HK1178774A - System for modifying eye tissue and intraocular lenses - Google Patents

Description

System for modifying eye tissue and intraocular lens

Background

Cataract extraction is one of the most commonly performed surgical procedures in the world. Cataracts are emulsions of the lens of the eye or its capsule, the lens capsule. It varies from slightly opaque to completely opaque which prevents the passage of light. During the early stages of the development of senile cataract, the power of the lens increases, causing myopia (myopia), and the gradual yellowing and emulsification of the lens reduces the perception of blue light, as these wavelengths are absorbed and scattered within the lens. Cataracts, if left untreated, often progress slowly causing vision loss and possible blindness.

Treatment is performed by removing the opaque lens and replacing it with an intraocular lens (IOL). An estimated 300 million cases are performed each year in the united states, and 1,500 million cases worldwide. This market is comprised of several parts including intraocular lenses for implantation, viscoelastic polymers to facilitate surgical procedures, disposable instruments including phacoemulsification tips, tubes, and various knives and forceps.

Modern cataract surgery is commonly performed using a technique termed phacoemulsification, in which an ultrasonic tip with associated irrigation and aspiration ports is used to sculpt the relatively hard nucleus of the lens to facilitate its removal through an opening in the anterior lens capsule termed anterior capsulotomy, or more recently, continuous tear circular capsulotomy (CCC). Finally, a synthetic foldable intraocular lens is inserted into the remaining lens capsule of the eye through a small incision.

The biggest technical challenge and critical step in this procedure is the capsulorhexis. This procedure has evolved from the earlier term open-can capsulotomy, in which a sharp needle is used to pierce the anterior lens capsule in a circular fashion, followed by removal of a circular fragment of the lens capsule, typically in the range of 5-8mm in diameter. This facilitates the next step, nucleus sculpting by phacoemulsification. Due to the various complications associated with the initial open-can capsulotomy technique, various attempts have been made by the authoritative specialist in the field to develop a better technique for removing the anterior lens capsule prior to the emulsification step.

The principle of continuous capsulorhexis is to provide a smooth continuous annular opening through which not only can phacoemulsification of the nucleus be safely and conveniently performed, but also can be used for convenient insertion of an intraocular lens. It provides an open central passageway for insertion, a permanent aperture for transmission of the image by the patient to the retina, and provides support for the IOL within the remaining lens capsule, thereby limiting potential misalignment.

Problems may arise relating to the surgeon's inability to adequately view the lens capsule due to lack of red light reflection, inability to grip the lens capsule firmly enough, inability to tear a smooth circular opening of appropriate size and location without creating radial tears and extensions. There are also technical difficulties related to the maintenance of the anterior capsule depth after the initial opening, the small size of the pupil, or the lack of red light emission due to lens opacity. Some of the problems of vision are minimized by the use of dyes such as methylene blue or indocyanine green. Other complications arise in patients due to the delicate zonules (typically elderly patients and very young children with very soft and elastic lens capsules) which are very difficult to rupture and tear controllably and reliably.

Many cataract patients have astigmatic vision errors. Astigmatism can occur when the corneal curvature is not uniform in all directions. IOLs are now used to correct astigmatism but require precise rotation and central positioning. In addition, even though many patients have more severe aberrations, IOLs have not been used to correct astigmatism beyond 5D. Higher correction of beyond 5D astigmatism requires reshaping the cornea to make it more spherical. There are a number of procedures that exist, including keratoplasty, astigmatic keratectomy, Corneal Relaxing Incisions (CRI), and Limbal Relaxing Incisions (LRI). Except for corneal transplants, all procedures allow the cornea to change shape to become more spherical by positioning the corneal incision in a well-defined manner and depth. These precision cuts are now positioned manually by indication on their limited precision.

But not only an incision is required for ophthalmic treatment. There is also a need for more flexible modifications to eye tissue that produce weakening of the mechanical properties of the tissue and/or changes in the optical properties of the treated tissue. In this case, the result should be sufficient flexibility to allow structural modification of the eye tissue without mechanical disruption. Ding et al (IOVS,2008(49),12, pp 5532-. However, practical application of Ding's technique is limited by the need to apply 100,000,000 laser pulses per cubic micron of tissue treated.

Vogel et al (US 2010/0163540a1) describe a method of machining and cutting a transparent material by a time-smoothed laser beam to produce a low density serum without forming a serum glow. In this teaching, they describe that it is particularly desirable to avoid linear adsorption of bare material, since it leads to random generation of seed electrons, which in turn generates random variations within the threshold of the serum. In addition, they describe that the formation of low density plasma is always associated with the formation of cavitation bubbles.

This is in sharp contrast to the present invention, where two working mechanisms are described. It was found that the use of laser wavelengths with some linear absorption in the target tissue can produce extremely low threshold effects. In addition, no time-smoothing pulse shape is required in the current invention. Also in one embodiment of the present invention, the formation of cavitation bubbles is not expected as a result of linear adsorption enhanced photolysis. And Vogel's data show that there is still more than an order of magnitude difference in the resulting plasma formation when comparing IR femtosecond laser and 355 sub-nanosecond laser. In our embodiment, the energy threshold for 355nm sub-nanosecond lasers is even slightly lower when compared to femtosecond lasers using the same numerical aperture optics due to linear adsorption of tissue endogenous chromophores (or by addition of exogenous chromophores).

Braun et al (DE 19855623C 1) describe a method of precision machining in glass using a laser that transmits wavelengths outside the plateau of the glass. The laser is then used specifically to create material defects in the glass without including the surface. This method allows them to locate material defects closer to the surface without damaging the surface itself. It does not describe surface effects. Since it is used only on glass (where no cavitation bubbles are formed), it also does not produce any cavitation events.

Koenig et al (WO 2007/057174) claim a system for eye surgery by using femtosecond laser pulses in the UV spectral range. In his teaching, he described using a higher numerical aperture of 0.8 for his invention, which significantly reduced the threshold to the nano-focus level. But it is difficult to transplant his system to a usable product because it is optically difficult to combine these numerical apertures with a 6-10mm wide scan range commonly used in ophthalmic applications. Also the generation of femtosecond UV laser pulses is technically challenging.

What is needed is a method, technique and apparatus for improving the standard of care of ophthalmic patients.

Disclosure of Invention

One embodiment is directed to an ophthalmic surgical system comprising a laser source configured to transmit a laser beam comprising a plurality of laser pulses having a wavelength between about 320 nanometers and about 430 nanometers and a pulse width between about 1 picosecond and about 100 nanoseconds; and an optical system operatively coupled to the laser source and configured to focus and direct the laser beam in a pattern to one or more intraocular targets in the patient's eye such that interaction between the one or more targets and the laser pulse is characterized by linear absorption enhanced photolysis without formation of plasma or related cavitation events. The wavelength may be about 355 nm. The pulse width may be between about 400 picoseconds and about 700 picoseconds. The pulse may have a pulse energy between about 0.01 microjoules and about 500 microjoules. The pulse may have a pulse energy between about 0.5 microjoules and about 10 microjoules. The plurality of laser pulses may have a repetition rate of between about 500 hertz and about 500 kilohertz. The optical system may be configured to focus the laser beam to produce a beam diameter of between about 0.5 microns and about 10 microns in the one or more intraocular targets. At least one of the one or more intraocular targets may be selected from the group consisting of a cornea, a capsular rim, a sclera, a lens capsule, a lens, and a synthetic intraocular lens implant. The pattern may be configured to produce one or more physical degenerations, such as cuts (incisions) and refractive index changes, in the intraocular target in a configuration selected from the group consisting of a corneal relaxing incision, a limbal relaxing incision, an astigmatic keratectomy, and a capsulotomy. The optical system and laser source may also be configured to structurally alter at least one of the one or more intraocular targets such that the refractive index of the altered tissue structure target changes.

Another embodiment is directed to an ophthalmic surgical system, comprising a laser source configured to transmit a laser beam comprising a plurality of laser pulses having a wavelength between about 320 nanometers and about 430 nanometers and a pulse width between about 1 picosecond and about 100 nanoseconds; and an optical system operatively coupled to the laser source and configured to focus and direct the laser beam in a pattern to one or more tissue structure targets within the patient's eye such that interaction between the one or more targets and the laser pulse is characterized by local formation of plasma, which is facilitated by linear absorption. The wavelength may be about 355 nm. The pulse width may be between about 400 picoseconds and about 700 picoseconds. The pulse may have a pulse energy between about 0.01 microjoules and about 500 microjoules. The pulse may have a pulse energy between about 0.5 microjoules and about 10 microjoules. The plurality of laser pulses may have a repetition rate of between about 500 hertz and about 500 kilohertz. The optical system may be configured to focus the laser beam to produce a beam diameter of between about 0.5 microns and about 10 microns in the one or more tissue structure targets. At least one of the one or more tissue structure targets may be selected from the group consisting of a cornea, a capsular rim, a sclera, a lens capsule, a lens, and a synthetic intraocular lens implant. The pattern may be configured to create one or more incisions in an intraocular target that is a tissue structure target in a configuration selected from the group consisting of a corneal relaxing incision, a limbal relaxing incision, an astigmatic keratectomy, and a capsulotomy.

Another embodiment is directed to an ophthalmic surgical system, comprising a laser source configured to transmit a laser beam comprising a plurality of laser pulses having a wavelength between about 320 nanometers and about 430 nanometers and a pulse width between about 1 picosecond and about 100 nanoseconds; and an optical system operatively coupled to the laser source and configured to focus and direct the laser beam in a pattern to one or more targets within the patient's eye such that interaction between the one or more targets and the laser pulse is characterized by linear absorption enhanced photolysis without formation of plasma or related cavitation events. The pattern may be configured such that operation of the optical system and the laser source causes physical changes to the one or more targets. The physical change may be manifested as a change in the index of refraction of the one or more targets or one or more kerfs. At least one of the one or more targets may be a cornea or an intraocular lens. The physical change may be configured to alter a refractive profile of the target.

Another embodiment is directed to an ophthalmic surgical system, comprising a laser source configured to transmit a laser beam comprising a plurality of laser pulses having a wavelength between about 320 nanometers and about 430 nanometers and a pulse width between about 1 picosecond and about 100 nanoseconds; and an optical system operatively coupled to the laser source and configured to focus and direct the laser beam in a pattern to one or more tissue structure targets within the patient's eye such that interaction between the one or more targets and the laser pulse is characterized by linear absorption enhanced photolysis without formation of plasma or related cavitation events; and an integrated imaging subsystem that captures, in a confocal arrangement, back-reflected light from the sample provided by the laser source. The laser pulses may induce fluorescence, which is collected by an imaging subsystem. The system may be configured to provide interleaved lower energy pulses for imaging and higher energy pulses for therapy. The imaging subsystem further includes an optical interference tomography imaging system, a Purkinje (Purkinje) imaging system, and/or a Scheimpflug (Scheimpflug) imaging system. The system also includes a controller configured to determine a position & shape of the ocular structure, determine a placement mode and/or laser parameters, and position the mode in a defined target.

Another embodiment is directed to an ophthalmic surgical system, comprising a laser source configured to transmit a laser beam comprising a plurality of laser pulses having a wavelength between about 320 nanometers and about 430 nanometers and a pulse width between about 1 picosecond and about 100 nanoseconds; and an optical system operatively coupled to the laser source and configured to focus and direct the laser beam in a pattern to one or more tissue structure targets within the patient's eye such that interaction between the one or more targets and the laser pulse is characterized by linear absorption enhanced photolysis without formation of plasma or related cavitation events; and an exogenous chromophore introduced to the target structure to create/enhance linear absorption. The exogenous chromophore may be trypan blue (trypan blue).

Another embodiment is directed to an ophthalmic surgical system, comprising a laser source configured to transmit a laser beam comprising a plurality of laser pulses having a wavelength between about 320 nanometers and about 430 nanometers and a pulse width between about 1 picosecond and about 100 nanoseconds; and an optical system operatively coupled to the laser source and configured to focus and direct the laser beam in a pattern to one or more intraocular targets within the eye of the patient such that interaction between the one or more targets and the laser pulse is characterized by linear absorption enhanced photolysis without formation of plasma or related cavitation events; and a second laser source configured to fragment the lens with a wavelength between about 800nm and about 1100 nm. The second laser may be a pulsed infrared laser. The second laser may have a pulse width between about 1 picosecond and about 100 nanoseconds. The second laser may be a Q-switched Nd: YAG laser.

Drawings

FIG. 1 illustrates a high level flow chart according to an embodiment of the present invention.

Fig. 2A and B are schematic diagrams of system embodiments.

Fig. 3 shows a flow chart of a method according to an alternative embodiment.

FIG. 4 is a schematic of the application of a line spectrum through the lens for depth range measurement of an axial section of the anterior capsule of the eye (OCT, confocal reflectance, confocal autofluorescence, ultrasound).

Figure 5 is a top view of a rotationally asymmetric capsulorhexis incision.

Figure 6 is a top view of a complementary rotationally asymmetric IOL.

Fig. 7 is a top view of the IOL of fig. 6 positioned in the lens capsule of fig. 5.

Figures 8 and 9 are side views of the rotationally asymmetric IOL of figure 6.

FIG. 10 illustrates a fragmentation pattern for an ophthalmic lens produced by one embodiment of the present invention.

FIG. 11 is a schematic of applying a line spectrum 501 through the cornea 504 and lens for depth range measurements (OCT, confocal reflectance, confocal autofluorescence, ultrasound) of an axial section of the anterior capsule of the eye. It passes through iris 502 and lens 402 (not shown).

Fig. 12 illustrates a survey scan pattern through the cornea and lens that can be used for depth range through OCT.

Fig. 13 illustrates a measurement scan pattern through the lens that can be used to pass the depth range of confocal autofluorescence using 320NM to 430NM laser light.

FIG. 14 is another illustration of a system according to an embodiment of the invention.

FIG. 15 shows a tissue cross-section of a corneal incision made by an embodiment of the present invention in which no cavitation bubbles are formed but the tissue is modified.

FIG. 16 shows a tissue cross-section of an open corneal incision made by one embodiment of the present invention in which no cavitation bubbles are formed, as in FIG. 15. The incision is effortlessly spread along the revised anatomy.

FIG. 17 shows a tissue cross-section of a corneal incision made by an embodiment of the present invention in which cavitation bubbles are formed.

Fig. 18 shows a schematic of the refractive index change 822 induced locally to the corneal tissue 504 by the described invention. It is seen from fig. 15 that no airborne bubbles will be generated in this case. This result will produce a change in the refractive index profile of the corneal tissue.

Figure 19 shows a high resolution SEM image of an ex vivo human lens capsule treated by the present invention. In contrast to fig. 20, the sample had a smoother edge quality and did not show any cavitation bubble effect.

Figure 20 shows a high resolution SEM image of an ex vivo human lens capsule treated by a femtosecond laser. The effect of each individual laser emitted at 5 mm intervals is visible due to cavitation mechanical effects causing rupture of the capsule tissue.

Detailed Description

The present invention relates to methods and systems for making incisions or altering the mechanical or optical properties of eye tissue.

As shown in the drawings for purposes of illustration, a method and system for making an incision or altering the mechanical or optical properties of ocular tissue is described. In varying embodiments, the method and system provide a number of advantages over current standards of care. In particular, a 320nm to 430nm laser can be used to make a fast and precise opening in the lens capsule to facilitate positioning and stabilization of the intraocular lens.

Other procedures that can be allowed by the techniques described herein include astigmatism therapy. Intraocular lenses (IOLs) are commonly used to correct astigmatism but require precise positioning, orientation, and stability. Complete and long-term correction with IOLs is difficult. It typically involves other surgical procedures to make the corneal shape more spherical, or at least less radially asymmetric. This will be achieved by making a corneal relaxing incision or limbal relaxation. Other procedures include the generation of corneal flaps for LASIK procedures, and the generation of corneal graft shapes that match the donor and recipient corneas. The invention is also useful for performing these precision cuts.

Fig. 1 is a flow chart of a method according to an embodiment. The first step 101 includes a laser system that produces a 320nm-430nm beam of light having at least a first pulse of light. A next step 102 includes passing the light beam through an optical element such that the light beam is focused at a predetermined depth in the eye tissue. By implementing this method, a fast and precise opening in the lens capsule can be allowed, thereby facilitating the positioning and stabilization of the intraocular lens.

The present invention can be implemented by a system 200 that projects or scans a light beam into a patient's eye 20, such as the system shown in FIG. 2. The system 200 includes an electronic controller 210, a light source 220, an attenuator 230, a beam expander 240, focusing lenses 250, 260, and a reflecting device 270. The electronic controller 210 may be a computer, microcontroller, or the like. Scanning may be achieved by using one or more movable optical elements (e.g., lenses 250, 260, reflecting device 270) that may also be controlled by electronic controller 210 through input and output devices (not shown). Another means of scanning may be allowed by electro-optical deflection means (single or dual axis) in the optical path.

In operation, the light source 200 generates the light beam 225 and the reflecting device 270 can be tilted to deviate the light beam 225 and direct the light beam 225 toward the patient's eye 20. The focusing lenses 250, 260 may be used to focus the light beam 225 to the patient's eye 20. The positioning and characteristics of the beam 225 and/or the scan pattern formed on the eye 20 may be further controlled by using an input device such as a joystick or any other suitable user input device.

The invention can alternatively be performed by a system 700 that additionally makes range measurements of the patient's eye 20, such as the system shown in fig. 14. System 700 includes electronic controller 210, light source 220, attenuator 230, beam expander 701, variable beam attenuator 230, individual focusing lens assemblies 704, and beam reflecting and scanning device 270. The light beam 225 of the light source 220 is focused to its target location 20 by the focusing lens 260. This will be controlled by the electronic controller 210, which is connected to the deflection unit 270. In addition, autofluorescence 725 of target structure 20 is sub-scanned (de-scan) by the preferred means of dichroic beam splitter 703 through the same optical path shared with laser 225, and focused by lens 720. The aperture stop 721 is placed at the focus of the shaped beam 725 as a conjugate of the laser beam 225 focused on the target structure 20. The intensity of the autofluorescence transmitted through the beam aperture 721 is detected and converted into an electrical signal that can be read out by the control unit 210. And imaging the image of the treatment area through the lens 711 onto the image capture device 710, which image capture device 710 can be a CCD or CMOS camera. The signal is also transmitted to the control unit 210.

In another variant of the system 700, the back reflected beam 225 from the sample 20 is detected confocally using a detection combining unit 703, 720, 721, 722.

The basic mechanism of the different embodiments is to use a laser source of 320nm to 430 nm. The UV spectrum is technically subdivided into three main spectral regions, UVA (400nm-315nm), UVB (315nm-280nm), UVC (280nm-100 nm). UVB and UVC are often associated with carcinogenic effects due to their high single photon energy, due to their ability to directly denature DNA. Although moisture remains transparent as low as 200nm, protein absorption increases strongly around 240 nm. Strong protein absorption in the UVC spectral region, which is also the major absorption by corneal tissue, is now used clinically in laser intracorneal remodeling (LASIK) to precisely ablate corneal tissue.

UVC lasers are capable of ablating biological tissue by photolysis, absorbing high energy photons to break bonds within organic molecules. A list of these general bonds together with their dissociation energy in terms of wavelength is given in the table below. The shorter the wavelength, the stronger the bond.

It is evident from the table that photolysis of biological material requires high energy photons, such as discussed in U.S. patent No.4,784,135 to Blum et al. This effect is the basis of numerous imaging medical systems, especially in ophthalmology, where 193nm excimer lasers are routinely used for corneal correction. Embodiments of the present invention utilize disparate physical phenomena and different spectral regions (UVA to green) to denature and/or ablate biological tissue, which are not present and contemplated in the prior art.

In an embodiment, the light source 220 is a 320nm to 430nm laser source, such as a Nd: YAG laser source, operating at a third harmonic wavelength of 355 nm. The transmission of the cornea at 355nm is about 85% and starts to sharply drop at 320nm (50% transmission) to about 2% transmission at 300nm, while the lens absorption is about 99%. And, for the elderly, the light scattering of the cornea is minimal, while the light scattering of the lens is considerably increased (cataracts).

The light scattering effect is wavelength sensitive. In the case of scattering centers smaller than all wavelengths, the scattering coefficient is of the order λ^-4. For larger scatterers with a range of dimensions in the size of the wavelength, the Mie approximation is well suited for describing the scattering function. For particles having a size of 350 to 700nm in size, the scattering coefficient is of the order λ^-1. The aged lens itself absorbs all wavelengths shorter than 420nm and is a strongly scattering substance. This means that shorter wavelengths can be used for laser cutting of the anterior part of the lens, especially the lens capsule, of elderly people, while protecting the retina by effectively attenuating the light that is ultimately disposed there.

Posterior capsular cataract opacity is conventionally treated with a Q-switched infrared laser having a power of several millijoules and in the IR spectral range (1064 nm). They do so by providing reliable plasma formation directly behind the posterior capsule of the lens. These pulses produce cavitation bubbles of several millimeters in size and kilobar peak pressure. The mechanical effect of cavitation bubbles of millimeter-scale size is a limiting factor for high precision cutting in liquid environments. In order to reduce the bubble size and the matching mechanical side effects that produce cuts with poor edge quality and thus poor mechanical strength, it is necessary to greatly reduce the laser pulse energy. However, this interaction will be well suited for application of lens accommodation.

Open angle glaucoma is conventionally treated with a Q-switched green laser with a pulse length of several millijoules and nanoseconds. This treatment, termed Selective Laser Trabeculoplasty (SLT), utilizes specific targeting of melanin chromophores that naturally occur in the trabecular meshwork. The laser itself uses a relatively large 200 micron spot size to cover most of the target tissue area. The laser also creates cavitation bubbles around the melanin absorber, but the effect is produced by linear heating rather than by plasma formation as used in posterior capsule cataract treatment by Q-switching IR laser pulses.

In embodiments of the present invention, the use of UV wavelengths greatly reduces the threshold for plasma formation and associated cavitation bubble formation, and also reduces the threshold energy required for linear absorption enhanced photolysis without cavitation bubble formation for several reasons. First, the focused spot diameter expands linearly with wavelength, which coincides with the peak of radiation exposure in the focal plane. Second, the linear absorption of the material itself allows for a lower threshold for plasma formation or low density photolysis due to the initial absorption of more laser energy in the tissue structure. Third, the use of UV laser pulses on the order of nanoseconds and sub-nanoseconds allows linear absorption enhanced photolysis and chromophore guided ionization.

Furthermore, the chromophore-guided ionization strongly lowers the ionization threshold in the case of whey formation, as well as the threshold for low-density photolysis for material denaturation or change without cavitation even at very weak absorption. Due to the high flux density, even the lowest linear absorption strongly lowers the threshold for a certain effect. It has been shown (Colombelli et al, Rev. Sci. Instrum, 2004, Vol 75, pp.472-478) that the threshold for plasma formation and cavitation bubble generation can be reduced by an order of magnitude when changing only from high purity water to water with a physiological NADH concentration of 38 mMol. Linear absorption also allows for specific treatment of local lens structures (e.g., the lens capsule) since the light penetration depth of the laser beam is limited by the linear absorption of the lens. This applies in particular to the aged lenses, whose absorption in the UV-blue spectral range is strongly increased compared to the young lenses.

Additionally, in another embodiment of the present invention, the linear absorption effect on the target structure can also be enhanced by the application of exogenous chromophores. One such useful chromophore is trypan blue, which is commonly used in surgery to stain the lens capsule in the absence of fundus red light reflection. Trypan blue has an increased linear absorption at wavelengths below 370 nm. This linear absorption also reduces the energy required to produce an impact on the lens capsule surface.

The method can also be used for the variation of the overall reflected energy of the human eye by:

i. making cuts (incisions) in the cornea to change its shape, and thus its refractive power;

ii. Modifying the refractive index of the corneal tissue to cause a change in its effective refractive power;

iii modifying the refractive index of the implanted synthetic IOL by writing a Fresnel lens or other similar lens into the IOL material, thereby changing its effective refractive power;

any combination of iv, i, ii and iii.

The inventive system allows surgical techniques that include the use of 320nm to 430nm pulsed lasers to perform high precision physical degeneration of ocular targets, including tissues (such as the lens, lens capsule, cornea, etc.) and synthetic intraocular lens implants. This can be achieved by two different surgical protocols, i.e. with or without cavitation bubble formation. Sub-cavitation schemes can also be used to modify the refractive index of the ocular target. Although the wavelengths used in the present invention are shorter than or comparable to those associated with retinal blue light toxicity, the absorption of 320nm to 400nm laser light in the elderly lens also minimizes the risk of retinal damage. As this light will be absorbed by the lens volume. Furthermore, the risk of damaging the corneal endothelium or other corneal structures is minimized. The threshold pulse energy will be E_th=Φ*d²[ 4 ] where Φ is the threshold radiation exposure, and d is the focal spot diameter. Here, the focal diameter D is D = λ F/D_bWhere λ is the wavelength, F is the focal length of the final focusing element, and D_bIs the beam diameter of the final lens. For stable and repeatable operation, the pulse energy should exceed the threshold by at least a factor of 2, howeverRather, the energy level can be adjusted to avoid damage to the corneal endothelium.

In a variant embodiment, the 320nm to 430nm laser provides a pulse energy between 0.01 μ J and 500 μ J, preferably between 0.1 μ J and 10 μ J, an overlap frequency of 500Hz to 500kHz, a pulse width between 1ps and 100ns, preferably a pulse width of 400 ps. Laser spot diameters of less than 10 μm, preferably 3 μm to 0.5 μm, are used. An example of the results of this system on a real human lens is shown in fig. 10. A 4 μ J, 400ps pulsed beam output from a laser operating at 355nm wavelength with a 0.5kHz pulse repetition rate was focused at NA =0.15, using an irradiance of approximately 120 gigawatts per square centimeter. This produces the capsulotomy pattern of the human lens shown in fig. 10. In this case, no cavitation bubbles are formed to induce cutting. This was confirmed visually under a microscope and by using hydrophones for acoustic detection from cavitation bubbles. For laser cataract surgery, the only high precision cut on the lens itself is the capsulotomy. This mode does not require high space constraints for softening or breaking the lens nucleus. Thus, for this reference, even with longer pulses, higher fluence and/or radiation thresholds are acceptable.

Fig. 3 is a flow chart of a method according to an alternative embodiment. The first step 301 comprises a laser system generating a 320nm-430nm beam. A next step 302 includes transforming the beam in a controlled manner in the eye tissue to form an incision. In embodiments, in performing a capsulorhexis, the incision is made in the anterior lens capsule of the eye tissue. Alternatively, the incision may be in the cornea for astigmatism correction or to create a surgical corridor. For example, clear corneal cataract instruments and puncture incisions may be used to provide a surgical pathway.

The electronic controller 210 and the light source 220 can be set to aim at the surface of a target structure in the eye 20 and ensure that the light beam 225 will be focused at the proper location and not accidentally damage non-target tissue. The imaging modalities and techniques described herein, such as, for example, Optical Coherence Tomography (OCT), Purkinje (Purkinje) imaging, Scheimpflug (Scheimpflug) imaging, autofluorescence imaging, confocal autofluorescence, confocal reflectance imaging, or ultrasound, may be used to determine position and measure the thickness of the lens and lens capsule, thereby providing greater accuracy to the laser focusing method, including 2D and 3D patterning. Laser focusing can also be achieved by using one or more methods including direct observation of an aiming beam, OCT, purkinje imaging, structured light irradiation, ultrasound or other known ophthalmic or medical imaging modalities and/or combinations thereof. It should be noted that the imaging depth includes only the anterior-most portion of the intraocular target and need not include the entire eye or even the anterior capsule.

Additionally, a confocal reflector can be used for modulation of the laser energy delivered during treatment, as it can be used to detect whether there is cavitation bubble generation after a laser pulse and to modulate the energy of subsequent laser pulses or to monitor laser induced changes in the refractive index of the tissue.

Thus, the three-dimensional application of laser energy can be applied to the lens capsule in various ways along the pattern created by the laser-induced effect. For example, the laser can be used to produce several cycles or other mode scans in succession at different depths in steps equal to the axial length of the effect region. Thus, with each sequential scan, the depth of focus (convergence) in the tissue gradually increases or decreases. The laser pulses are sequentially applied to the same traversing pattern at different depths of the tissue, e.g., using an axial scan of the focusing element or adjusting the light intensity of the focusing element to optionally simultaneously or subsequently scan the traversing pattern.

The adverse consequences of laser beam scattering on bubbles, crevices and/or tissue fragments before reaching the focal point can be avoided by first creating a pattern/focus at the maximum desired depth of the tissue, and then subsequently focusing on a shallower tissue space in subsequent passes. This "inversion" treatment technique not only reduces undesirable beam attenuation in tissue above the target tissue layer, but also helps protect tissue below the target tissue layer. These defects help protect the underlying retina by scattering laser radiation transmitted beyond the focal point on the bleb, fissure, and/or tissue fragment (which resulted from the previous scan). Similarly, when the lens is segmented, the laser can be focused on the posterior-most portion of the lens and moved further forward as the procedure continues.

The present invention may be implemented by a system that projects or scans a light beam into a patient's eye 68, such as the system 2 shown in FIG. 2B, which includes a therapeutic light source 4 (e.g., a 355nm short pulse laser). Using this system, the beam in the patient's eye can be scanned in three dimensions X, Y, Z. Safety limits on unintentional damage to non-target tissue define upper limits on repetition rate and pulse energy; while the threshold energy, time and stability of completing the surgical procedure define the lower limits of pulse energy and repetition rate.

The laser 4 is controlled by an electronic controller 300 via an input and output device 302 to produce a light beam 6. Electronic controller 300 may be a computer, microcontroller, or the like. In this example, the entire system is under the control of a controller 300 and data is moved through an input/output device IO 302. A graphical user interface GUI 304 may be used to set system operating parameters, process User Inputs (UI)306 on the GUI 304, and display collected information such as images of eye structures.

The resulting treatment beam 6 travels toward the patient's eye 68, which passes through the half-wave plate 8 and the linear polarizer 10. The polarization state of the beam can be adjusted so that the desired amount of light passes through half-wave plate 8 and linear polarizer 10, which together act as a variable attenuator for the therapeutic beam 6. In addition, the orientation of linear polarizer 10 determines the polarization state of the incident light incident on beam combiner 34, thereby optimizing the flux of the beam combiner.

The treatment beam travels through shutter 12, aperture 14 and pickup device 16. The system controlled shutter 12 ensures on/off control of the laser for procedural and safety reasons. The aperture sets a useful outer diameter for the laser beam and the pickup device monitors the output of the useful beam. The pick-up device 16 comprises a partially reflective mirror 20 and a detector 18. The detector 18 may be used to measure pulse energy, average power, or a combination thereof. This information can be used to feed back to the half-wave plate 8 for attenuation and to verify whether the shutter 12 is open or closed. In addition, shutter 12 may have a position sensor to provide redundant status detection.

The beam passes through a beam conditioning stage 22 where beam parameters such as beam diameter, divergence, circularity and astigmatism can be modified. In the illustrated example, the beam conditioning stage 22 includes a 2-element beam expanding telescope that includes spherical optics 24 and 26 to facilitate achieving the desired beam size and collimation. Although not shown here, anamorphic or other optical systems may be used to achieve the desired beam parameters. 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 site. In addition, the optical system 22 may be used to image the aperture 14 to a desired position (e.g., a center position between 2-axis scanning devices 50 described below). In this way, it is ensured that the amount of light passing through the diaphragm 14 can pass through the scanning system. Subsequently, the pick-up device 16 makes a reliable measurement of the available light.

After exiting the conditioning stage 22, the light beam 6 reflects off the fold mirrors 28, 30 and 32. These mirrors may be adjusted for alignment purposes. The beam 6 is then incident on the beam combiner 34. The beam combiner 34 reflects the treatment beam 6 (and transmits the OCT 114 and aim 202 beams described below). For efficient operation of the beam combiner, the angle of incidence is preferably kept below 45 degrees and the polarization at which the beam is likely to be is fixed. The orientation of the linear polarizer 10 provides a fixed polarization for the treatment beam 6.

After the beam combiner 34, the beam 6 continues on the Z-adjust or Z-scan device 40. In this illustrative example, the z-adjustment comprises a Galilean telescope with two lens groups 42 and 44 (each lens group comprising one or more lenses). The lens group 42 moves along the z-axis about the alignment position of the telescope. Thus, the focal position of the spot in the patient's eye 68 is shifted along the z-axis as shown. Generally, there is a fixed linear relationship between the movement of the lens 42 and the movement of the focal point. In this case, the z-adjust telescope has approximately 2x beam expansion, and the movement of the lens 42 is 1: the relation 1. Alternatively, lens group 44 may be moved along the z-axis to facilitate z-adjustment and scanning. z-accommodation is a z-scan device used for treatment in the eye 68. Which may be controlled automatically and dynamically by the system and selected to be independent or to interact with the X-Y scanning device described below. Mirrors 36 and 38 may be used for aligning the optical axis with the axis of z-adjust device 40.

After passing through z-adjust device 40, beam 6 is directed by mirrors 46 and 48 to the x-y scanning device. Mirrors 46 and 48 can be adjusted for alignment purposes. The X-Y scanning is achieved by a scanning device 50, which scanning device 50 preferably uses two mirrors 52 and 54, which are rotated in a vertical direction using motors, galvanometers, or any other known optical movement device, under the control of an electronic controller 300. The mirrors 52 and 54 are positioned near the telecentric position of the objective lens 58 and contact lens 66 combination, as described below. These mirrors 52/54 are tilted so that they deflect the light beam 6, causing a lateral displacement in the plane of the treatment focus located in the patient's eye 68. Objective lens 58 may be a complex multi-element lens element, as shown, and is represented by lenses 60, 62, and 64. The complexity of the lens 58 will be dictated by the scan area size, the focal spot size, the available working distance on the proximal and distal sides of the objective lens 58, and the amount of aberration control. One example is a focal length of 60mm, operating an objective lens 58 with a 20mm diameter input beam size over a 7mm field of view. Alternatively, the X-Y scanning by scanner 50 may be accomplished by using one or more movable optical elements (e.g., lenses, gratings) that are controlled by electronic controller 300 via input and output devices 302.

The scanner 50 is capable of automatically generating aiming and treatment scan patterns under the control of the controller 300. Such patterns may include a single point of light, multiple points of light, a continuous light image, multiple continuous light images, and/or any combination thereof. Furthermore, the aiming mode (using aiming beam 202 described below) need not be the same as the treatment mode (using beam 6), but preferably is at least bounded so as to ensure that the treatment light is delivered only in the desired target area for patient safety reasons. This may be accomplished, for example, by having the targeting mode provide a profile of the desired treatment pattern. In this way, the user can be made aware of the spatial extent of the treatment pattern, even if the exact position of each focus point itself is not known, and thus optimize the speed, efficiency and accuracy of the scan. The aiming pattern may also be made to be perceived blinking in order to further enhance its visibility to the user.

An optical contact lens 66 may be used to help further focus the light beam 6 into the patient's eye 68 while at the same time helping to stabilize the eye position, the optical contact lens 66 may be any suitable ophthalmic lens. The positioning and characteristics of the beam 6 and/or the scanning pattern that the beam 6 forms on the eye 68 may further be controlled by using an input device such as a joystick or any other suitable user input device (e.g., GUI 304) to position the patient and/or the optical system.

The treatment laser 4 and controller 300 may be set to aim at the surface of a target structure in the eye 68 and ensure that the beam 6 is focused in the proper location and does not accidentally damage non-target tissue. The imaging modalities and techniques described herein, such as, for example, Optical Coherence Tomography (OCT), purkinje imaging, puru imaging, structured light illumination, confocal back-reflection imaging, fluorescence imaging, or ultrasound, may be used to determine position and measure the thickness of the lens and lens capsule to provide greater accuracy for laser focusing methods, including 2D and 3D patterning, or other known ophthalmic or medical imaging modalities and/or combinations thereof. In the embodiment of fig. 1, an OCT apparatus 100 is described. An OCT scan of the eye will provide information about the axial position of the anterior and posterior portions of the lens capsule, the boundaries of the cataract nucleus, and the depth of the anterior chamber. This information is then loaded into electronic controller 300 and used to program and control subsequent laser-assisted surgical procedures. This information may also be used to determine a number of parameters related to the surgical procedure, such as, for example, upper and lower limits of the focal plane for modifying the lens capsule, the cornea, and the synthetic intraocular lens implant, etc.

The OCT apparatus 100 in figure 1 includes a broadband or swept optical source 102 divided by a fiber coupler 104 into a reference arm 106 and a sampling arm 110. The reference arm 106 includes a module 108 that contains the reference reflection and appropriate dispersion and path length compensation. The sampling arm 110 of the OCT apparatus 100 has an output connector 112 that serves as an interface to the rest of the treatment laser system. The signals returned from reference arm 106 and sampling arm 110 are then directed by coupler 104 to detection device 128, which employs time domain, frequency domain, or single point detection techniques. In fig. 1, a frequency domain technique is used, with an OCT wavelength of 920nm and a bandwidth of 100 nm.

After exiting the connector 112, the OCT beam 114 is collimated using a lens 116. The size of the collimated beam 114 is determined by the focal length of the lens 116. The size of the beam 114 is determined by the desired NA at the focal point in the eye and the magnification of the beam train directed to the eye 68. Generally, in the focal plane, the OCT beam 114 need not have the same high NA as the treatment beam 6, so at the location of the beam combiner 34, the OCT beam 114 has a smaller diameter than the treatment beam 6. Following the collimating lens 116 is an aperture 118 that further modifies the composite NA of the OCT beam 114 at the eye. The diameter of the aperture 118 is selected to optimize the intensity of the OCT light and the return signal incident on the target tissue. Polarization control element 120, which may be active or dynamic, is used to compensate for polarization state changes that may be caused by individual differences in corneal birefringence, for example. Mirrors 122 and 124 are then used to direct the OCT beam 114 to beam combiners 126 and 34. Mirrors 122 and 124 can be adjusted for alignment purposes, and in particular for overlaying OCT beam 114 over treatment beam 6 after beam combiner 34. Similarly, a beam combiner 126 is used to combine the OCT beam 114 with the aim beam 202 described below.

Once combined with the treatment beam 6 after the beam combiner 34, the OCT beam 114 follows the same path as the treatment beam 6, through the rest of the system. Thus, the OCT beam 114 indicates the position of the treatment beam 6. The OCT beam 114 passes through the z-scan 40 and x-y scan 50 devices, then through the objective lens 58, contact lens 66, and into the eye 68. Reflections and scattering off structures inside the eye provide a return beam that is folded back through the optical system into connector 112, through coupler 104, and to OCT detector 128. These back reflections provide the OCT signal, which is then interpreted by the system as the position of the therapeutic beam 6 focal point at X, Y, Z.

The working principle of the OCT apparatus 100 is to measure the difference in optical path length between its reference arm and sampling arm. Thus, passing OCT through z-adjust 40 does not extend the z-range of OCT system 100 because the optical path length does not change with movement of 42. The OCT system 100 has an inherent z-range associated with the detection scheme, and in the case of frequency domain detection, it is particularly associated with the spectrometer and the position of the reference arm 106. In the case of the OCT system 100 used in FIG. 1, the z-range is approximately 1-2mm in a liquid phase environment. Extending this range to at least 4mm involves adjustment of the path length of the reference arm in OCT system 100. The z-scan in the sample arm to pass the OCT beam 114 through the z-adjust 40 allows for optimization of the OCT signal intensity. This is accomplished by focusing the OCT beam 114 on the target structure while accommodating the extended optical path length by matching increasing the path in the reference arm 106 of the OCT system 100.

Because of fundamental differences in OCT measurements with respect to the therapeutic focusing device due to aberrations such as immersion index, refraction, and color and monochrome, the OCT signal must be analyzed with the therapeutic beam focus position in consideration. A calibration or registration procedure should be implemented as a function of X, Y, Z in order to match the OCT signal information to the treatment focus position and to the absolute size quantities involved.

Observation of the aiming beam may also be used to assist the user in directing the therapeutic laser focus. Furthermore, assuming that the aiming beam accurately represents the infrared beam parameters, a macroscopic aiming beam may help with alignment instead of the infrared OCT and treatment beams. The targeting subsystem 200 is employed in the configuration shown in fig. 1. An aim beam 202 is generated by an aim beam light source 201, such as a helium-neon laser operating at a wavelength of 633 nm. Alternatively, laser diodes in the 630-650nm range may be used. An advantage of using a helium neon 633nm beam is its long coherence length, which would allow the use of the aiming path as a Laser Unequal Path Interferometer (LUPI) to measure the optical quality of the beam train, for example.

It should be noted that the therapeutic beam can be attenuated to the nano-focus level and used in place of the OCT system described above. This configuration provides the most direct relationship between the positioning of the focal positions for imaging and therapy (which are the same beam). The attenuated probe beam may be used directly in a back reflection measurement configuration or even indirectly in a fluorescence detection scheme. Both methods have advantages as one will see an increase in backscatter and fluorescence in the tissue structure. They can also be used to transmit sparse patterns to limit patient exposure while still identifying a reasonable mapping of intraocular targets.

Furthermore, because the precision and inclusion (inclusion) size required for lens accommodation is much less critical compared to corneal incisions and the lens capsule, the present invention contemplates adding a short pulse IR laser source to the above system for lens treatment, using a millifocal pulse from a Q-switched Nd: YAG laser for post-emulsification treatment as described above. This pulse energy will cause greater inclusion, which is not suitable for the robust isolation of the cataractous lens that the capsular and corneal incisions can provide. NIR wavelengths are not strongly absorbed or scattered by the lens, as opposed to short wavelengths. The second treatment source may have its beam combined with the first treatment beam by another beam splitter. The large difference in wavelength makes this a rather simple design. However, the same spectral difference would require a different registration for the imaging and/or ranging modality, as discussed above with respect to fig. 2B.

FIG. 4 is an illustration of a line spectrum of an OCT measurement across an axial section of the anterior capsule of the lens for eye 20. OCT imaging of the anterior capsule of the eye 20 can be performed along a simple linear scan across the lens using the same laser and/or the same scanner as the spectrum that produced the cut. This can provide information about the axial position of the anterior and posterior capsules of the lens, the boundaries of the cataract nucleus, and the depth of the anterior capsule. This information is then loaded into the laser scanning system and used to program and control the subsequent laser assisted surgery. This information may be used to determine various parameters about the procedure, such as, for example, upper and lower axial limits for cutting the lens capsule and segmenting the focal plane of the lens cortex and lens nucleus, the thickness of the lens capsule, and the like.

Fig. 5-9 illustrate various aspects of embodiments of the invention, which may be implemented using the scanning system 200 described above. As shown in fig. 5, capsulorhexis incision 400 (which may be created using system 200) is modified for an astigmatic intraocular lens (IOL). Such an astigmatic correcting IOL needs to be positioned not only in the correct location in the lens capsule 402 of the eye 20, but also oriented at the correct rotational/timing angle. Thus, they have an inherent rotational asymmetry, unlike spherical IOLs. The incision 400 shown in this example is elliptical; however other shapes are also useful. Incision 400 may be formed continuously or in sections to largely maintain the structural integrity of the lens capsule device of patient eye 20.

Such incomplete incisions 400 may be considered perforated incisions and have them removed gently so as to minimize their potential for inadvertent extension capsulorhexis. Regardless, incision 400 is a closed incision, which for purposes of this disclosure means that it begins and ends at the same location and surrounds a volume of tissue therein. The simplest example of a closed incision is a circular incision, in which the incision surrounds a disc of tissue. The next closed treatment pattern (i.e., the treatment pattern generated by the system 200 to form the closed incision) thus also starts and ends at the same location and defines a defined space surrounded thereby.

One key feature of closing incision 400 is that it includes registration features to orient the IOL to be placed therein. For the illustrated elliptical incision 400, its elliptical shape is its registration feature, which allows for precise placement of the IOL by virtue of its inherent rotational asymmetry, unlike the desired circular results of manual CCC. The major 404 and minor 406 axes of the ellipse of incision 400 are shown. The major axis 404 and the minor axis 406 are not equal. Incision 400 may be formed at any rotational angle relative to patient's eye 20, although it is shown in this example as being in a horizontal direction in the plane of the iris with its long axis 404. Incision 400 is intended to match one or more complementary registration features on the IOL. The surface of the lens capsule 402 to be cut can be precisely defined using the system 200. This may nominally serve to isolate the laser pulse from the vicinity of the target lens capsule 402 itself, thus minimizing the energy and treatment time required and, correspondingly, increasing patient safety and overall efficiency.

As shown in FIG. 6, IOL408 includes an optic portion 410 for focusing light and a haptic (haptic)416 for positioning IOL 408. Optical portion 410 is a rotationally asymmetric lens (about its optical axis) that includes an elliptical peripheral sidewall or edge 412, a complementary registration feature that mates with elliptical cutout 400. In this example, the elliptical edge 412 includes a major axis 418 and a minor axis 420. Major axis 418 and minor axis 420 are unequal. IOL408 further includes a surface 414 for retaining haptic elements 416 and providing a support platform for lens capsule 402 to secure optic 410 of intraocular lens 408 in the proper orientation and position within lens capsule 402 of eye 20 of a patient. Surface 414 is shown as being elliptical, but this is not required.

Haptics 416 provide stability and may be used to position edges 412 of intraocular lens 408 in incision 400 by applying a retaining force to the anterior portion of lens capsule 402. The haptics 416 may be arranged in any orientation. The direction of the cylindrical correction of optic 410 of intraocular lens 408 may be made to coincide with its major axis 418 or its minor axis 420. In this manner, IOL408 and optic 410 may be manufactured in a standard manner and the rotational orientation of incision 400 and the spherical and cylindrical optical power of optic 410 may be varied to conform to the individual optical prescription of patient's eye 20.

Figure 7 shows that once intraocular lens 408 is installed in lens capsule 402, matching the engaged registration feature edges 412 and incision 400, and on surface 414, intraocular lens 408 is properly momentarily deployed. Major axis 404 and major axis 418 are of unequal lengths. The minor axis 406 and the minor axis 420 are also unequal in length. This is provided to accommodate the fact that the lens capsule 402 may shrink slightly after capsulorhexis incision. The difference between the lengths of these axes tends to allow for the contraction of lens capsule 402 and still better position intraocular lens 408 in lens capsule 402 via incision 400. These differences should be limited to allow for reasonable foreshortening, but not so much as to allow significant rotation of intraocular lens 408. Typical values for these length differences may range from 100 μm to 500 μm, for example.

Fig. 8 shows a side view of the same intraocular lens 408 shown in fig. 6 and 7. In this schematic representation, edge 412 is shown on the same side of optic 410 as surface 424 of intraocular lens 408. Surface 422 on intraocular lens 408 serves to maintain the integrity of the fit between edge 412 and incision 400. The edge 412 is visible in the alternative views shown in fig. 6 and 7 as a projection of the surface 422. The optical axis 411 of the optic 410 is shown. In this view, pad 416 is positioned in line of sight.

Fig. 9 is a side view of the lens structure of fig. 8, but rotated 90 degrees to show that the display surface 426 is not curved in both directions (i.e., shaped as a cylindrical lens). This cylindrical or toric optical system of optic 410 provides a cylindrical correction of the patient's astigmatism. In this view, the contact 416 is positioned perpendicular to the line of sight.

As shown in fig. 15, the system can be used to alter the structure of, for example, corneal tissue without generating cavitation bubbles as shown in fig. 16, and these changes in corneal tissue can be used to shape the refractive index profile of the cornea 504 itself as illustrated in fig. 18. A number of small local changes 822 can be introduced into the cornea, which will change the refractive profile by changing the refractive index itself as well as the mechanical strength of the corneal tissue. It is thus possible to use not only the change in refractive index but also the change in corneal topography. This is achieved by tightly controlling the lateral spacing of the laser effect, which is achieved by the focusing unit 260 using the beam deflection unit 270 and the focus displacement unit 704.

As shown in the drawings for purposes of illustration, a method and system for making a physical degeneration (structural change) or incision in eye tissue is described. In various embodiments, the method and system provide a number of advantages over existing standards of care. In particular, a 320nm to 430nm laser is used to facilitate positioning and stability of the intraocular lens for a fast and precise opening in the lens capsule. But also the variation of the refractive power of the corneal tissue is achieved by locally changing the refractive index and reshaping the corneal topography.

Claims (10)

1. An ophthalmic surgical system, comprising:

a. a laser source configured to transmit a laser beam comprising a plurality of laser pulses having a wavelength between about 320 nanometers and about 430 nanometers and a pulse width between about 1 picosecond and about 100 nanoseconds; and

b. an optical system operatively coupled to the laser source and configured to focus and direct the laser beam in a pattern to one or more intraocular targets in the patient's eye such that interaction between the one or more targets and the laser pulse is characterized by linear absorption enhanced photolysis without formation of plasma or related cavitation events.

2. The system of claim 1, wherein the wavelength is about 355 nm.

3. The system of claim 1, wherein the pulse width is between about 400 picoseconds and about 700 picoseconds.

4. The system of claim 1, wherein the pulse has a pulse energy between about 0.01 microjoules and about 500 microjoules.

5. The system of claim 1, wherein the pulse has a pulse energy between about 0.5 microjoules and about 10 microjoules.

6. The system of claim 1, wherein the plurality of laser pulses have a repetition rate of between about 500 hertz and about 500 kilohertz.

7. The system of claim 1, wherein the optical system is configured to focus the laser beam to produce a beam diameter of between about 0.5 microns and about 10 microns in the one or more intraocular targets.

8. The system of claim 1, wherein at least one of the one or more intraocular targets is selected from the group consisting of a cornea, a capsular rim, a sclera, a lens capsule, a lens, and a synthetic intraocular lens implant.

9. The system of claim 8, wherein the pattern is configured to produce one or more incisions in an intraocular target in a manner selected from the group consisting of a keratotomy, limbal relaxing incision, astigmatic keratectomy, and capsulotomy.

10. The system of claim 1, wherein the optical system and laser source are configured to structurally alter at least one of the one or more intraocular targets such that a refractive index of the altered target changes.