target 90, and at the wavelength of the coaxial illuminator 120. This is possible since the visible range, the fixation target 90, and the illuminator 120 are adjacent in wavelength and far from the wavelength of laser 20. At both mirrors 50 and 60 the transmitted beams suffer small lateral displacements due to oblique incidence and the finite thickness of the mirror substrates, but these fixed displacements are easily compensated for in the design of the apparatus, as would be apparent to a person of ordinary skill in the art.

In addition, mirror 130, shown between beams 15 and 16 of FIGS. 1 and 2, is also preferably partially transmitting, although not dichroic. The coating on mirror 130 nominally has approximately equal reflectance and transmission characteristics at the wavelengths of the eye tracker light source 120 and throughout a significant portion of the visible spectral region. In this manner, a portion of the energy of beam 15 can be redirected as beam 18 into video camera 140, discussed above. It is understood that a beamsplitting prism, typically in the form of a cemented two-element cube with a partially-reflecting coating on an internal surface can be employed to provide the function of mirror 130.

SCANNING PATTERNS

As previously indicated, the surgical apparatus 200 of FIGS. 1 and 2 preferably provides a computer-controlled scanning motion of the focused laser beam 14 for sequentially irradiating contiguous small areas of the central portion of the cornea of eye 70 with pulses of mid-infrared laser radiation in predefined patterns, such as those illustrated in FIGS. 3(a) and 3(b). In each case, the region to be treated has a diameter of up to 9 mm. The size of the focused spot of laser radiation is preferably on the order of a 0.5 to 2.0 mm circumscribed diameter.

As shown in FIG. 3(a), a rectilinear or raster-scan 310 of the scanning spot of laser beam 14 covers a square area centered on the desired treatment region 315. The laser beam 14 is modulated "off" when the computer 110 predicts that the energy would impinge upon corneal tissue outside the desired treatment region 315. As shown in FIG. 3(b), the laser beam 14 scans in a concentric-circle pattern 322 that is centered on the desired treatment region 325. While the path of the laser beam 14 may be continuous from start to finish, as indicated in the illustrative modes of Figs. 3(a) and 3(b), an alternative operational mode divides the pattern into a list of location coordinates and covers the entire area in a discontinuous fashion in order to minimize residual thermal effects of the area adjacent to the scan path by cumulative irradiation in rapidly sequenced locations of the beam. In this embodiment, the scanner would have random access capability to each location.

In the illustrative modes shown in FIGS. 3(a) and 3(b), or in other continuous or discontinuous scan patterns which would be apparent to persons of ordinary skill in the art, based on the disclosure herein, adjacent scan paths nominally overlap in a controlled manner. In this manner, the entire treatment region 315, 325 is uniformly irradiated with minimal discernible lines of overexposed or underexposed tissue lying between the scans. It is noted that the discontinuous property of the sequence distributes the pulses over the entire area in each time interval which is short compared to the entire sequence, thereby better distributing any residual heat to the entire surface and minimizing the buildup of heat and any temperature rise in any localized area. Once the pattern is defined by the computer 110, the implementation of the delivery can be discontinuously distributed across the corneal surface for maximum surface smoothness and minimum thermal effect. Scanning of the laser beam over the cornea surface is accomplished by a controlled tilting of the partially -reflecting mirror 50 about two axes so the reflected beam is deviated in an appropriate manner. This scanning motion is imparted to electrically-driven tilting mechanisms attached to mirror 50 under control of computer 110 upon initiation of the surgical treatment.

The velocity of the scan motion is varied at different points within the treatment area 315, 325 in accordance with an algorithm prescribed by the surgeon 55 to cause more or less ablation to take place locally, thereby causing the desired changes in refractive power of the cornea's anterior surface to correct the patient's vision defects. Correction of astigmatic, or cylindrical, errors can be accomplished by driving the scan mirror at different speeds as a function of rotational location about the propagation axis in the pattern. This allows the laser beam 14 to selectively ablate more tissue near one meridian of the corneal surface than near the orthogonal meridian. The nonsymmetric scan motion can be superimposed upon the normal symmetric pattern to simultaneously correct spherical and cylindrical refractive errors.

As shown in FIG. 4(a), the intensity profile of the focused laser beam 14 at the corneal surface ideally is contoured as a rotationally-symmetric trapezoid, in order to facilitate uniform irradiation of the treatment region 315, 325. The essentially gaussian profile shown in FIG. 4(b) approximates the idealized intensity profile illustrated in FIG. 4(a). It is noted that for smaller beam diameters, i.e., up to 2 mm, impinging on the corneal surface, the tissue removal profile for excimer ablation approximates a gaussian shape, independent from the beam intensity profile. For intermediate diameters, however, i.e., from 2 to 4 mm, the ablation profile approximates the beam intensity profile of the excimer laser beam. For larger diameters, i.e., from 4 to 7 mm or more, the ablation profile is deeper at the edge than the center compared to the beam intensity profile.

Photospallation is similar to the excimer ablation mechanism described above in that the beam intensity profile is generally not critical to the design or ablation pattern when using a spot size of 2 mm or smaller. Unlike photovaporization, where the tissue ablation mechanism is photothermal, the tissue ablation mechanism for photospallation is photomechanical. Therefore, the ablation pattern depends on the beam diameter, rather than a specific intensity profile. Thus, as a further advantage, since the present invention depends on pulse diameter and is not particularly sensitive to minor variations in the beam intensity profile, laser design issues may be relaxed.

BEAM TRANSFER OPTICS

As previously indicated, laser beam 10 is transferred to the main portion of the surgical apparatus 200 by means of beam transfer optics 30, shown in greater detail in FIGS. 5(a) and 5(b). It is noted that for the often crowded environment of an operating room, a flexible arrangement, whereby the beam delivery is effectively decoupled from the laser system, is preferred. As shown in FIG. 5(a), the beam transfer optics preferably includes a focusing lens 160 to condense the laser beam 10 into the entrance aperture of a decoupled guided means 162, such as a flexible fiberoptic cable. The fiber-optic cable 162 should preferably be capable of transmitting the intense infrared laser radiation over some distance, i.e., across an operating room, without damage to the fiber-optic cable itself, or significant loss of laser energy.

The fiber-optic cable 162 can be embodied as a single- or multiple-fiber bundle, and comprised of a material that safely transmits the specific wavelength of the laser 20, such as glass, sapphire, or another crystal. It is noted that in the infrared wavelength range, the additional losses associated with the added components required by the decoupled beam transfer optics 30 will generally be quite small. Alternatively, the laser beam can be coupled to the scanner system by means of a flexible hollow waveguide (not shown).

Preferably, the fiber-optic cable 162 connects the laser 20 to the main portion of the surgical apparatus 200 in a manner that permits convenient location of the laser 20 in the vicinity of the surgical apparatus 200, but not necessarily in a specific location. As shown in FIG. 5(a), the laser radiation exiting the output aperture

163 of the fiber cable 162 is captured by a relay lens 164 that forms an image of the output aperture 163. As shown in FIG. 1, this image is then propagated along paths 11, 12, 13 and 14 by means of partially-reflecting mirrors 50 and 60, to position the image at the anterior surface of the cornea of eye 70. The image plane of relay lens

164 is positioned during assembly of the apparatus so as to lie at the plane of best focus of microscope 80. The fiber-optic cable 162 may be embodied as the SapphIRe product, commercially available from Saphikon, Inc., or in accordance with the teachings of U.S. Patent No. 5,349,590.

An alternate embodiment of the beam transfer optics 30 is shown in FIG. 5(b). The alternate arrangement of FIG. 5(b) replaces the fiber-optic cable 162 of FIG. 5(a) with a flexible articulated arm 166. The flexibility of the articulated arm 166, by rotation about axes B-C, C-D, D-E, E-F, and/or F-G, allows convenient location of the laser source 20 with respect to the main portion of the surgical apparatus 200, again without requiring the laser source 20 to occupy a specific location. Condensing and relaying of the laser radiation at input and output apertures of the articulated arm are accomplished by means of lenses 168 and 170 in a manner substantially as described for the corresponding optical components in FIG. 5(a). The articulated arm 166 may be embodied as the Light Guiding Arm, commercially available from Dantec Measurements Technology, or in accordance with the teachings of U.S. Patent No. 4,896,015.

Another alternate embodiment for the beam delivery system would place the laser on the arm of the surgical microscope in a fixed location with respect to the main portion of the surgical apparatus 200. Such an arrangement would require certain rigid relay means to transport the radiation, which may require greater care in optical alignment, while imposing additional packaging constraints. For these and other reasons, the decoupled means of FIG. 5(a) and Fig. 5(b) are preferred.

EYE TRACKER The present invention may be used in connection with an apparatus and/or method that tracks and compensates for movement of the eye during surgery. One such tracking device is disclosed commonly assigned, copending United States Patent Application entitled "Apparatus And Method For Tracking And Compensating For Eye Movements", filed in the names of Telfair et al. and incorporated herein by reference. The invention disclosed there is directed toward a system and method for compensating for movement of an eye of a patient during a surgical procedure. The eye has a feature and a visual axis associated therewith, wherein the feature is illuminated with ambient light. The surgical procedure includes directing a laser beam upon the eye using a mirror. The laser beam has an optical axis associated therewith. The system includes illumination means for illuminating at least the feature of the object with a tracking light. The system also includes detection means for detecting an image of the feature and for outputting signals corresponding to movement of the image, wherein the signals have a first component due to the tracking light and a second component due to the ambient light. A filter means is also included for filtering the second component from the signals and for outputting the first component of the signals so that the ambient light is discriminated from the tracking light. The system further includes logic means for receiving the filtered signals and for generating tracking signals based thereon. The system also includes means for directing the laser beam upon the eye based on the tracking signals to maintain a substantially centered condition between the optical axis of the laser beam and the visual axis of the eye. TOPOGRAPHIC MEASUREMENTS As previously indicated, a corneal topography device 180 may be used to assist in pre-op and post-op measurements of the eyes' shape or curvature. Any commercially available topographic instrument may be used for this purpose as long as it is modified to include reference targets for fixation as utilized by the present invention. An alternate embodiment would include in this location a Spatially Resolved Refractometer (SRR) 190 to measure true refraction across the cornea.

The ability to establish a common reference frame between different ophthalmic instruments is of further importance in consideration of the desirability of integrating the method of corneal surgery that is the subject of the invention with independent refractive and/or topographic measurements of the cornea. It is generally recognized that accurate measurement and determination of the refractive status of the eye is desirable for a successful outcome of any refractive surgical procedure.

Corneal topographic devices, such as those manufactured by EyeSys and Computed Anatomy, have had some utility in providing evaluation of pre- and postoperative shape of the cornea. Other instruments that have recently become available, such as the OrbScan product by Orbtek, Inc., may provide information about the local shape of the cornea which can be highly useful for optimizing the correction of certain types of refractive errors, such as astigmatism. For any of these instruments to be effective, however, it must be compatible with repeated measurements being referenced to the same location in the eye. This aspect can be provided by an eye tracking or fixation technique, in the manner described above, that is unique to a patient and not to an instrument. Inclusion of such an alignment feature may also allow intraoperative measurement of corneal topography which could be used as an active feedback during the procedure for the purpose of enhancing the precision of surgery and eliminating undesirable variables affecting predictability. Prior art as described by U.S. Patent No. 5,350,374 to Smith shows the possibility of integrating an active feedback control loop based on a particular type of topographic instrument with a corneal surgery procedure.

In various embodiments, the present invention also seeks to include topographic feedback that is compatible with any number of available corneal measurement devices thus incorporating many of the advantageous features of the prior art devices, but enlarging their scope to include PRK surgery with a mid-infrared laser using a scanning beam delivery system.

An alternative to the shape mapping of these topography devices is the refraction mapping device and method called Spatially Resolved Refractometer (SRR). For a detailed discussion of SRR, see Webb, R.H., Murray Penny, C, Thompson, K.P. , "Measurement of Ocular Local Wavefront Distortion with a Spatially Resolved Refractometer," Applied Optics, 31, 19, 3678-3686 (1992). The SRR device measures the refraction at each point on the cornea over the pupil by having a patient align two fixation sources through a small pinhole. This pinhole is translated across the cornea to map each point of the cornea with a separate refraction measurement. Since the purpose of PRK is to correct the refractive error of a patient, the SRR map is the ideal input for correction by the PRK system, providing an improvement over the refraction measured in a refracting lane, as well as the power map from a topography system. This preoperative input data may be used to help define the ablation profile and pattern. Alternatively, SRR may to used to map the eye during a procedure.

It is to be understood that the embodiments and variations shown and described herein are illustrative of the principles of this invention only and that various modifications may be implemented by those skilled in the art without departing from the scope and spirit of the invention.

Claims

We claim:

1. A method for removing corneal tissue from an eye of a patient, said method comprising the steps of: generating a pulsed beam of laser radiation for ablating said corneal tissue, wherein said beam comprises mid-infrared radiation at a wavelength approximately corresponding to a corneal absorption peak, wherein said beam has a pulse duration of at least about 1 nanosecond, and wherein said beam has a clinical fluence in the range of 20-200 mJ/cm²; and scanning said beam across an area of said corneal tissue in a predefined pattern to remove portions of said corneal tissue primarily by photospallation.

2. The method according to Claim 1, wherein said beam has a pulse duration of less than about 50 nanoseconds.

3. The method according to Claim 1, wherein said beam has an ablation threshold in the range of 10-75 mJ/cm².

4. The method according to Claim 1, wherein collateral thermal damage on the tissue is less than 1 micron thick.

5. The method of Claim 1, further comprising the step of stabilizing the eye with an eye tracker.

6. A method for removing corneal tissue from an eye of a patient, said method comprising the steps of: generating a pulsed beam of laser radiation for ablating said corneal tissue, wherein said beam comprises mid-infrared radiation at a wavelength approximately corresponding to a corneal absorption peak, wherein said beam has a pulse duration of at least about 1 nanosecond, and wherein said beam results in an ablation threshold in the range of 10-75 mJ/cm²; and scanning said beam across an area of said corneal tissue in a predefined pattern to remove portions of said corneal tissue primarily by photospallation.

7. The method according to Claim 6, wherein said beam has a pulse duration of less than about 50 nanoseconds.

8. The method according to Claim 6, wherein said beam has a clinical fluence in the range of 20-200 mJ/cm².

9. The method according to Claim 6, wherein collateral thermal damage on the tissue is less than 1 micron thick.

10. The method of Claim 6, further comprising the step of stabilizing the eye with an eye tracker.

11. A medical apparatus for removing corneal tissue from an eye of a patient, said apparatus comprising: a laser source that produces pulses of mid-infrared radiation, wherein said infrared radiation has a wavelength approximately corresponding to a corneal abso╧åtion peak, wherein said pulses have a duration of at least about 1 nanosecond and a clinical fluence in the range of 20-200 mJ/cm²; and a scanner-deflection means to direct the pulsed radiation across an area of said corneal tissue in a predefined pattern to remove portions of said corneal tissue primarily by photospallation.

12. The apparatus according to Claim 11, wherein said pulses have a duration of less than about 50 nanoseconds.

13. The apparatus according to Claim 11, wherein said pulses have an ablation threshold in the range of 10-75 mJ/cm².

14. The apparatus according to Claim 11, wherein collateral thermal damage on the tissue is less than 1 micron thick.

15. The apparatus according to Claim 11, further comprising an eye tracker for stabilizing the eye.

16. A medical apparatus for removing corneal tissue from an eye of a patient, said apparatus comprising: a laser source that produces pulses of mid-infrared radiation, wherein said infrared radiation has a wavelength approximately corresponding to a corneal abso╧åtion peak, wherein said pulses have a duration of at least about 1 nanosecond and results in an ablation threshold in the range of 10-75 mJ/cm²; and a scanner-deflection means to direct the pulsed radiation across an area of said corneal tissue in a predefined pattem to remove portions of said corneal tissue primarily by photospallation.

17. The apparatus according to Claim 16, wherein said pulses have a duration of less than about 50 nanoseconds.

18. The apparatus according to Claim 16, wherein said pulses have a clinical fluence in the range of 20-200 mJ/cm².

19. The apparatus according to Claim 16, wherein collateral thermal damage on the tissue is less than 1 micron thick.

20. The apparatus according to Claim 16, further comprising an eye tracker for stabilizing the eye.

21. A method for removing corneal tissue from an eye of a patient, said method comprising the steps of: generating a pulsed beam of laser radiation for ablating said comeal tissue, wherein said beam comprises mid-infrared radiation at a wavelength approximately corresponding to a corneal abso╧åtion peak; and scanning said beam across an area of said corneal tissue in a predefined pattern to remove portions of said comeal tissue primarily by photospallation so that collateral thermal damage on the tissue is less than 1 micron thick.

22. The method according to Claim 21, wherein said beam has a duration of at least about 1 nanosecond and less than about 50 nanoseconds.

23. The method according to Claim 21, wherein said beam has a clinical fluence in the range of 20-200 mJ/cm².

24. The method to Claim 21 , wherein said beam has a duration of at least about 1 nanosecond and an ablation threshold in the range of 10-75 mJ/cm².

25. The method of Claim 21, further comprising the step of stabilizing the eye with an eye tracker.

26. A medical apparatus for removing corneal tissue from an eye of a patient, said apparatus comprising: a laser source that produces pulses of mid-infrared radiation, wherein said infrared radiation has a wavelength approximately corresponding to a comeal abso╧åtion peak; and a scanner-deflection means to direct the pulsed radiation across an area of said corneal tissue in a predefined pattern to remove portions of said corneal tissue primarily by photospallation so that collateral thermal damage on the tissue is less than 1 micron thick.

27. The apparatus according to Claim 26, wherein said pulses have a duration of at least about 1 nanosecond and less than about 50 nanoseconds.

28. The apparatus according to Claim 26, wherein said pulses have a clinical fluence in the range of 20-200 mJ/cm².

29. The apparatus according to Claim 26, wherein said pulses result in an ablation threshold in the range of 10-75 mJ/cm².