JP2019520966A - Photo-adjustable hydrogel and bioanalogous intraocular lens - Google Patents

Abstract

A living body-similar implantable ophthalmic lens ("BIOL") that removes the natural lens (NCL), implants the BIOL into the posterior chamber, and places it in the capsular bag vacated by the NCL, followed by its various essentials. A bio-similar implantable ophthalmic lens that can replace NCL in terms of function. At least the rear surface of the lens has a convex shape and is made of a transparent and flexible hydrogel material. At least the front optical surface and the rear optical surface are defined by rotation of one or more conical sections along the main optical axis, the surface defined by the rotation being perpendicular to the axis and the conical surface symmetrical by the axis. A flat surface. A hydrogel implantable ophthalmic lens whose optical parameters can be optimized and / or customized by controlled absorption of electromagnetic radiation that will result in a change in the refractive index of the irradiated hydrogel.

Description

(Cross reference to related applications)
This application claims the benefit of priority of US Provisional Patent Application No. 61 / 752,685, filed January 15, 2013, the US domestic phase of International Application No. PCT / IB2013 / 060869, filed December 12, 2013. Claim priority to US patent application Ser. No. 15 / 190,715, filed Jun. 23, 2016, which is a continuation-in-part of US patent application Ser. No. 14 / 760,868 filed Jul. 14, 2015 . The entire contents of all of the above applications are incorporated herein by reference.

FIELD OF THE INVENTION The present invention is a hydrogel-embedded ophthalmic lens, the optical parameters of which are electromagnetic radiation such as laser light in the visible and / or near infrared region, and more particularly in pulses shorter than 1 nanosecond. The invention relates to an ophthalmic lens which can be optimized and / or customized by controlling the absorption of emitted radiation (so-called femtosecond laser, FSL) and changing the refractive index of the irradiated hydrogel .

BACKGROUND OF THE INVENTION Intraocular lenses (IOLs) are surgically implantable lenses that replace or supplement the optical function of NCL. The so-called "posterior intraocular lens", ie the PC IOL, has replaced the NCL by the so-called "clear lens exchange", ie CLE, in cases of cataract or more recently in the case of presbyopia. Other implantable lenses are placed in the anterior chamber (AC IOL), placed in the cornea (corneal implant or intrastromal implant), or placed between the NCL and the iris (so-called "embedded" Contact lenses "or ICL). So far, most of these IOLs have been designed simply to replace or supplement the basic optical functions of NCL. It can be seen that NCL in the human eye depicted in FIG. 1 is a complex structure with several functions. The main part of the eye is the cornea 101, iris 102, NCL 103, back capsule 104, ciliary muscle 105, ciliary zonule 106, vitreous body 107, and retina 108.

  The basic optical function of the NCL 103 is to assist the cornea 101 to focus incident light so that distant objects can be projected onto the retina 108. Another important optical function is accommodation, ie, adjusting the refractive power of the lens so that objects at different distances can be projected onto the retina 108. Several theories exist to explain accommodation mechanisms. For example, L. Werner et al., Physiology of Accommodation and Presbyopia, ARQ. BRAS. OFTALMOL. 63 (6), DEZEMBRO / 2000-503.

  The most rigidly established theory refers to FIG. 1, where the relaxed ciliary muscle 105 creates tension in the ciliary zonules 106 pulling the periphery of the lens 103 outward, making the NCL 103 suitable for far vision The von Helmholtz theory, which illustrates keeping the deformed (flat) shape providing even lower refractive power. Focusing on the near object relaxes the ciliary zonules 106 and NCL 103 in its "native" form with smaller diameter, thicker chunk center thickness and smaller radius of curvature in both the anterior and posterior surfaces It is induced by tension in the ciliary muscle 105 to be acquired. As a result, the refractive power of the NCL is enhanced, allowing the projection of the image of the near object on the retina 108.

  Most common intraocular lenses have a spherical surface that is rather easy to manufacture. For a while, it has been assumed that NCL 103 is required to be spherical. However, spherical lenses are not exactly single focus, but instead so-called "spherical" in which light rays incident through the center bend to a focal point slightly further from the lens than light rays incident through the lens perimeter "Aberration". Thus, a spherical lens is refracted somewhat more strongly at its periphery than at its center. This change is continuous and such a lens does not have a single focus but has many focuses at short distance intervals (focusing range) between the longest focal length and the shortest focal length. In other words, a spherical lens is negative multifocal (its focal length decreases from the center to the periphery). A lens that is not spherical but has an elliptical surface (such as a surface generated by solidification of a static liquid meniscus) exhibits stronger negative multifocality than a spherical lens because it has a clearer spherical aberration.

  Some artificial intraocular lenses include hyperboloids aligned with other quadratic surfaces such as spherical or even ellipsoidal surfaces with negative multifocality and very contrasting optical effects. More importantly, in the prior art, the quadratic surface (or conical cross section) is generally not clearly specified, and a quadratic surface with positive spherical aberration (but not a surface with hyperbolic aberration) Combined with a meniscoid surface that is only approximating.

  For example, in U.S. Pat. No. 4,971,732 Wichterle claims a meniscoid surface approximating a flat ellipse, and in U.S. Pat. No. 5,674,283 Stoy considers the meniscoid surface to be an approximation of a sphere, both of which are negative multifocality It is considered to have Combining the surfaces with positive and negative multifocality reduces or negates the benefits of surfaces with positive multifocality.

  In addition, U.S. Pat. No. 4,971,732 to Wichterle describes that the monomer solidifies in an open mold, one side of the lens (back) having the shape of the forming cavity and the front side solidifying (possibly negative) of the liquid meniscus We describe a method of making an intraocular lens that approximates a flat oval with multifocality and that has a shape that appears to be between a nearly pure sphere and a pure ellipsoid. The shaping cavity has the shape of a secondary surface that can include a hyperboloid. It can be appreciated that each optical surface is made in a different manner, one by solidification of the polymer precursor to the solid surface and the other by solidification at the liquid-vapor interface. It is known to those skilled in the art that the surface quality of two optical surfaces formed in such different environments may be quite different in both optical and biological aspects.

  In U.S. Pat. No. 4,846,832 Wichterle has the shape (possibly similar to a flat elliptical shape with negative multifocality) of the solidified liquid meniscus on the back side of the lens and the front side on It describes another method of manufacturing an intraocular lens formed as an imprint of a solid mold formed as a secondary surface that may also implicitly include a hyperboloid. Again, it can be seen that each optical surface is created in a different manner, one by solidification of the polymer precursor to the solid surface and the other by solidification at the liquid-vapor interface.

  U.S. Pat. No. 5,674,283 to Stoy is a two-part mold in which one part is similar to the Wichterle's type and the other part is used to form a modified meniscoid of shorter diameter on the anterior lens surface. A modification of the method described by Wichterle US Pat. No. 4,971,732 used is disclosed. The meniscoid optical surface, despite its short diameter, has the same characteristics as the meniscoid obtained from Wichterle, US Pat. No. 4,971,732, and thus probably seems to be more spherical than an ellipsoid. In any case, such surfaces have negative multifocality. The rear side is formed as a solid-type imprint shaped as a secondary side which may include hyperboloids, the other optical side is formed by solidification of the liquid polymer precursor at the liquid-vapor interface.

    In the international application CZ 2005/000093, Michalek and Vacik describe a method of making an IOL using an open-type rotomolding method. The mold filled with the monomer mixture rotates along its vertical axis while the polymerization proceeds. One of the optical surfaces is made as an imprint of the solid mold surface and the other is formed by rotation of the mold. The imprinted surface has a shape (which may include a hyperboloid shape) formed by rotation of the conical cross-section along the vertical axis. On the other side, it is shaped as a meniscoid corrected by centrifugal force that moves a portion of the liquid precursor from the center to the periphery. In the case of a convex meniscus, the centrifugal force flattens the center and creates a sharp curvature at the periphery, ie increases the spherical aberration of the surface. In the case of a convex meniscus, the centrifugal force will create a meniscus with a shorter central radius and will modify the surface to approximate the shape between spherical and parabolic. In any case, hyperbolic aberration can not be achieved for either the convex or concave meniscoid surface.

  In U.S. Pat. Nos. 4,994,083 and 4,955,903, Sulc et al disclose an intraocular lens whose front surface projects forward so as to be in permanent contact with the iris which is to be centered on the lens. There is. The rear and front surfaces may both have shapes (spherical, parabolic, hyperbolic, elliptical) obtained by rotation of the conical section around the optical axis. The iris contact portion of the lens is a hydrogel that is soft and deformable in nature, having a very high water content (at least 70% and preferably more than 90% water). Thus, the optical surface deformed by contact with the iris may be a surface that may not be exactly conical in cross section, but perhaps has a somewhat smaller central radius close to a spherical shape, which may vary with the pupil diameter. That is, this situation is similar to the lens according to another reference which achieves a reduction in central diameter by pressing the deformable gel-filled lens against the pupil-like opening in the iris-like artificial element (Nun, U.S. Patent No. 7220279). U.S. Pat. No. 7,220,279 to Nun does not mention or imply the use of hyperboloidal optical surfaces. In U.S. Patent Application Publication Nos. 2007/0129800 and 2008/0269887, Cummings is a hydraulically controlled IOL that results in optical surface change and accommodation by pushing liquid into the internal IOL chamber by the action of the ciliary body device. Is disclosed.

  In U.S. Pat. No. 7,350,916 and U.S. Patent Application Publication No. 2006/0244904 Hong et al. Disclose an aspheric intraocular lens having at least one optical surface having spherical aberration to compensate for positive spherical aberration of the cornea. ing. Negative spherical aberration is achieved by the hyperboloid shape of the optical surface.

  Hong et al. In US Patent Application Publication No. 2006/0222786 disclose an IOL shape factor that is optimal for the human eye, and a certain range of “shape factor” of −0.5 to +4 (this shape factor Defined by Hong as the ratio of their sum to the difference of the back and the back curve), and preferably at least one of the optical surfaces has a conic constant of -76 to -27. It is a sphere.

  Hong et al. In U.S. Pat. No. 7,350,916 describe an IOL having at least one optical surface with negative spherical aberration in the range of about -0.202 microns to about -0.190 microns over the entire output range.

U.S. Pat. No. 4,971,732 U.S. Pat. No. 5,674,283 U.S. Pat. No. 4,846,832 International application CZ2005 / 000093 U.S. Pat. No. 4,994,083 U.S. Pat. No. 4,955,903 U.S. Pat. No. 7,220,279 US Patent Application Publication No. 2007/0129800 U.S. Patent Application Publication No. 2008/0269887 U.S. Pat. No. 7,350,916 U.S. Patent Application Publication No. 2006/0244904 US Patent Application Publication No. 2006/0227286

L. Werner et al., Physiology of Accommodation and Presbyopia, ARQ. BRAS. OFTALMOL. 63 (6), DEZEMBRO / 2000-503

  In at least one embodiment, the present invention is an artificial lens that can be implanted in the posterior chamber of a human eye to replace a natural lens, comprising a main optical axis 1A, a central optical part 2 and a peripheral support part 3 And the overall shape of the implant is defined by the implant's anterior surface 4, posterior surface 5 and the transition surface 6 between the upper boundary 7A of the anterior surface and the upper boundary 7B of the posterior surface, with the boundary 9A and the anterior apex 10A An artificial lens (see FIG. 3) having a central anterior optical surface 8A, a central posterior optical surface 8B having a boundary 9B and a posterior apex 10B, and a front peripheral support surface 11A and a rear peripheral support surface 11B is provided.

  The back of the human eye to replace the innate lens, which reproduces the shape, size, optical properties and material properties of the NCL as closely as practicable, taking into account the need for surgical implantation by making small incisions An artificial lens that can be embedded in a tuft.

  The artificial lens according to at least one embodiment of the present invention has at least a posterior surface that approximates the shape and size of the posterior surface of the native lens to achieve substantially complete contact with the posterior capsule of the eye. As defined in this context, the term "substantially" means that at least about 90% of the posterior BAIOL surface is in contact with the posterior capsule, or at least about 75 of the posterior capsule (having a larger diameter than the lens). It can mean that% is in contact with the rear surface of the lens. At least that part of the artificial lens according to the invention in contact with the posterior capsule is made of a transparent, flexible hydrogel material which approximates the optical, hydrophilic and electrochemical properties of the tissue forming the natural lens. The anterior side is designed to avoid permanent contact with the iris.

In at least one embodiment, the anterior surface is shaped to avoid permanent contact with the iris with the anterior peripheral support surface 11A being concave.
In at least one embodiment, an artificial lens according to the present invention is a liquid polymer precursor defined by rotation of one or more conical cross sections along the optical axis and in contact with a solid wall of a mold, preferably a hydrophobic plastic type It has at least a major portion of its front and back surfaces, including both optical surfaces, formed by the solidification of the body.

  One embodiment of the invention relates to a hydrogel comprising a UV absorbing dopant moiety and an activator moiety which is negatively charged at physiological pH, wherein the hydrogel in the fully hydrated state is exposed to electromagnetic radiation This results in one or more structural changes in the hydrogel and two-photon absorption which induces a change in refractive index. In one embodiment, the hydrogel is a covalently cross-linked hydrogel. In one embodiment, structural changes are achieved without substantially changing the volume of the treated hydrogel segment after reaching equilibrium with the liquid medium around the lens. One convenient method of determining volume change can be configured by the following procedure: Several equal areas in a hydrogel (e.g. 50 x 50 microns) are treated with different laser settings to vary from area to area Achieve phase shift. After a sufficient time to reach equilibrium, the linear dimensions of the treatment area do not change more than 20% from the original area dimensions and the dimensional change is less than 10% for most conditions. Since the depth of the treatment area can not be measured easily or directly, the volume change is isotropic, and it is assumed that the change in treatment volume is due to the same relative expansion or contraction for all dimensions. One embodiment of the hydrogel comprises a polymer comprising (meth) acrylic acid derivatives and / or monomer units of (meth) acrylic acid. In one embodiment of the hydrogel, the dopant and activator moieties are pendant groups of a polyacrylate polymer or polymethacrylate polymer. In one embodiment, the dopant moiety is a UV absorbing compound that does not strongly absorb light of about 400 nm wavelength. In one embodiment, the dopant moiety is a compound selected from the group consisting of rhodamine, benzophenone, coumarin, fluorescein, benzotriazole and derivatives thereof. In one embodiment, the UV absorbing moiety is a carbonyl group conjugated to an aromatic system, a phenolic hydroxyl group conjugated to an aromatic system, or preferably a carbonyl group conjugated to an aromatic system and a phenolic hydroxyl Includes both of the groups. In one embodiment, the activator moiety is a compound comprising a carboxylate group, a sulfonate group, a sulfate group, a phenolate group or a phosphate group. In one embodiment of the hydrogel, the one or more structural changes comprise partial depolymerization of the hydrogel. In one embodiment, partial depolymerization results in the formation of moisture-filled voids in the hydrogel. In one embodiment of the hydrogel, the change in refractive index is a negative change. In one embodiment, the depth of partial depolymerization of the hydrogel is dependent on the accumulated energy absorbed at the predetermined position of the hydrogel. In one embodiment, the hydrogel comprises a polymer comprising monomer units selected from the group consisting of acrylic acid derivatives, methacrylic acid derivatives, acrylic acid, methacrylic acid and mixtures of two or more thereof.

  Another embodiment of the present invention relates to an ophthalmic implant comprising the above hydrogel.

  Yet another embodiment of the present invention is an in situ controlled hydrogel ophthalmic implant comprising an acrylate copolymer hydrogel or a methacrylate copolymer hydrogel, said copolymer comprising at least four comonomers: a) at least one pendant hydroxyl An acrylic or methacrylic ester containing a group; b) a polyol acrylic ester or polyol methacrylic ester or amide with at least two acrylate or methacrylate groups per polyol ester or amide; c) at least one pendant carboxyl group Derivatives of acrylic acid or methacrylic acid having the following; and d) vinyl monomers having pendant UV absorbing groups, acrylic monomers and An ophthalmic implant comprising a methacrylic monomer, wherein the refractive properties of the implant are adjusted by controlling the absorption of target electromagnetic radiation by the hydrogel to change the refractive index at selected locations of the implant . In one embodiment, the ester of component a) has at least one pendant hydroxyl group in the alcohol part of the ester. In one embodiment, the copolymer is covalently crosslinked. In one embodiment, the implant has anterior and posterior refractive surfaces that form a lens with positive or negative refractive power. In one embodiment, the lens is an intrasubstrate lens. In another embodiment, the lens is an anterior chamber lens. In yet another embodiment, the lens is a phakic lens for placement between the iris and the native lens. In yet another embodiment, the lens is a posterior chamber lens for at least partially replacing the native lens. In one embodiment of the hydrogel ophthalmic implant, at least one of the refractive surfaces is a non-spherical surface with negative spherical aberration. In another embodiment of the hydrogel ophthalmic implant, the monomer comprising pendent carboxyl groups is neutralized or partially neutralized methacrylic acid. In one embodiment, the monomers comprising pendent carboxyl groups are present at a concentration of 0.1 mole% to 5 mole% based on total monomer units of the copolymer. In another embodiment, monomers comprising pendent carboxyl groups are present at a concentration of 0.5 mole% to 2 mole% based on total monomer units of the copolymer. In one embodiment, the monomer comprising the pendant UV absorbing group is present at a concentration of 0.1 mole% to 5 mole% based on total monomer units of the copolymer. In another embodiment, the monomer comprising pendant UV absorbing groups is present at a concentration of 0.2 mole% to 2.5 mole% based on total monomer units of the copolymer. In one embodiment of the hydrogel ophthalmic implant, the pendant UV absorbing group comprises a carbonyl group conjugated to an aromatic group, and in one embodiment, the monomer comprising the pendant UV absorbing group comprises all the monomer units of the copolymer It is present at a concentration of 0.1 mol% to 5 mol% based. In one embodiment of the hydrogel ophthalmic implant, at least one of the UV absorbing pendent groups is selected from the group consisting of derivatives of benzophenone, derivatives of benzotriazole, derivatives of coumarin and derivatives of fluorescein. In one embodiment, the pendent carboxyl group and the pendent UV absorbing group are present in a molar ratio of about 0.25-5, and in another embodiment the pendent UV absorbing group is about 0.5 to about 3. Present in a molar ratio of 5. In one embodiment of the hydrogel ophthalmic implant, the copolymer comprises at least two different comonomers comprising different UV absorbing groups, and in one embodiment at least one of the UV absorbing groups is benzophenone.

  In one embodiment of the ophthalmic lens, the pendant carboxyl groups are ionized and the molar ratio of ionized pendant carboxyl groups to UV absorbing pendant groups is about 0.5 to about 3.5. In another embodiment of the ophthalmic lens, at least the main part of the hydrogel polymer is a derivative of methacrylic acid, and in one embodiment at least the main part of the methacrylic acid derivative is a hydrophilic derivative of methacrylic acid. In one embodiment, the hydrophilic methacrylic acid derivative is a glycol ester of methacrylic acid. In one embodiment of the ophthalmic lens, the covalently cross-linked hydrogel comprises a liquid in a proportion greater than 30% by weight under equilibrium physiological conditions. In one embodiment, the covalently cross-linked hydrogel comprises liquid at a proportion of less than 55% by weight under equilibrium physiological conditions. In another embodiment, the covalently cross-linked hydrogel comprises liquid in a proportion of 35% to 47.5% by weight under equilibrium physiological conditions. In one embodiment of the ophthalmic lens, at least the posterior optical surface provides refraction at negative spherical aberration. In another embodiment, the posterior optical surface provides refraction at negative spherical aberration of -0.1 microns to -2 microns. In a further embodiment, the negative spherical aberration is -0.5 microns to -1.5 microns. Alternatively, the negative spherical aberration is -0.75 microns to -1.25 microns. One embodiment of an ophthalmic lens comprises both a UV absorbing pendant group comprising a carbonyl group conjugated to an aromatic system and a UV absorbing group comprising a benzotriazole structure. In one embodiment, a UV absorbing pendant group comprising a carbonyl group conjugated to an aromatic system and a UV absorbing group comprising a benzotriazole structure are located in separate layers of an ophthalmic lens. In one embodiment, an ophthalmic lens is implanted in the cornea. In another embodiment, the lens is implanted in the anterior chamber of the eye between the cornea and the iris. In yet another embodiment, the lens is a phasic lens embedded between the iris and the native lens. In yet another embodiment, the lens is implanted in the posterior chamber of the eye and replaced, at least in part, with a native lens.

  Another embodiment of the present invention is a method of modulating the refractive properties of the fully hydrated hydrogel of the present invention, comprising intensively irradiating the hydrogel with electromagnetic radiation such that two-photon absorption occurs. The present invention relates to a method in which partial decomposition polymerization and / or decomposition is carried out on the polymer component of hydrogel, and a part of the polymer scaffold is effectively removed to create a void.

  Another embodiment of the present invention is a method for the in situ adjustment of optical parameters of a hydrogel ophthalmic implant, comprising the steps of: a) including the hydrogel ophthalmic implant according to claim 13 in the eye, and b) Irradiating a portion of the hydrogel ophthalmic implant with electromagnetic radiation using a femtosecond laser, thereby depolymerizing and / or ablating a portion of the hydrogel copolymer and adjusting the optical parameters of the implant It is about In one embodiment of the method, the optical parameter comprises a refractive index. In one embodiment of the method, the irradiation results in an elongated cavity or voxel inside the hydrogel ophthalmic implant. In some embodiments, the voxel depth is approximately 20-30 microns or more. In one embodiment, increasing the depth of the voxels also increases the phase shift, but the refractive index remains approximately constant, and in one embodiment the refractive index is 1.3335 or greater. In some embodiments of the method, the phase shift is up to three wavelengths of green light. In another embodiment of the method, the depolymerized material that has been ablated by irradiation comprises a soluble, easily diffusing, low toxicity compound. The low toxicity of the method is also evident, as the procedure only releases very low concentrations of the depolymerised compound into the intraocular space. In one embodiment of the method, the modified optical properties are provided by forming in the hydrogel a pattern of elongated voxels of various depths while keeping the modified refractive index substantially constant. In one embodiment of the method, the phase shift is controlled by changing the voxel depth rather than by changing the refractive index.

BRIEF DESCRIPTION OF THE DRAWINGS The accompanying drawings, which are incorporated in and constitute a part of this specification by reference, illustrate presently preferred embodiments of the invention, and together with the summary and the detailed description given above. , Serve to explain the features of the present invention. In the drawings:
FIG. 1 illustrates the internal placement of the eye with major structures including the cornea, sclera, iris, NCL, vitreous, retina and lens suspension devices (capsule, zonules and ciliary muscle). FIG. 2 illustrates the distribution of refractive power in a lens having one hyperboloid. FIG. 3A is a cross-sectional view of a biosimilar intraocular lens according to an exemplary embodiment of the present invention. FIG. 3B is a top view of the lens of FIG. 3A. FIG. 4A is a top view of another exemplary embodiment of a lens having circular optics and an oval support. FIG. 4B is a top view of another exemplary embodiment of a lens with a circular support truncated at a straight cut. FIG. 4C is a top view of another exemplary embodiment of a lens with a circular support tipped off by two symmetrical crescent-shaped cuts. FIG. 4D is a top view of another exemplary embodiment of a lens with a circular support tipped off with a straight cut and two crescent cuts. FIG. 4E is a top view of another exemplary embodiment of a lens with a circular support tipped off with four symmetrical crescent-shaped cuts. FIG. 4F is a top view of another exemplary embodiment of a lens with a circular support tipped off by two straight parallel cuts and a cylindrical lens with a cylindrical axis 1B at an angle α with respect to the cutting direction It is. FIG. 5A is a top view of an exemplary lens in which the optical surface is divided into two or more optical areas. FIG. 5B is a top view of an exemplary lens in which the optical surface is divided into two or more optical areas. FIG. 5C is a top view of an exemplary lens in which the optical surface is divided into two or more optical areas. FIG. 6A is a cross-sectional view of an alternative lens according to the present invention composed of two or more materials. FIG. 6B is a cross-sectional view of an alternative lens according to the present invention composed of two or more materials. FIG. 6C is a cross-sectional view of an alternative lens according to the present invention composed of two or more materials. FIG. 7A is an enlarged view that reveals alternative aspects of the support perimeter of a typical lens. FIG. 7B is an enlarged view that reveals alternative aspects of the support perimeter of a typical lens. FIG. 7C is an enlarged view that reveals alternative aspects of the support perimeter of a typical lens. FIG. 8 illustrates a schematic arrangement of molds for manufacturing a lens according to an exemplary embodiment of the present invention. FIG. 9A shows a comparison of Raman spectra of representative hydrogels of the invention before and after two-photon absorption (TPA) using a femtosecond laser. FIG. 9B shows a graph of the relevant parameters of the Raman spectrum. FIG. 9C shows a graph of the relevant parameters of the Raman spectrum.

DETAILED DESCRIPTION OF THE INVENTION In the present application, the terms "(meth) acrylic ((meth) acrylic)" and "(meth) acrylate" are any of acrylic / acrylate or methacrylic / methacrylate. Indicate In some embodiments, the polymer comprises an acrylic / acrylate moiety. In another embodiment, the polymer comprises methacrylic / methacrylate moieties. In yet another embodiment, the polymer comprises both acrylic / acrylate and methacrylic / methacrylate moieties. In a preferred embodiment, the polymer comprises methacrylic / methacrylate moieties.

  Also, as used herein, the term "ablation" refers to the removal of fragments of the polymeric structural support of the hydrogel, preferably by degradation of the polymer to diffusive low molecular weight fragments. The term "depolymerization" constitutes a special type of ablation in which the fragments are monomers.

  There are many types of implantable ophthalmic lenses used at various locations in the eye, from the corneal stroma, through the anterior chamber to the posterior chamber. One of the problems with implanted lenses is that the choice of the correct refractive properties (so-called biometry) is complicated and the results of the biometry are found to be false or the optical requirements of the eye are It is difficult to replace or correct the refractive properties if they change over time. This problem has triggered the industry to work towards the development of implantable lenses whose optical parameters can be adjusted non-invasively and after surgery. Changes in the optical properties of the artificial lens either before or after surgery can be achieved by changing the refractive index of the lens material.

  Changes in refractive index due to the action of two-photon absorption (TPA) or multi-photon absorption (MPA) for implantable lenses have been reported in the prior art as positive for hydrophilic polymers and negative for hydrophobic polymers. This is consistent with a mechanism that is presumed to be accompanied by a decrease in water content as a result of the increased crosslink density in hydrophilic acrylates and hydrogels, and conversely, an increase in hydrophilicity in hydrophobic acrylates. The cause of these changes is presumed to be a local temperature rise.

  Changes in the optical properties of innate (eg, human cornea) or artificial lenses by altering the refractive index of their materials have also been reported in numerous articles, patents and patent applications, such as: Phillips, A . J. , System and Method for Treatment of Hyperopia and Myopia, U.S. Patent No. 6,102,906, Bille J. et al. F. U.S. Pat. Nos. 8,292,952; 8,920,690; 9,192,292; Sahler; Ruth et al .: "Hydrophilicity alteration system" U.S. Pat. Nos. 9,023,257; 9,186,242 and 9,107,746; Sahler; Ruth et al .: "intraocular lens (IOL) fabrication system and method" US patent application Publication No. 20160074967; Smith, T. et al. U.S. Pat. Optical Material And Method For Modifying The Refractory Index, U.S. Patent Nos. 8,932,352; 8,337,553; 7,789,910 B2; Knox, Wayne H. et al. U.S. Patent Application Publication No. 20130138093; U.S. Patent No. 20130178934; U.S. Patent No. 201200298933; U.S. Patent No. 20080001320; U.S. Patent No. 20090143858; Et al: Method For Modifying The Refractive Index Of An Optical Material And Resulting Optical Vision Component, U.S. Patent Application Publication No. 20120310340, Knox, Wayne H. et al. Et al: Method For Modifying Refractive Index Of Ocular Tissues, U.S. Patent Nos. 8,486,055; 8,512,320; 8,617,147; No. 20140107632; Knox, Wayne H. et al. Et al: Methods For Modifying Refractive Index of Ocular Tissues And Applications Thereof, U.S. Patent Application Publication No. 20120230223, each incorporated herein by reference. None of these prior art references suggest ablation or depolymerization as an optical control mechanism in hydrogels.

  Applicant's co-pending international application, PCT / IB2016 / 052487, filed May 2, 2016, Method and Device for Optimizing Vision Via Customization of Spherical Aberration of Eye (optimization of visual acuity by customizing spherical aberration of the eye) The method and apparatus for) includes related disclosure.

  The wavelength of the laser beam is usually in the near infrared, in the range of about 800 nm to 1300 nm, and more typically in the visible and near infrared range of about 660 nm to about 1100 nm. It is usually preferred to use higher wavelengths, due to safety concerns. One method of changing the refractive index of a lens material by femtosecond laser means (abbreviated "FSL procedure") is, in principle, spherical refractive power, cylindrical refractive power, spherical aberration, etc. Many changes in refractive properties can be achieved. In principle, this procedure can selectively change any of the coefficients in the Zernike polynomial ("Zernike coefficients") repeatedly even in the embedded lens. This is described by Gustavo A. Gandara-Montano et al. "Femtosecond laser writing of freeform gradients in hydrogel-based contact lenses", October 1, 2015 | Volume 5, No. 10 | 5. 002257 | Certified by OPTICAL MATERIALS EXPRESS 2257. The authors noted negative phase shifts in ETAFILCON® contact lens hydrogels when treated with a femtosecond laser at 800 nm. However, none of these systems has been used clinically, especially because the change in refractive index by using the FSL procedure is unexpectedly small for currently available ophthalmic lens materials.

  The FSL procedure can be performed on both hydrophilic IOL materials (refractive index RI is usually enhanced) and hydrophobic IOL materials (RI is usually reduced).

  Even in the currently available hydrogels, the change in refraction is higher than in the case of hydrophobic materials, so the FSL procedure is more advantageously implemented for hydrogel ophthalmic lenses, in particular for various types of implantable lenses obtain. Furthermore, the FSL procedure in hydrophobic acrylates can lead to other problems associated with the formation of so-called "glistening" or hydrophilic "penetration cells" in hydrophobic materials, at least in theory. Therefore, various hydrogels were tested as substrates for microfabrication with femtosecond laser. The sensitivity of hydrogels can be enhanced by various "dopants" designed to respond to electromagnetic radiation of various wavelengths. So far, several dopants have been described in the patent and scientific literature. All currently known dopants are single compounds capable of UV absorption, but none of the known dopants are capable of one-photon absorption at the wavelengths used for FSL procedures.

  We have added dopants for TPA and MPA in acrylic and methacrylic hydrogels by adding certain TPA activators, such as comonomers containing negatively charged pendant groups, in particular organic carboxylates. It has been found that the effect can be improved, corrected and enhanced. The dopant and its activator may both be covalently linked, preferably to the polymer chain, more preferably to the chain forming the polymer network of the hydrogel. The dopant and its activator may both be attached to the same polymer chain or to different polymer chains.

  More specifically, the hydrogel of the invention comprises a minor portion of an activator in the form of a pendant group comprising an ionized salt of an acid, a major portion of a methacrylate neutral hydrophilic derivative, in particular a glycol or glycerol ester of methacrylic acid and acrylic acid or Together with the minor part of the polyol with at least two of the hydroxyl groups esterified by methacrylic acid, it comprises a combined dopant in the form of a minor part of a pendant UV absorbing structure which absorbs less visible light. These hydrogels are materials particularly suitable for adjustment of refractive index by absorption of visible or near infrared femtosecond pulses (femtosecond laser (FSL) processing). Absorption of electromagnetic energy induces controlled degradation and depolymerization of the polymeric components of the hydrogel, thereby forming domains of reduced refractive index. By virtue of this activator, the domain can be formed even when the hydrogel is irradiated with a femtosecond laser at a very high "writing speed" or "scanning speed" with relatively low energy pulses.

  As used herein, a covalently cross-linked bioanalogic containing a negatively charged group, in particular a carboxylate group, and also including a monomer with a pendant UV absorbing group such as, for example, methacryloyloxybenzophenone (MOBP) Discloses acrylic and methacrylic hydrogels. Also herein, hydrogels that can be implanted using a femtosecond laser to depolymerize a portion of the hydrogel, such as a toric lens, to form an internal cavity that will form a new refractive member Discloses a method of depolymerizing said polymer by absorption of electromagnetic radiation to adjust the optical properties of said polymer. Implicitly, but certainly, such newly formed cavities in the hydrogel matrix are necessarily filled with water or an aqueous fluid (including the residue of the degraded polymer component alone or in the hydrogel). Thus, it will have a different (especially lower) refractive index than the parent hydrogel material. It also discloses the possibility of performing this in situ lens modulation in the implant, as such degradation products are water soluble compounds that are low toxic and can diffuse slowly through the hydrogel.

  We note that the UV absorbing groups used in the hydrogel act as "dopants" to enhance the rate of depolymerization, while negatively charged groups further to the efficiency of the dopant " It confirmed that it acted as "activator". In addition, activators that act as quenchers protect the material from the undesirable "charring" or "burning" that can occur to a certain amount of energy absorbed per dose. This reduces the overall amount of energy required to achieve a certain refractive change, so that high-energy visible light can be safely used rather than low-energy near infrared (NIR) light.

  As not intended to limit the scope of the present invention by any particular hypothesis or theory, the following is a proposed explanation of the phenomenon of hydrogel depolymerization by electromagnetic radiation. Focused femtosecond lasers are known to promote two-photon absorption (TPA) (or even multiphoton absorption (MPA)) in their focal volume. The volume of material that TPA influences is usually referred to as "volume pixel" or "voxel". The size of the voxels varies with different parameters and generally increases with the absorbed energy. The smallest possible voxel size corresponds to an "collective volume" or ellipsoid with a volume that is approximately cubic of the wavelength of the absorbed laser beam. The diameter of the voxel is considerably shorter than its depth, typically around 500 nm, thanks to the self-focusing of the laser beam in the case of TPA or MPA. The voxel depth is quite large and increases with the stored energy. The maximum voxel depth reported is about 6 microns for currently known materials.

  When radiation of a certain wavelength is absorbed through TPA or MPA, the dopant will store excitation energy that is equivalent to the wavelength of the incident light divided by the number of photons absorbed and one photon absorption. For example, laser light at 400 nm absorbed by TPA through the dopant will correspond to 200 nm light absorbed through SPA. The 200 nm wavelength light is hard UV light (UVC band) with sufficient energy to induce chemical bond breakdown and deep structural rearrangement. Of course, three-photon absorption or four-photon absorption, although much less likely than TPA, will accumulate to much higher energy concentrations. Because energy dissipates rapidly through one of several possible paths (most commonly the conversion to thermal energy), highly energy stored excited states are ephemeral. The usual effects of femtosecond laser treatment of tissue or synthetic hydrogels depend on the amount of energy absorbed, but conversion of the object to plasma by heating sufficient to cause charring (eg, via laser ablation) Starting with, with a local temperature increase that can trigger an effect leading to a more elaborate additional crosslinking via various mechanisms such as re-esterification, disproportionation or dehydration with the formation of ether bonds.

  However, the mechanism of this process is altered by the presence of activators (negatively charged pendant groups, in particular carboxylate groups). We believe that this is triggered by some co-operation between the activator group and the dopant. In this cooperation, activation of the dopant further promotes TPA absorption and its effect is increased. Activators such as carboxylic acid groups of methacrylic acid or its salts appear to be in an interactive relationship with the dopant, which, due to the presence of the activator, changes the UV / visible region spectrum of the dopant (ie UV This is because a subtle shift (from a region to a visible region) is induced. Thereby, the "TPA cross section" of the dopant can be increased and the absorption efficiency of the dopant can be enhanced.

  However, the main role of the activator is to direct the absorbed energy of the two (or more in the case of MPA) absorbed photons from the dopant to the main polymer chain, resulting in a covalent bond in the main chain It can trigger a cleavage. From the general characteristics of this process, and from the known depolymerization kinetics of methacrylate polymers, this cleavage is homopolar and free radical that initiates the depolymerization process in the vicinity of negatively charged activators (eg carboxylates) Give rise to This free radical mechanism can greatly increase the quantum yield of photolysis, inducing large-scale changes in local composition and structure even with relatively low concentrations of dopant (on a molar basis), resulting in local refractive index It helps to explain why it can trigger considerable change.

  In addition, the activator group diverts the energy absorbed from the dopant to quench the excited state, helping to "consume" this energy for the cleavage of covalent bonds. This "energy conduit" function of the activator will "recycle" the dopant molecule and help preserve it for future TPA cycles in order to prevent excessive heat build up in the vicinity of the dopant moiety. This function also helps to protect the polymer structure from burning and charring, and helps to increase the efficiency of the entire absorption process.

  Furthermore, this proposed mechanism is such that in the absence of activator, TPA in hydrogels with only dopant proceeds by a different mechanism, the refractive index decreases in the presence of activator and the refractive index in the absence of We can explain the fact that the opposite result of increasing can be obtained. One possible reason for the increased refractive index is that water has the lowest refractive index of all the hydrogel components, so that the water content decreases and the further crosslinking which results in an increase in refractive index It appears to be.

  Another consequence of the stated mechanism, which entails an activator, is that the TPA process reduces (or even eliminates) the volume change. That is, if additional crosslinking takes place, the water content is reduced while the amount of polymer component remains approximately the same. Thus, the mass and volume of the amount of hydrogel treated must be reduced, with various deleterious effects on the product, such as the occurrence of internal stresses, changes in external morphology and changes in mechanical properties. Conversely, if the polymer becomes more hydrophilic and attracts more water, the processing volume of the polymer must be expanded (again, the detrimental effects on the external morphology, optical properties and stress in the material) Accompany).

  The mechanism specified in the present invention entails exchange of water with some parts of the polymer mass, albeit with little, if any, change in actual volume. Depolymerization of the hydrogels encompassed by the present invention results in low toxic monomers and / or their fragments, mainly 2-hydroxyethyl methacrylate, methacrylic acid and ethylene glycol. All these degradation products are well soluble in water, can diffuse through the surrounding intact hydrogel network, and are released from the implant over time at very low concentrations.

  The proposed mechanism may explain the very large phase shift achievable in the hydrogel according to the invention. The phase shift is determined by the change in refractive index and the length of the light path through the material with the changed refractive index, in other words the voxel depth. The refractive index in the hydrogel can not be substantially lower than the refractive index of water, or about 1.3335, as the lowest refractive index conceivable can be achieved in an aqueous liquid-filled cavity. The cavities in the hydrophilic polymer matrix can not be filled with gas for some basic thermodynamic reasons, and there is no known or easily imagined mechanism to create hydrophobic cavities in the hydrogel.

  Thus, a large phase shift (more than one wavelength) will require the formation of voxels with a large depth. It is usually estimated that the voxel depth can reach at most about 5-6 microns. However, if the refractive index in the collection volume decreases, the elongated voxels will act as a light guide to guide additional light pulses to be absorbed at the bottom of the voxels. Thus, the depth of voxels may increase gradually with increasing number of pulses, and even if the index of refraction is approximately constant, remaining equal to or higher than 1.3335, the phase shift will increase as well. Thus, this mechanism is different from previously reported mechanisms in which the voxel size remains nearly constant but the refractive index change increases with increasing energy absorbed (eg the number of pulses achieved). The presently proposed mechanism is believed to be able to provide voxel depths greater than 10 microns, and phase shifts corresponding to voxel depths as high as 25 or 30 microns have been observed. This is not a limitation on the present invention, but is stated to demonstrate the fundamental differences of the present invention with the prior art. One of the practical consequences of this difference is as follows: In the case of the materials used in the prior art, refractive or gradient structures (GRIN) are created to create refractive or diffractive structures, according to the invention In the case of hydrogels according to the same, it is possible to achieve similar refractive or diffractive effects by forming patterns of various voxel depths while keeping the modified refractive index nearly constant. Furthermore, it is believed that this refractive index value will first approach, and eventually approximate, the refractive index of isotonic saline.

  As a result of this new mechanism, the newly created refractive or diffractive structure forms a system of parallel waveguides of different lengths (ie elongated voxels with a lower refractive index than the surrounding hydrogel) , Whereby the phase shifts are controlled by the varying voxel depth rather than their varying refractive index.

  A further feature of the invention is to create a gradient of refractive index in the vicinity of the individual voxels. The monomers and other fragments released by depolymerization in voxels move radially by diffusing through the surrounding hydrogel. When the temperature drops below the ceiling temperature (around 200 ° C. in the case of methacrylates), at least a portion of the released monomers and / or fragments re-polymerize and have a higher density with a higher refractive index than the parent hydrogel It can create a network structure. This mechanism has two beneficial consequences: First, it reduces the amount of compound diffused and metabolized outside the implant, and secondly, by the gradient of the refractive index thus formed The light guiding properties of the voxels are improved.

  That is, the monomers released by depolymerization can not all diffuse to the outside of the hydrogel, but can partially repolymerize near voxels. In that sense, the method is somewhat similar to FS laser ablation in that it both removes some polymer clumps and instead holds water, but the term "ablation" inherently means that the polymer and water are It often refers to the process of decomposition that is converted to plasma (gas). The processes disclosed herein are even more gentle, so they do not create bubbles that would break the polymer degradation.

  Depolymerization of the hydrogel copolymer is believed to be the primary mechanism involved in the process disclosed herein, but this involves small water soluble fragments such as hydrolysis or oxidation of (meth) acrylate pendant groups It can be compensated by other decomposition reactions that form.

  The dopant concentration in the hydrogel according to the invention varies from about 0.05 mol% to about 5 mol%, preferably from about 0.1 mol% to about 2.5 mol%. Alternatively, different optimum dopant concentrations may be used for the various dopants, for example 0.25 to 0.55 mole% for benzophenone derivatives or 0.1 to 0.2 mole% for benzotriazole derivatives. The preferred dopant is a low absorbance for visible light, ie a UV absorber greater than a wavelength of about 390 nm. Examples of suitable dopants are vinyl, acrylate or methacrylate derivatives of benzophenone, benzotriazole and coumarin, but one skilled in the art will ensure that other suitable UV absorbers act as dopants within the meaning of the invention. It should be identifiable.

  The activating group is present at a concentration of about 0.25 mol% to about 5 mol%, preferably about 0.75 to about 3.5 mol%. In some embodiments, the preferred molar concentration of methacrylic acid is 1-1.25 mole%. Preferred activator groups are derivatives of acrylic acid or methacrylic acid comprising a pendant acidic group comprising a carboxylate group, a sulfate group, a sulfonate group and a phosphate group. Such acidic groups are preferably neutralized by suitable organic or inorganic cations. The molar ratio of activator to dopant groups should be about 0.75 to 10, preferably about 1 to about 5. Alternatively, different optimum molar ratios of activator / dopant will be different for different dopants. For example, the optimum ratio for benzophenone derivatives is about 2 to 4 and the optimum ratio for benzotriazole derivatives is about 5 to about 7.

  The preferred composition of the hydrogel according to the invention comprises a major fraction of methacrylate monomers with pendant neutral hydrophilic groups. By "major fraction" is meant at least 50 mol% of the total monomer units in the hydrogel. In some embodiments, the mole fraction of the hydrophilic methacrylate monomer is 90 mol% to 99.5 mol%, often 97.5 mol% to 99 mol%. Such hydrophilic methacrylate monomers may be esters of polyol aliphatic compounds such as glycols, glycol ethers, glycerol and sugars. The best known of this group is 2-hydroxyethyl methacrylate (2-HEMA). Alternative monomers to the above mentioned esters include the amides of methacrylic acid such as methacrylamide, N-isopropylmethacrylamide or N- (2-hydroxyethyl) methacrylamide. This major fraction of hydrogel polymer may also comprise a mixture of the above hydrophilic monomers. Alternatively, minor portions (but up to 25 mol%) of methacrylate monomers can be replaced by similar acrylic acid derivatives. In some embodiments, 0.5 mol% to 5 mol% can be replaced with an acrylate monomer.

  The minor portion of the monomer unit is formed by the above-described activator monomer with a negatively charged pendant group, and the further minor portion is formed by the above-described dopant monomer.

  The polymers are preferably covalently crosslinked. Crosslinking may be achieved by any of the methods known to those skilled in the art, such as radiation crosslinking, crosslinking by formation of ether linkages between pendant OH groups. A preferred method of crosslinking is, for example, copolymerization of a polyol with a minor fraction of methacrylate or acrylate cross-linked diesters or triesters, such as triethylene glycol dimethacrylate.

  The refractive index of the hydrogel according to the invention is about 1.38 to about 1.48, preferably about 1.40 to about 1.45. In some embodiments, RI is about 1.40, or about 1.41, or about 1.42, or about 1.43, or about 1.44, or about 1.45.

  Preferred hydrogels according to the present invention comprise about 25 wt% to about 85 wt% liquid in equilibrium with the native intraocular environment. In the case of an intraocular lens that supplements or replaces the native lens, a more advantageous equilibrium fluid concentration is 35% to 50% by weight, and more specifically 40% to 47% by weight. In some embodiments, the equilibrium water content is 41% ± 0.75%, or 42.5 ± 1 wt%, or 44.5 ± 1 wt% by weight. The person skilled in the art has a common understanding that the equilibrium liquid content in the hydrogel is influenced by many variables, such as body temperature, the composition of the body fluid or the pressure exerted by the surrounding tissue or body structure. Also, these values are illustrative, as measurement of liquid content in small hydrogel implants can be subject to certain measurement errors. Since negatively charged activator groups tend to increase the equilibrium liquid content, ie, to decrease the refractive index of the hydrogel, higher concentrations of activators reduce the water content and the refractive index to the extent desired for ophthalmic implant design It may be corrected by the addition of hydrophobic acrylates or methacrylates, such as methyl methacrylate, ethyl methacrylate, benzyl methacrylate, isobornyl methacrylate or ethoxyethyl methacrylate which will enhance the Typical concentrations of the above-mentioned added methacrylates or acrylates are up to about 40 mol%. In some embodiments, the concentration of hydrophobic monomer added is about 5 mol% to 25 mol%.

  The dopant group and the activating group may be located on the same polymer chain, for example on separate polymer chains in close contact, eg in the polymer mixture or on two different segments of the polymer network In some cases. One may be on the graft, or one on the graft and the other on the base strand. Dopant groups and activating groups can be present in a single molecule in the proper conformation to allow their interaction. Also, such "dopant-activator complexes" may be covalently attached to the polymer chain, mixed into the polymer, or be part of the main polymer chain.

  The negatively charged activator group has improved biocompatibility, resistance to protein adsorption, resistance to biofilm formation, resistance to calcification, and resistance to cell adhesion and diffusion (back capsule hardening Have some additional advantages over intraocular implants, such as resistance to and resistance to posterior capsule opacification.

  The hydrogel implants according to the invention can be placed at various positions in the eye along the light path. It may have the form of a "blank" in which a refractive or diffractive lens is created, or it may be of a refractive or diffractive lens whose optical properties have been modified by refractive index changes in selected locations of the implant It may have a form.

  The intraocular lens can be replaced partially or completely with the native lens. Intraocular hydrogel lenses have several advantages over other IOL types, and various types and associated manufacturing methods are described, for example, in the following patents and patent applications: Stoy, V. et al. Et al: Bioanalogic Intraocular Lens, International Patent Application No. WO 2014111769; Wichterle, O. et al. Method Of Molding An Intraocular Lens, US Pat. No. 4,846,832; Wichterle, O. et al. Stoy, V .: Soft and Elastic Intracameral Lens And Method For Manufacturing Thereof, U.S. Pat. No. 4,846,832; U.S. Pat. No. 5,674,283; Sulc, J .: Implantable Ophthalmic Lens, A Method Of Manufacturing Same And A Mold For Carrying Out Said Method. Et al: Soft Intracameral Lens, U.S. Patent Nos. 4,994,083 and 4,955,903; and Michalek, J. et al. Et al .: Method Of Manufacturing An Impantable Intraocular Planar / Convex, Biconvex, Planar / Concave Or Convex / Concave Lens And Using A Lens Made Using This Method, US Patent No. 8409481. Each of these is incorporated herein by reference.

  Another type of implantable ophthalmic lens is the so-called "embedded contact lens" (or ICL). The ICL is a non-quama lens which is placed between the iris and the natural lens. It is described in several patents, for example, Fedorov et al., Intraocular lenses for correcting moderate to severe hypermetropia, US Pat. No. 5,766,245; Intraocular contact lenses and methods of implantation, UP 5, 913, 898; Intraocular reflective correction lenses, US Pat. No. 6, 106, 553, each of which is incorporated herein by reference.

  These implantable lenses are made from hydrogels that incorporate biological components, usually collagen. These so-called "colamers" are described by Feingold et al., Biocompatible, optically transparent, ultraviolet light absorbing, polymeric material based upon collagen and method of making, US Patent No. 5, 910, 537; Feingold et al., Biocompatible optically transparent polymeric material upon reaction No. 5,654,349, U.S. Pat. No. 5,654,388 and U.S. Pat. No. 5,661,218; Fedorov et al., Biocompat Various patents have been described, such as ble polymeric materials, methods of preparing such materials and uses thereof, U.S. Pat. No. 5,993,796; and methods of preparing a biological material for use in ophthalmology, each of which is incorporated herein by reference. Incorporated into

  Another approach to refractive correction is an intra- or intra-corneal implant or inlay, such as Miller, Aspherical corneal implant, US Pat. No. 7,776,086; Lang, Alan, Design of Inlays With Intrinsic Power, US Pat. Application Publication No. 2007/0255041; Dishler; Jon et al., Small Diameter Inlays, US Patent Application Publication No. 2007/0203577; Lang, Alan et al., Intracorneal Inlays, US Patent Application Publication No. 2007/0129797; and Dishler et al., Method of using small diameter intracorn al inlays to treat visual impairment, are described in U.S. Pat. No. 8,057,541, incorporated herein by reference.

  Some intra-substrate implants are designed for post-operative refractive power adjustment with a laser, for example Peyman, Gholam A. et al. , Intrastromal corneal modification via laser, U.S. Patent Application Publication No. 2001/0027314; Adjustable ablatable inlay, U.S. Patent Application Publication Nos. 2002/0138069 and 2002/0138070; Ablatable intracorneal inlay with predetermined refractory properties, U.S. Patent Application Publication No. 2003/0093066; and Bifocal implant and method for altering the reflective properties of the eye, described in US Patent Application Publication No. 2005/0222679, Each is incorporated herein by reference.

  Also, an intraocular implant using two lenses side by side to achieve improved accommodation, modularity of design or reduction in size of the embedded incision when assembling the implant in situ from the individual parts There is also.

  All of the above mentioned ophthalmic lens types can be made from hydrogels according to the present disclosure and then corrected by using femtosecond laser processing.

  Again, as not wishing to be limited by any particular theory, the partial depolymerization of the method disclosed herein is not limited to the present invention when exposed to a focused femtosecond laser (FSL). It is believed to be the mechanism of the large negative phase shift observed in hydrogels. As indicated above, this process appears to produce a system of parallel longitudinal "voxels" which are further stretched as the number of FSP pulses absorbed is increased. The voxel length can not be reduced below the RI value of water, which is 1.3335, to control the achievable phase shift for a given refractive index. Furthermore, it is not the refractive index itself but the phase shift that is actually affected in the hydrogels and methods disclosed herein. This mechanism is different from that disclosed by the prior art (the estimated crosslinks result in a decrease in water content leading to an increase in refractive index, resulting in a corresponding positive phase shift).

  A further difference from the prior art is that the hydrogels and methods disclosed herein do not reduce the water content of the hydrogel (such as by polymer crosslinking) by modifying the copolymer properties, but rather the hydrogel formation. Replacing part of the copolymer with water or an aqueous fluid.

  In the drawings, like numerals indicate like elements throughout. Certain terms are used herein for convenience only and are not to be regarded as a limitation on the present invention. Hereinafter, preferred embodiments of the present invention will be described. However, based on this disclosure, it should be understood that the present invention is not limited by the preferred embodiments described herein.

NCL has a very complex structure that evolves over time. One of the structural features is the asphericity of the posterior and anterior surfaces of NCL 103. Recently, E.I. L. Markwell et al., MRI study of crystalline shape with accommodation and accommodation in humans, Journal of Vision (20110 11 (3); 19, 1-16); M. Dubbelman et al., Change in shape of human crystalline lenses with accommodation, Vision Research 45 (2005), 117-132; F. Manns et al., Radius of curvature and affinity of the anterior and posterior surfaces of human cadaver crystall in eye lens, Experimental Eye Research 78 (2004), 39-51; M. Dubbelman et al., The shape of the aging human lens: Curvature, equivalent refractive index and the lens paradox, Vision Research 41 (2001) 1867-1877. As such, the anterior and posterior surfaces of the juvenile lens are both hyperbolic and may be characterized by the following equation:
Y-Yo = X ^ 2 / {Ro * (1 + 1-h * (X / Ro) ^ 2) ^ 0.5} Equation 1

  Where Y is the coordinate in the direction of the main optical axis 1A, X is the distance from the main optical axis 1A, Yo is the position of the vertex on the main optical axis 1A, Ro is the central radius of curvature, h Is the conic constant (or shape parameter). Equation 1 describes the curve of any conical cross-section that varies with the shape parameter h value, and is parabolic if h = 0, circular if h = 1, and h <0 Is a hyperboloid shape, is an oblong oval shape in the case of 0 <h <1, and is an oblong oval shape in the case of h> 1.

  For a typical young NCL, the anterior surface is more hyperbolic than the posterior surface, hyperbolicity is significantly increased by accommodation, and the human lens grows with age so that the old NCL can be approximately spherical. It has been found that the hyperbolicity is reduced.

  In the referenced study, the dimensions of typical NCLs for selected population samples were mapped. According to these references, the anterior central radius of a typical human lens is in the range of about 5 to 13 mm, and the average anterior cone parameter is about -4 (ranging from about -22 to +6). The aft center radius is in the range of about 4 to 8 mm and the mean aft cone parameter is about -3 (ranging from about -14 to +3).

  The central thickness of relaxed NCL in young people is typically in the range of about 3.2 mm to about 4.2 mm, with a thickness of about 3.5 mm to about 5.4 mm with aging and / or short focus adjustment To increase. The posterior depth of the NCL is typically the same as or greater than the anterior depth. Thus, the sagittal depth of the posterior lens surface is typically about 1.75 mm to about 2.75 mm on an equatorial diameter of about 8.4 mm to about 10 mm. This defines the basic dimensions of the posterior capsule in the "native" state.

  Although the above references do not mention any specific relationship between the geometrical and optical properties of NCL, we note that the hyperboloid has a lens whose power is at its maximum at its central part. It turned out by mathematical modeling that it changed to multifocality which decreases gradually towards the periphery. One direct consequence predicted from such multifocality is the large depth of focus of the lens so that nearby objects can be projected onto the retina without any specific lens shape change . Another result inferred from the modeling is that the average power of the lens increases with decreasing aperture. Thus, to the conclusion that the short focus can be improved by the contraction of the pupil, which is so-called "pupil reflex" or "near myosis", which can be observed practically at the short focus Reach. Another result, inferred from the hyperbolic nature of the innate lens, is a natural neurofit of the human (especially young) brain that correctly interprets the image formed by projecting onto the retina through a hyperbolic lens It is an ability.

  This accommodation mechanism, which exploits a certain type of multifocality, is worthy of further explanation as follows.

  A lens having at least one hyperboloid exhibits "hyperboloid aberration" which is the opposite of the spherical aberration: the ray incident through the center is bent to the focal point closest to the lens, which focal point is the lens center Gradually from the lens due to light rays incident at increasing distances from the lens towards the lens periphery.

  Thus, a lens with a hyperboloid is positive multifocal: it has the shortest focal length at the center (ie the highest refractive power) and the focal length increases from the center towards the lens periphery (Ie, the refractive power decreases). The focal range of the hyperboloid lens can be rather large and can be controlled by so-called conic constants or shape parameters defining a hyperboloid shape.

An example of the distribution of refractive power in a lens with a hyperboloid is shown in FIG. 2 in which the local refractive power indicated by dioptre m ⁻¹ is plotted against the distance from the optical axis indicated in mm. The above optical aspects of decreasing refractive power with distance from the optical axis, or with the aperture of the imaging system, or with the pupil diameter of the eye, may already be present in the implanted lens, or It can be seen that this can be produced by laser hydrogel correction after lens implantation.

  Based on tests in the present invention, it can be seen that positive multifocality and its changes in the innate lens help the eye to adjust in several ways:

  The native lens causes the images of all objects in the field of view at all distances covered by the focal region of the lens to be simultaneously projected onto the retina. This is enough to create an image in which all objects are well in focus (with many out of focus images that the brain learns to suppress), which significantly increases the depth of focus of the eye (1 diopter Increase).

  The native lens increases its hyperbolicity by accommodation, which further increases the focal area of the lens and thus furthers the depth of focus.

  The eye helps to focus on near objects by narrowing the pupil. This so-called "pupil reflex" or "short-distance migraine" has two consequences: first it increases the depth of focus of the eye as an optical system to reduce the aperture (aperture Narrowing blocks the light rays which are far from the axis and incident at an acute angle with respect to the axis) and increases the average power of the lens by using only its central part with the highest refractive power.

  It is clear from our tests that short-distance pupils can only aid short-focus for lenses with hyperbolic aberration, ie positive multifocality. Monofocal parabolic lenses have little effect, negative multifocal lenses have opposite effect: spherical or elliptical (eg, meniscoid) lenses are short lenses rather than stronger lenses required for short focus It becomes a weak lens with low refractive power due to the distance pupil.

    The artificial lens according to the present invention is a hydrogel device that can be implanted in the posterior chamber of the human eye to replace the native lens. The artificial lenses according to the present invention are designed to mimic or replicate the essential physiological and optical functions of the innate lens without causing the problems that can be brought about in some circumstances by previous attempts. It is important to recognize that this is achieved with a novel combination of features that may not be very successful and may have been applied individually or in different combinations before. The innate lens also achieves its function by a combination of balanced features, rather than by a single feature.

  Features contributing to the overall function and combined according to the invention include the size and shape of the implant, the material properties, the surface properties, the optical properties, the embedding method and the manufacturing method. We describe the various features below and provide a specific configuration of how the individual features interact to provide a beneficial effect. The implant may combine some of the described features to achieve the desired effect, but recognizes that the present invention is not limited to the exemplary configurations described below but includes various combinations of features Is important.

  Referring to FIGS. 3A and 3B, the implant has a main optical axis 1A with a central optical portion 2 and a peripheral support portion 3. The overall shape of the implant is defined by its anterior surface 4, posterior surface 5 and transition surface 6 between upper boundary 7A and upper boundary 7B of each of the anterior and posterior surfaces. Each face is composed of two or more faces. The front central optical surface 8A has a boundary 9A, and the central rear optical surface 8B has a boundary 9B. Each face may be divided into two or more sections by boundaries (shown as 13A and 13B in FIGS. 5A-5C) that are circular, linear or otherwise defined shapes. The top 10A of the center anterior optical surface and the top 10B of the center posterior optical surface are disposed on the main optical axis 1A. The front peripheral support surface is 11A and the rear peripheral support surface is 11B.

  With reference to FIGS. 6A, 6B and 6C depicting lenses comprising several different materials or layers, it goes without saying that any layer or structure in the light path can be formed by the hydrogel according to the invention. A preferred layer may be the layer closest to the cornea, ie the layer forming the anterior optical surface of the lens. The optical structure formed by the hydrogel optical modification may be a refractive structure, a diffractive structure, a Fresnel lens as shown in FIG. 6B, a refractive index gradient lens or the like. Such structures may be formed in the hydrogel or on the surface thereof, preferably on the anterior optical surface.

  The boundaries 7A and 7B can be distinguished as the top cut of the front surface 4 and the rear surface 5, respectively. Such a break is at the inflection point of the surface in the optical axis direction or at the discontinuity of the second derivative of the surface in the optical axis direction. The boundaries may be curvilinear or continuous, but preferably the boundaries are formed by sharp edges or ends. The advantage of the sharp end is to create an obstacle to the migration of cells such as fibroblasts along the surface of the capsule (the usual cause for posterior capsule opacification).

  The overall diameter is defined as the larger of the boundaries 7A and 7B. The diameter of the lens optical zone is defined as the smallest diameter of the boundary 9A and the boundary 9B. The posterior sagittal depth is the vertical distance between the posterior apex 10B and the plane defining the posterior boundary 7B. The center thickness is the distance between the vertex 10A and the vertex 10B. The front depth is the vertical distance between the front vertex 10A and the plane defining the front boundary 7A.

  If the main optical axis 1A is defined by the circles in the plane perpendicular to the optical axis, the boundaries 7A and 7B and the boundaries 9A and 9B are defined by a circle, and the central optical part 2 is symmetrical, for example a cylindrical component If not, it may be a symmetry axis. Such an implant having a symmetrical circular footprint is shown in FIG. 3B. However, the edges and boundaries may have other than circular footprints, for example the oval footprint shown in FIG. 4A, or have single, two, three or four tip cuts 12A-12D. It may have a footprint shaped as the tipping circle in FIGS. 4B-4E. These tipping footprints serve several purposes.

  The footprint provides better access to the space behind the lens during implantation. Before closing the surgical incision it is important to clean this space in order to remove the visco-elastic polymer or lubricant or other auxiliaries.

  The footprint prevents lens rotation after the lens capsule has contracted around the IOL. This is particularly important in the case of toric lenses.

  The footprint facilitates folding and insertion with small incisions.

  If the optical system has cylindrical components, the cylindrical axis 1B is to be arranged in a defined manner as to the asymmetry of the outer edge, for example at an angle α to the tip cuts 12A and 12B shown in FIG. 4F. Become. It goes without saying that the distal cuts 12A to 12D do not necessarily have to be straight cuts, but preferably they may be shaped, for example, in the shape of a crescent, and their number may exceed four. Also, the tip cuts do not have to be the same length and may not be arranged symmetrically. It can be seen that a footprint with a truncated edge facilitates folding of the implant and facilitates its insertion with a small surgical incision. Furthermore, the non-symmetrical edge footprint will prevent rotation of the implant once the capsular bag has settled around it. This is particularly important in the case of toric lenses with cylindrical components designed to correct astigmatism.

  The posterior surface 5 is shaped and sized to approximate the shape and size of the posterior surface of the native lens and to achieve contact with at least a major portion of the posterior capsule of the eye. This is important for several reasons.

  The implant, due to its optimal optical performance, will keep the posterior capsule free of wrinkles and smooth in its native shape.

  Intimate contact between the lens capsule and the implant will prevent migration of fibroblasts that can induce posterior capsule opacification, which is highly hydrated on the posterior surface and has a negative fixed charge. It is especially effective when it has.

  The implant prevents the vitreous from advancing by occupying the space left by the posterior side of the native lens, thereby reducing vitreous pressure on the retina (promoting retinal detachment and / or cystic macular edema ) Can be suppressed.

  Close contact between the implant and the posterior capsule will result in fibrosis of the posterior capsule and consequent hardening, clouding that will interfere with (or even displace) implant function, as described later herein. It should be noted that the contact surface of the implant is hydrophilic and is particularly beneficial if it has a negative fixed charge to prevent contraction and contraction.

  In a preferred embodiment of the invention, at least the main part of the rear surface 5 is a generally smooth convex surface formed by the rotation of the conical cross section around the optical axis or a combination of such surfaces. The periphery is preferably formed by a conical surface or a hyperboloid, and the central optical surface is preferably formed by a hyperboloid, a paraboloid or a sphere (or a combination thereof). The sagittal depth of the back surface (i.e. the vertical distance between the back center optical surface vertex 10B and the boundary of the back surface 7B measured on the principal optical axis 1A) is 1.1 mm so that the lens can perform its function well Should be larger. The aft sagittal depth should be greater than 1.25 mm, advantageously greater than 1.75 mm, and preferably greater than 2 mm, in order to function well throughout the refractive range, but in any event about Should be less than 2.75 mm.

  The overall outer diameter (LOD) of the implant is important for its centrality, positional stability and lens capsule filling ability. The outer diameter of the rear face 5, ie the largest dimension of the rear outer boundary 7B (in a plane perpendicular to the main axis 1A), should be greater than 8.4 mm, preferably at least 8.9 m and preferably at least 9.2 mm is there. The maximum allowable outer diameter is about 11 mm, but should desirably be less than 10.75 mm, preferably less than 10.5 mm. The considerable flexibility in the external dimensions is due to several factors, namely the flexibility of the lens and, in particular, of the outer peripheral support part 3 which can adjust different capsule sizes and capsule contractions without deforming the central optical part 2 It is made possible by the flexibility.

  The central optical surface may be constituted by one or more geometrically different areas. The areas may be concentric, in which case the back boundary 13B between them in FIG. 5A is circular. The areas may also be divided by straight boundaries, in which case the areas may have a crescent-shaped or wedge-shaped footprint. Various examples are shown in FIGS. 5A-5C. The area may be on the anterior optical surface or may be on the posterior optical surface. FIG. 5A shows that the posterior optical surface is divided by the boundary 13B into two concentric optical areas, a central optical area 8B1 and an outer optical area 8B2. For example, the posterior optical surface of central optic area 8B1 may be a spherical or parabolic area used for clear near vision, and the hyperbolic outer area plays a role for intermediate and distance vision. Alternatively, both areas may have hyperboloids with different central radii Ro and / or different conic constants. Each optical surface may also be divided into more than two areas. The example in FIG. 5B shows a top view of a lens in which the anterior optical surface 8A is divided into two optical areas, an area 8A1 and an area 8A2 that are equal to a linear boundary 13A. Each of these areas has different shapes with different optical parameters. The example in FIG. 5C is a top view of a lens having a front optical surface 8A divided into four pairs of optical zones 8A1 and 8A2 by two straight boundaries 13A and 13B, each having different areas and different optical parameters. Show. For example, 8A1 has a higher power than 8A2 and may play a role for short focus. One of the zones may have a cylindrical component.

  Both optical surfaces (or their areas or segments) are surfaces formed by rotation of the conical cross section along the optical axis, or a combination thereof. One or both of the optical surfaces may include one or more spherical optical areas. Advantageously, at least one of the optical surfaces preferably comprises at least one hyperboloid in the outer optical area. Preferably, both optical surfaces each comprise at least one hyperboloid area. Such hyperboloids are similar to the plane of NCL and mimic some of their useful optical properties. Even more preferably, the posterior and anterior optical surfaces are both hyperboloids or combinations of two or more concentric hyperboloid areas. A lens with at least one hyperboloid has so-called hyperbolic aberration, which is the exact opposite of the spherical aberration of a lens with spherical, elliptical or meniscoid surface. A lens with hyperbolic aberration has the highest refraction at the center, which decreases gradually with distance from the optical axis. (In lenses with spherical aberration, refractive power increases with distance from the optical axis.) Hyperbolic aberration helps the eye to adjust through some of the above mechanisms. It is clear that hyperbolic aberrations can be produced not only by geometrically accurate hyperboloids, but also by similar surfaces whose surface steepness generally decreases with distance from the optical axis. is there. Thus, "hyperboloid" also refers to other hyperbole-like surfaces that approximate this property.

  Because spherical aberration is measured at the aperture (or pupil diameter), the negative spherical aberration of the hyperboloid-like surface increases in absolute value (i.e., more negative values increase). Spherical aberration corresponds to the deflection of the wavefront as shown in microns, or as a steeping of the decrease of the local power from the optical axis, by the decrease of the power with the increase of the aperture, or to such an optical aspect It can be expressed in various alternative ways, such as by the value of the conic constant or shape parameter. One skilled in the art could easily convert one of such values to the corresponding value represented in a different manner. The implant according to the invention may inherently have any spherical aberration, since the value of the spherical aberration can be adjusted after surgery using the method of the invention.

  The spherical aberration in the final embedding step (i.e. after laser hydrogel correction) will generally be about -0.1 to -2 microns at an aperture of 4.5 mm. Preferably, the final spherical aberration will be about -0.5 microns to -1.5 microns for an aperture of 4.5 mm, and even more preferably the spherical aberration will be about -0.75 for an aperture of 4.5 mm It is micron to -1.25 micron.

  In order to mimic the optical properties of NCL, the conical constants of the anterior and posterior optical surfaces, from the highest value at the central optical part 2 to the lowest value at the peripheral part of the central optical part 2, at the optical axis Choose to decrease generally.

  The steepness of the decrease of the refractive power with the distance from the optical axis changes with the shape parameter (cone constant) of the hyperboloid. The conical parameter has an average decrease in optical power of -0.25 Dpt / mm to -3 Dpt / mm, advantageously -0.5 Dpt / mm to -2.5 Dpt / mm, preferably about -1 Dpt / mm to -2 Dpt / mm. Should be chosen to be mm.

  The back central radius of curvature (at the point where the optical axis intersects the back apex) is advantageously 2.5 to 8 mm, preferably about 3.0 to 5 mm. The conic constant of the rear face is advantageously selected from the range of about +3 to about -14, preferably about -1 to -8 reported for NCL.

  The central radius Ro of the anterior optical surface 8A is selected to be greater than about +3 mm or less than about -3 mm, preferably greater than about +5 mm or less than about -5 mm.

  The conical constant of the anterior optical surface 8A is selected from the range of +6 to -22 reported from human NCL, preferably from the range of about -1 to -8 mm.

  The anterior optical surface 8A may be partially or completely formed by a spherical or paraboloid. In that case, the central posterior optical surface 8B should preferably be a hyperboloid with conical parameters selected in such a range that the whole lens has hyperbolic aberration.

  However, preferably at least the main part of the front optical surface 8A is a hyperboloid, in particular the outer optical area. The central optical zone of the front optical surface having a diameter of about 1.5 to 4 mm, preferably about 2 to 3.5 mm, may be formed by a paraboloid or a spherical surface to further improve the short focus resolution.

  FIG. 2 schematically shows an example of a preferred optical aspect of a lens according to the invention. It should be understood that different eyes have different required refractive power of the implanted lens.

  Most of the current IOLs are designed to mimic only the basic optical functions of the NCL, ie to provide the basic power needed to focus on distant objects on the retina Because it is not a biosimilar type. Depending on the specific eye, the basic refractive power is usually 15 to 30 Dpt, with some variation on either side. An approximately monofocal (usually spherical) rigid lens located somewhere near the main plane of the NCL can meet this requirement. Most IOLs are NCL, as the most detailed image is projected onto a relatively small part of the retina (macular) located at the optical axis, and most of our activity takes place in the small eye opening (miosis) It is significantly smaller (4.5-6 mm for most IOLs versus 9.5-10.5 mm for NCL). Some small size optical systems are preferred by some IOL manufacturers, as such IOLs are more easily adapted to implantation by small incisions. For the same reason, most IOLs are manufactured from a soft, elastic material that allows for small incision implantation in a deformed (folded, rolled, etc.) shape. However, this deformability has nothing to do with the optical function.

  However, small sized optical systems have their disadvantages. The IOL end may reflect light rays with a large pupil opening (e.g. during driving at night), which may lead to glare, halos and other harmful effects. Moreover, small optics, especially with large pupils, can not image all marginal and off-axis rays as NCL does. Finally, the small size of the optics interferes with the clear visibility of the retinal periphery that may be needed for diagnosis and treatment. For those reasons, large optics similar in size to NCL are preferred over the small optics used in most current IOLs. Importantly, the entire large optical area must have a clear geometry to be optically useful. Lenses with meniscoid optical surfaces have shapes that are poorly defined, especially in the peripheral region. This can cause unexpected disturbing optical phenomena.

  Some current IOLs are designed to mimic NCL accommodation or pseudoaccommodation to some extent (ie, the eye can be focused on both distant and nearby objects). is there. Different IOLs use different means to achieve this goal: Some use bifocal, multifocal or multifocal optics, and allow for the forward-backward shift of the IOL optics with respect to the eye Some also use designs that allow changes in optical power by changing the position of the two lenses relative to each other. Some lenses may even have a change in the pressure of the ciliary muscle and / or vitreous, a change in the position of the head, or a small pump acting to move the liquid within the lens to change the refractive power.

  These designs are sometimes rather complex and novel tricks and are very different in size, shape and material properties from NCL. This results in various problems such as lens capsule fibrosis or cell ingrowth or protein deposition on their surface that interferes with their function. In addition, their bulky and complicated design precludes the requirement for all current IOLs to be implantable with a small incision. This increases glare and halo problems as it requires design with small diameter optics and the use of high refractive index materials that are more refractive than NCL.

  In most cases, these lenses use small diameter optics, typically 4.5-6 mm, with elongated and flexible "haptics" to place the optics in the center of the light path There is. In addition, deformable materials are used that allow folding or winding for implantation with small incisions. The surface properties of such IOLs may be modified to achieve better biocompatibility (e.g. A. M Domschke, U.S. Patent Application Publication 2012/0147323, J. Salamone et al., U.S. Patent Application Publication 2008/0003259).

  This common design allows the folding of the IOL for implantation with a relatively small incision (usually 2-3 mm). However, this small IOL size has its own drawbacks.

  Insufficient light conditions cause large eye openings (causes night glare, halos, peripheral vision limitations etc) or eccentric IOL ("sunset syndrome" or other problems) A small optical system, 6 mm in diameter or less, may not completely replace the 9-10.5 mm diameter lens.

  Smaller optics may not be able to reflect all the peripheral and off-axis rays that the NCL does, so imaging performance may be reduced, especially with large open pupils (e.g. needed for night vision peripheral vision) .

  Small optics may complicate or even prevent examination and treatment of the retina (which may be important, especially in the case of diabetes).

  In addition, the small IOL size essentially leaves the space originally occupied by the much larger NCL. As a result, the vitreous is advanced and its pressure on the retina is partially relieved. This is described in J. A. Rowe, J.J. C. Erie, K .; H. Baratz et al. (1999). “Retinal detachment in Olmsted County, Minnesota, 1976 through 1995”. As reported by Ophthalmology 106 (1): 154-159, it may be responsible for the increased likelihood of retinal detachment after cataract surgery. The same effect may also induce or promote cystic macular edema (CME). Steven R. See Virata, The Retina Center, Lafayette, Indiana: Cistoid Macular Edema, WEB page.

  Another disadvantage exists in conventional IOL designs with small optics and haptics: IOLs with optics suspended in relatively open spaces by relatively fragile haptic means are incidental It may be susceptible to damage and / or displacement in the event of an impact (such as a slip on a slippery surface, a car crash, a punch hit, etc.).

  Some problems stemming from the small bulk IOL and small diameter optics are addressed by the IOL design which fills the space left and right by the NCL. There are several approaches to this, each having its own advantages and disadvantages.

  Fill the capsular bag with a liquid that can solidify into a clear, flexible solid such as silicone rubber. It was expected that this approach would restore the regulatory function of the native lens as long as the filler material had a deformability similar to that of NCL (eg, Gasser et al., US Pat. No. 5,224,957). However, materials used to date often induce fibrosis and clouding of the lens capsule. Moreover, it is difficult to control the shape and optical parameters of the in situ formed IOL.

  Implants of large, bulky IOLs with highly deformed shapes that allow for implantation by a reasonably small incision and fill a significant portion of the lens capsule. This approach has been tested with hydrophobic memory polymers that can be "frozen" in a highly deformed shape for implantation and return to their original functional shape when heated to body temperature (Gupta, US Pat. No. 4, 834,750 and U.S. Patent No. RE 36,150). However, this hydrophobic memory polymer is a very heterogeneous material and induces problems similar to similar materials used to fill the lens capsule.

  We also tried a similar approach with hydrogels. Very large IOLs that mimic the size and shape of the innate lens were embedded in the open lens capsule (eg, Wichterle 732 and Stoy 283). The problem with these particular IOLs was their special optics. These lenses had a meniscoid anterior optical surface strongly derived from the geometry of the NCL. This meniscoid shape was formed by the solidification of the free surface of the monomer mixture and had problems with the control of the optical properties of such IOLs. Furthermore, these lenses were often too bulky to implant with a small incision. Furthermore, some of the hydrogels used in these lenses lack a negative fixed charge, but such hydrogels have a tendency to calcify at some point after their implantation. Some other lens capsule-filled lenses (Sulc et al. '083 and' 903) stabilize the lens at approximately the center position by touching the iris but block the liquid flow, deform the lens optics, and irritate the lens Some have a forward protrusion that causes various problems such as

  Another approach has been the implantation of hollow lenses (or lens shells) that are filled after implantation by a liquid that solidifies in situ (eg, Nakada et al., US Pat. Nos. 5,091,121 and 5,035, 710).

  Another approach is the implantation of a dual optics IOL, with two lenses, one in contact with the anterior capsule and the other in contact with the posterior capsule, both lenses being separated by a flexible member or connector (US Patent No. 4,946,469; US Patent No. 4,963,148; US Patent No. 5,275,623; US Patent No. 6,423,094; US Patent No. 6,488, U.S. Pat. No. 6,761,737; U.S. Pat. No. 6,764,511; U.S. Pat. No. 6,767,363; U.S. Pat. No. 6,786,934; U.S. Pat. U.S. Pat. No. 6,846,326; U.S. Pat. No. 6,858,040; U.S. Pat. No. 6,884,261).

  Such implants, which essentially fill the entire capsular bag of the original lens, also have several problems.

  The front face of the IOL should touch the iris and induce erosion, depigmentation, hemorrhage or inflammation if not made from a highly biocompatible material with hydration and negative charge similar to NCL There is also.

  Some materials are more biocompatible by having high equilibrium water content. However, in that case, the refractive index will be reduced to much lower than the optimum value (the value for young NCL).

  Important to the IOL are not only the shape and type of optics, but also its materials. NCL is composed of a complex natural hydrogel structure comprising water, salt and polymer components containing collagen proteins, polysaccharides and proteoglycans. Importantly, the polymer component contains acidic ionic groups such as carboxylates or sulfates at significant concentrations. These groups provide the lens material with a negative fixed charge. Hydration and negative charge affect the interaction between NCL and proteins in intraocular fluid. Furthermore, its surface properties influence the interaction between the lens and the cells. Synthetic hydrogels containing surfaces with a negative fixed charge do not attract proteins and cells, making them more resistant to calcification (Karel Smetana Jr. et al., "Intraocular biocompatibility of Hydroxyethyl methacrylate" and Methacrylic Acid Copolymer / Partially Hydrolyzed Poly (2-Hydroxyethyl methacrylate), "Journal of Biomedical Materials Research (1987) vol. , "The Influence of Hydrogel Functional Groups on Cell Behavior", Journal of Biomedical Materials Research (1990) vol.24 pp.463-470) it is known. Many IOL manufacturers avoid materials with carboxylate groups based on the assumption that carboxylates induce calcification by attracting calcium ions, but some references to hydrogel IOLs that contain carboxylate groups (Wichterle '732, Sulc et al.,' 083 and '903, Stoy, U.S. Patent No. 5,939,208, Michalek and Vacik,' 093).

  For example, as described in Stoy '208 and Sulc et al., US Pat. No. 5,158, 832, carboxylate groups are uniformly dispersed in the hydrogel, or primarily due to swelling and charge density gradients. It can be concentrated on the forming surface. Typically, the NCL material contains an average of about 66% water by weight. However, NCL is built with a higher density core and a higher hydration jacket, and hydration of NCL varies with age and from individual to individual. Therefore, it is not possible to assign a single water content value other than the average to NCL.

  Similarly, the various layers of NCL have different refractive indices. The refractive index of the lens varies from approximately 1.406 in the central layer of the lens to 1.386 in the less dense layer. For example, Hecht, Eugene. Optics, 2nd ed. (1987), Addison Wesley, ISBN 0-201-11609-X. p. See 178. Therefore, the optically meaningful equivalent refractive index, ie, ERI, is given as a characteristic of NCL. Both refractive index and water content change with lens age. Average ERI = 1.441-3.9 x 10 ^-4 x age, ie decreasing from about 1.441 at birth to about 1.414 at 70 years of age. M. See Dubbelman et al., "The Shape of the Aging Human Lens: Curvature, Equivalent Refractive Index And The Lens Paradox", Vision Research 41 (2001) 1867-1877, FIG.

  Furthermore, ERI increases by about 0.0013 to 0.0015 per diopter due to modulation. M. Dubbelman et al., “Change in Shape of the Aging Human Crystalline Lens With Accommodation”, Vision Research 45 (2005), 117-132 Ref pp. 127-128. This change in refractive index can be expected to be associated with a change (decrease) in water content due to lens deformation during accommodation. Ignoring these complex relationships, we will use the average ERI = 1.42 unless otherwise stated.

  It is interesting to see that finding a synthetic hydrogel with the same water content and at the same time the same refractive index as the NCL material appears to be very difficult, if not impossible. Specifically, synthetic hydrogels containing 66% by weight water typically have a refractive index of about 1.395 rather than the 1.42 expected for hydrogels containing water at a percentage close to 50%. Will have a rate.

  The average liquid content for ERI = 1.441 (very young average NCL) will be 40% water, but for ERI = 1.414 (geriatric average NCL) a water content of about 55% by weight It is necessary to have a hydrogel. Because we believe that mimicking the refractive index is more important than the water content of NCL for biosimilar IOL materials, a range of desirable average water content of the IOL according to a specific embodiment of the present invention Were selected to be 40% to 55% by weight. Of course, this is an average water content similar to that of NCL, and the lens has different layers with different water content, eg an inner layer with higher refractive index and an outer layer with lower refractive index Good.

  Although some prior art references cite IOLs with hydrogels with high water content, they do not recognize the relationship between water content and refractive index values. For example, Wichterle '732 specifies that the desired refractive index value is around 1.4 (broadly 1.37 to 1.45, for known synthetic hydrogels whose water content is specified) Clearly impossible, at least 60%, preferably 65 to 70%, with a refractive index range of 1.39 to 1.405). The examples show formulations having a low content of carboxylate groups.

  Sulc et al. '083 and' 903 disclose a water content of at least 70%, preferably at least 90% at the surface or part thereof and mentions 55-70% water content in prior art IOLs . A core with a higher refractive index and a casing with a lower refractive index are mentioned, the core may have the form of a Fresnel lens. Gradients for both hydration and refractive index are sometimes obtained by NaOH treatment to achieve reorganization of the hydrogel covalent network. Example 1 of this reference shows an IOL having a moisture content of 88.5%, Example 2 shows an IOL having a moisture content of 81%, and Example 4 shows a lens having a moisture content of 91%. Water content is not given for Example 3.

Charles Freeman, in US Patent Application Publication No. 2009/0023835, has a water content lower than 55%, a refractive index higher than 1.41, and a sodium ion flow rate of about 16 microequivalent-mm / hr / cm ² to about 20 A hydrogel material is described that is in the range of microequivalents-mm / hr / cm ² and is particularly useful for the posterior chamber IOL of the pterite. There is no mention at all of the carboxyl or acidic groups, but it is known that their presence increases the ion diffusion flux through the hydrogel.

  The hydrogel properties of NCL materials have some possible, less obvious but potentially important consequences, ie their water content depends on the pressure on the lens. As a result, far-range-tuned NCL may have a different water content than a near-field adjusted relaxed lens, and thus may have a different refractive index. Since the stress in NCL adjusted at long distances is not evenly distributed, a gradient of expansion and a gradient of refractive index may occur. This will cause subtle changes in optical properties in addition to the multifocality of the NCL surface. These subtle changes may be important for vision and to replicate those changes in ways other than by using hydrogels of similar physicochemical and optical properties and similar geometry to those of NCL Should be difficult. In particular, hydrogels of NCL substitutes should have similar refractive index and the ability to change the water content due to external stress that can reasonably be expected to act on NCL. Thus, the hydrogels used in biosimilar IOLs should have a hydraulic flow ability to water.

  Thus, at least the portion of the implant in contact with the posterior capsule is made of a transparent, flexible hydrogel material that approximates the optical, hydrophilic and electrochemical properties of the tissue forming the innate lens.

  The anterior portion of the IOL can prevent or even block the flow of the vitreous humor, which can result in increased IOP and ultimately glaucoma. This design often requires prior iris removal.

  The large area contact between the lens capsule and the artificial material used in the current IOL is sometimes a lens, unless manufactured from a highly biocompatible material with hydration and negative charge similar to NCL. It may cause capsular opacification, fibrosis and so on. These problems are currently solved by the biosimilar intraocular lens according to the present invention.

  The central optical part 2 has an equilibrium water content of about 35% to 65%, advantageously about 38% to 55%, and preferably about 40% to 50% (all percentages are by weight unless otherwise stated) The equilibrium water content is made of a deformable elastic material such as a hydrogel having a water content that is in equilibrium with the intraocular fluid.

  The deformability of the optic is weight for both the small incision implantation and its adjustment function. The optic can be constructed as a hydrogel shell with a core composed of liquid or soft gel as shown in FIG. 6A. FIG. 6A shows a cross-sectional view of a lens having a posterior hydrogel jacket 14, a softer core 15 and an anterior shell 16. The rear hydrogel jacket 14 is advantageously integrated with the peripheral support portion 3 of the lens and comprises a negative fixed charge on at least its rear face. The core 15 may advantageously be made of a hydrophobic liquid, such as mineral oil or silicone oil, or of a soft silicone or an acrylic, slightly cross-linked gel which can easily be designed and produced by a person skilled in the art. Alternatively, the core can be made of hydrophilic fluid or soft hydrogel. The anterior shell 16 may be made of the same or different material as the posterior hydrogel jacket 14.

  In one embodiment, the hydrogel jacket and soft core 15 have essentially the same refractive index so that the main part of refraction is performed at the external optical surface of the lens rather than at the internal contact surface. This can be accomplished, for example, by making the core from a silicone fluid or gel having a refractive index around 1.42 and making a jacket from a hydrogel having a water content of about 41% to 45% water. By correctly formulating the hydrogel, one skilled in the art can adjust the water content in the hydrogel to achieve a substantial reconciliation of refractive index. Alternatively, the core and the jacket may have different refractive indices so that part of the refraction takes place on the internal interface between the materials.

  FIG. 6B, a cross-sectional view of a lens having an internal contact surface between core 15 and an adjacent optical medium 16, shaped to form a complex lens, eg, a Fresnel lens. The material of the core 15 and the material of the optical medium 16 have different refractive indices, one of which is advantageously a fluid capable of improving both deformability and refraction. The areas 15 or 16 (which have the lower refractive index) can be produced by modifying the hydrogel according to the invention with a laser. The hydrogel correction can be performed preoperatively or postoperatively. The advantage of this arrangement is that it can use hydrogels with high water content and low refractive index as the basic building material and achieve a relatively low center thickness of the lens which allows implantation by small incisions It is.

  FIG. 6C shows an alternative design of a lens comprising two different materials. The material at the posterior side 14 is a hydrogel with a high hydration rate and containing negatively charged groups. This is the same for the optical part and the support part. The front side material of the core 15 is a material with lower water content and higher refractive index. The contact surfaces of these two materials are refractive.

  Both the central optical front surface 8A and the central rear optical surface 8B have a diameter of greater than about 5.6 mm, advantageously a diameter of greater than about 6.5 mm, preferably a diameter of greater than about 7.2 mm. The larger optimum diameter of these two optical surfaces is greater than about 7.5 mm, preferably about 8 mm, to approximate the size of the NCL optical system. Such a large optical system is usually suitable for the uneven or plano-convex central optical part 2. For biconvex optical portions, the anterior optical diameter is usually chosen smaller to minimize the center thickness of the optical portion. In any case, the diameter of the front optical surface 8A is preferably not larger than the diameter of the central rear optical surface 8B.

  The central optical surfaces 8A and 8B are surrounded by boundaries 9A and 9B which are not necessarily circular. Also, the boundaries 9A and / or 9B may be elliptical or have the shape of a truncated circle to facilitate lens folding and implantation by a small incision. Non-circular optical surfaces are particularly suitable for lenses having cylindrical components.

  The rear peripheral support surface 11B is formed by a convex surface, preferably a hyperboloid or conical surface having an axis which is identical to the main optical axis 1A. This surface is highly hydrophilic and has a negative fixed charge because of the content of the acid group such as carboxylate group, sulfo group, sulfate group or phosphate group. This combination of hydration and negative charge prevents permanent adhesion to the lens capsule, prevents migration of cells, especially fibroblasts, along the interface between the lens and lens capsule and reduces irreversible protein adsorption To prevent fibrosis and opacity of the lens capsule. The posterior peripheral surface is advantageously limited by the sharp edge 7B which further impedes cell migration towards the optical zone.

  The front peripheral support surface 11A is a concave surface, the top of which is located on the optical axis, and is preferably symmetrical along the axis 1A. Advantageously, it is a conical surface or hyperboloid whose axis coincides with the main optical axis 1A. Said surface is preferably highly hydrophilic and has a negative fixed charge, in order to prevent cell adhesion and migration as well as anterior capsule fibrosis. The anterior peripheral surface is advantageously limited by the sharp end 7A which further impedes cell migration.

  The front peripheral support surface 11A and the rear peripheral support surface 11B together with the connecting surface 6 define the shape of the peripheral support portion 3. The peripheral support portion is convex on the rear side and concave on the front side, and the average distance between these two surfaces is about 0.05 mm to 1 mm, advantageously about 0.1 mm to 0.6 mm, preferably It is in the range of about 0.15 mm to 0.35 mm. The optimum distance varies depending on the moisture content, the negative charge density, the crosslink density and the stiffness of the material depending on other parameters.

  If the back and front faces are formed by faces of similar geometry, such as hyperboloids, the peripheral support portion 3 will have the same thickness. The arrangement shown in FIG. 7A is easily deformable and adjustable for various sizes of lens capsule and has the advantage that the two sharp ends 7A and 7B prevent the migration of fibroblasts towards the optical zone .

  Also, the peripheral support portion 3 can be made more or less deformable by increasing or decreasing its thickness from edge to center, as shown respectively in FIGS. 7B and 7C. Also, these figures show various alternative arrangements of the ends 7A and 7B.

  The front face 4 of the implant is shaped to avoid any permanent contact with the iris which may cause iris erosion, pupil block, iris pigment transfer to the implant and other problems. Such contact may also cause a reversal of intraocular pressure as it impedes the flow of intraocular fluid. Such contact may also interfere with pupil contraction so as to prevent so-called near-field miosis, which compensates for the short focus with both the natural lens and the implant according to the invention. Thus, the front central optical surface 8A portion is partially sunk by the front peripheral support surface 11A recess and by the positioning of the boundary 9A under the plane defined by the front boundary 7A. The central anterior surface 8A has a front apex 10A, advantageously no more than the upper edge, preferably more than about 0.25 mm below the highest point of the lens (the higher of 7A and 7B), preferably at least 0. It is a flat, convex or concave surface with a front apex 10A of about 1 mm below the highest point 7A.

  At least the main part of both the front face 4 and the rear face 5 (including the central optical faces 8A and 8B) is defined by rotation of one or more conical cross sections around the main optical axis 1A, where the term " Conical section "includes a line segment for the purpose of the present application. The plane defined by said rotation will comprise a plane perpendicular to the conical plane which is symmetrical by the axis and the main optical axis 1A. The peripheral support portion is convex on the rear side and concave on the front side, and the average distance between these two surfaces is about 0.05 mm to 1 mm, advantageously about 0.1 mm to 0.6 mm, preferably It is in the range of about 0.15 mm to 0.35 mm.

  In at least one embodiment, a lens according to the present invention is manufactured by solidification of a liquid polymer precursor. In a preferred embodiment, solidification takes place in contact with a mold of solid type, in particular of a hydrophobic plastic. It goes without saying that the surface microstructure of the polymer depends on the environment in which its solidification takes place. The surface microstructure is different when solidification occurs at the solid-liquid interface and when solidification occurs at the liquid-liquid interface or the liquid-gas interface. Preferably, at least all optical surfaces are made by solidification of the precursor on a solid interface. Even more preferably, the entire surface of the implant is formed by solidification of the liquid precursor to a solid surface, in particular a hydrophobic plastic surface. The preferred plastic for the mold is a polyolefin, and a particularly preferred plastic is polypropylene. Polyolefins have low polarity and low interaction with highly polar monomers used as hydrogel precursors. Similarly, the hydrogel formed by liquid precursor solidification has very low adhesion to the mold surface and can be peeled off cleanly without even surface damage at the microscopic level. This is important for both the optical properties of the implant and the long-term biocompatibility.

  It is difficult to manufacture relatively large lenses of precise shape by molding. It is appreciated by those skilled in the art that solidification of liquid precursors is all accompanied by volumetric shrinkage that may even exceed 20 percent. With a constant volume closed mold, such shrinkage would cause vacuoles, bubble formation, surface deformation and other defects as it impedes replication of the inner mold surface. This is the main reason for using the meniscus casting method for IOL molding. Other inventors have reported methods and mold designs that allow excess monomer to be transported from the adjacent space by the suction created by volumetric contraction (Shepherd T., US Pat. No. 4,815,690). issue). However, this method can not be used if the liquid precursor gels with low conversion (eg, 5 to 10 percent) for cross-linking polymerization.

  The inventors have found a different way of compensating for the volume shrinkage, i.e. the reduction of the internal mold cavity volume by the deformation of a certain mold part. The mold depicted in FIG. 8 is made up of two parts 18A and 18B, part 18A being used for molding of the front face 4 and part 18B being used for molding of the back face 5.

  The shaped surface 19B of the portion 18B has the shape required to form the back optical surface 8B of the lens. The peripheral portion 22B of the shaping surface has a diameter greater than that of the lens, and preferably has the shape of a hyperboloid or cone.

  The portion 18A has a central portion 21A which forms the front optical surface 8A of the lens, and a shaped surface 19A which is divided into peripheral portions 22A of a diameter which is larger than the diameter of the lens. The peripheral portion 22A preferably has the shape of a hyperboloid or a cone. The peripheral surface 22A is substantially parallel to the corresponding surface 22B of the portion 18B.

  The diameter during molding of the shaped parts 18A and 18B is larger than the diameter of the lens, preferably they are the same. One of the faces for 22A or 22B comprises a relatively thin deformable barrier 20 with an inner surface corresponding to the geometry of the face 6 of the lens. The height of portion 20 is typically about 0.05 mm to 1.3 mm, and its thickness is less than its height. The sides of the part 20 are preferably bowl-shaped or triangular. At least one of the faces of the part 20 is preferably parallel to the optical axis 1A. The barrier 20 may be separate from the parts 18A and 18B but is advantageously an integral part of one of the parts. Advantageously, this portion 20 is located on the concave surface 22B. In the preferred mode of operation, the liquid precursor is filled into the concave molding portion 18B in a slight excess to reach the barrier 20 and then covered by the portion 18A. The mold is constructed such that only contact between portion 18A and portion 18B results in the formation via portion 20. Solidification of the precursor causes its contraction, which results in a reduction of pressure in the molding cavity. At low conversion rates, additional liquid precursor is drawn into the forming cavity. Once the gel point is reached by crosslinking, the precursor can not flow any further. The reduced pressure induces deformation of portion 20, resulting in a decrease in the distance between portions 18A and 18B, resulting in a decrease in the shaped cavity volume. The two-part mold for an IOL according to the invention is preferably made by injection molding from a polyolefin, preferably polypropylene.

  Preferred liquid precursors for the present invention are mixtures of acrylic and / or methacrylic monomers including crosslinkers, initiators and other components well known to those skilled in the art. Preferred precursor compositions are mixtures of acrylic monoesters and / or methacrylic monoesters and diesters of glycols, where the monoester is the hydrophilic component and the diester is the crosslinker. Preferred precursors also include acrylic acid and / or methacrylic acid or salts thereof. Preferred precursors also advantageously comprise UV-absorbing molecules having polymerizable double bonds, such as methacryloyloxybenzophenone (MOBP). Other possible derivatives of acrylic or methacrylic acid are their esters, amides, amidines and salts.

  Also, part of the hydrogel structure is an ionic group having a negative charge such as carboxylate pendant group, sulfuric acid pendant group, phosphoric acid pendant group or sulfonic acid pendant group. The ionic groups can be introduced by copolymerization with suitable monomers carrying the above-mentioned groups, such as methacrylic acid or acrylic acid. In this case, the ionogenic functional groups are uniformly dispersed in the hydrogel. Particularly advantageous are hydrogels having ionogenic groups concentrated on the side with a gradient resulting mainly from expansion and charge density. Such gradients can be produced by post-treatment of molded lenses, for example by the method described in Stoy ‘208 and Sulc et al., US Pat. Nos. 5,0806,83 and 5,158,832.

  Other methods include, for example, grafting of monomers containing ionogenic groups on the lens surface. It will be appreciated that only part of the lens surface may be treated to contain high concentrations of ionogenic groups, or different parts of the surface may be treated in different ways.

  The lens according to the invention can be implanted in a deformed, partially dewatered state. Achieve controlled partial dehydration by contacting the lens with an appropriate aqueous solution of magnesium ions or physiologically acceptable salts of salts such as chlorides, sulfates or phosphates of monovalent ions such as sodium or potassium, etc. can do. The salt concentration can be adjusted to achieve hydration of about 15% to 25% by weight of the liquid. The lenses in hypertonic solution can be sterilized, preferably by autoclaving.

  Another method of preparing a hydrogel lens for implantation by reduced size dissection is glycerol or dimethyl dimethylate so that the plasticized hydrogel has a softening point above ambient temperature but below eye temperature It is a method of plasticizing hydrogel with non-toxic organic water miscible solvent such as sulfoxide. Such compositions and processes are described, for example, in Sulc et al., US Pat. No. 4,834,753, which is incorporated herein by reference.

  A lens according to at least one embodiment of the present invention is advantageously implanted in an osmotic non-equilibrium state so as to temporarily attach to tissue. This osmotic imbalance allows the lens to be centered by attaching the lens to the posterior capsule while the capsule shrinks around the lens. Once the lens is enveloped by the lens capsule, its position is stabilized. Osmotic non-equilibrium can be achieved in various ways, ie by immersing the lens in a hypertonic salt solution, for example in a solution of 10% by weight to 22% by weight, preferably 15% by weight to 19% by weight NaCl, prior to implantation. Prior to implantation, replace the water with a lower concentration of a water-miscible solvent such as glycerol or dimethylsulfoxide, or the lens in a state in which the ionogenic groups are not completely ionized, ie in the acidic state prior to neutralization Implantation can be achieved by spontaneously advancing neutralization in situ by cations from body fluids. The lens spontaneously reaches its osmotic non-equilibrium several hours to several days after implantation.

  The lens shape is preferably formed by cross-linking copolymerization of methacrylic acid ester and / or acrylic acid ester and salt in a closed two-part mold.

  By removing certain parts of the lens, it is possible to adjust the shape of the lens after molding, for example by cutting out a part of the support part, by drilling holes in the lens outside the optical zone, etc. Shape adjustment may be done in the hydrogel state or in the xerogel (i.e. non-hydrated) state. We use the negatively charged hydrogel material to even allow the use of methods developed primarily for living tissue (including NCR), such as ultrasound phacoemulsification, ablation or femtosecond laser treatment I found out. These methods allow shape control even in the fully hydrated hydrogel state. Femtosecond lasers may even be used for the formation of internal cavities of hydrogel lenses that may be used to form new refractive members in lenses, for example as refractive cylindrical lenses to compensate for astigmatism . If the one removed by shape adjustment (e.g. by laser treatment) is water soluble and substantially non-toxic, such optical adjustment may possibly be achieved even after surgery in situ. The composition of the hydrogel in at least the treated part of the lens should preferably be based on an ester of polymethacrylic acid. Such polymers are known to be capable of depolymerizing into their parent monomers (such as 2-hydroxyethyl methacrylate or methacrylic acid), which are well soluble, easily diffusing, low toxicity compounds. Other polymers, such as polyacrylates, polyvinyl compounds or polyurethanes do not have this advantage.

  The invention is further illustrated by the following examples which are meant to provide additional information without limiting the scope of the invention.

  The following monomer mixture was prepared: 98 parts by weight of 2-hydroxyethyl methacrylate (HEMA), 0.5% by weight of triethylene glycol dimethacrylate (TEGDMA), 1% by weight of methacryloyloxybenzophenone (MOBP), 1% by weight Methacrylic acid, 0.25% by weight camphorquinone (CQ) and 0.05% by weight trieethanolamine (TEA). This mixture is degassed by the use of carbon dioxide, 18B being part of the mold for molding the posterior lens surface and 18A being part of the mold for shaping the anterior portion of the lens surface, FIG. The two-part plastic mold shown schematically in FIG. Both parts are injection molded from polypropylene (PP). The shaped surface 19B of the portion 18B has a shape formed by two concentric hyperboloids. The central part of this surface is a hyperboloid with a diameter of 3 mm, a central radius of 3.25 mm and a conic constant of -3.76, and the peripheral part with a central radius of 3.25 mm and a conic constant of -6.26. The mold surface is provided with a projecting circular barrier 20 with a diameter of 8.5 mm and a height of 0.2 mm with asymmetric triangular sides. This edge is designed to form the connection surface 6 in FIG. 3A.

  The portion 18A has a shaped surface 19A divided into a central portion 21 of diameter 6.8 mm and a peripheral portion 22A of diameter 13 mm. The peripheral part is formed by a hyperboloid having a central radius of 3.25 mm and a conic constant of -6.26. The peripheral hyperboloid is parallel to the corresponding surface of the portion 18B. The central portion of the portion 18A has a central radius of curvature of -20 mm and a conical constant h = 1.

  About 0.1 ml of the monomer mixture is pipetted into portion 18B and then carefully centered and covered with a lightly pressed portion 18A of small weight. The only direct contact between the parts is a circular contact between the barrier 20 and the peripheral part of 22A. The mold is then irradiated with blue light at a wavelength of 471 nm for 10 minutes. This light initiates the polymerization of the monomer with a relatively low conversion gelation and volume shrinkage roughly proportional to the conversion. The contraction of the soft gel creates a weak vacuum which draws both parts of the mold together. The conical peripheral portion 22A of the mold 18A squeezes the barrier 20 causing it to deform slightly and approach the portion 18B to reduce the volume of the molding cavity. This compensates for the volumetric shrinkage due to polymerization. The above mold design is particularly suitable for producing relatively bulky IOLs from materials with relatively low conversion rates and high polymerization shrinkage reaching gel points.

  The mold parts are separated and the xerogel lens, which is an exact duplicate of the molding cavity, is neutralized with sodium bicarbonate solution and extracted with an isotonic solution. The linear expansion coefficient between the xerogel lens and the hydrogel lens is 1.17. After evaluation of the optical properties, the lenses were immersed in 18% by weight aqueous NaCl solution in a sealed blister package and sterilized by autoclave.

  The following monomer mixture was prepared: 94 parts by weight of 2-hydroxyethyl methacrylate (HEMA), 0.5% by weight of triethylene glycol dimethacrylate (TEGDMA), 4.5% by weight of methacryloyloxybenzophenone (MOBP), 1 Wt% methacrylic acid and 0.25 wt% dibenzoyl peroxide. The mixture was degassed using nitrogen and carbon and filled into a two-part plastic mold as schematically shown in FIG. The shaped surface 19B of the portion 18B has a shape formed by two concentric planes. The central part of this surface is hyperbolic with a diameter of 3 mm, a central radius of 3.00 mm and a conic constant of 1, and the peripheral part with a central radius of 3.25 mm and a conic constant of -6.26. The mold surface is provided with a projecting circular barrier 20 of diameter 8.8 mm and height 0.15 mm with asymmetric triangular sides. The inner surface of the barrier 20 is designed to form the connecting surface 6 in FIG. 3A.

  The portion 18A has a shaped surface 19A divided into a central portion 21 of diameter 7.1 mm and a peripheral portion 22A of diameter 13 mm. The peripheral part is formed by a hyperboloid having a central radius of 3.25 mm and a conic constant of -6.26. The peripheral hyperboloid is parallel to the corresponding surface of the portion 18B. The central portion of the portion 18A is a plane perpendicular to the optical axis 1A.

  About 0.1 ml of the monomer mixture is pipetted into portion 18B and then carefully centered and covered with a lightly pressed portion 18A of small weight. The only direct contact between the parts is a circular contact between the barrier 20 and the peripheral part of 22A. The mold is then heated to 75 ° C. for 6 hours.

  The mold part is separated and the xerogel lens, which is an exact duplicate of the molding cavity, is neutralized with sodium bicarbonate solution and extracted three times with ethyl alcohol and five times with isotonic solution. The lens was yellow due to complete absorption of UV light and some of the blue visible light. The linear expansion coefficient between the xerogel lens and the hydrogel lens is 1.13. After evaluation of the optical properties, the lenses were immersed in a 15% by weight aqueous NaCl solution in a sealed blister package and sterilized by autoclave.

  The following monomer mixture was prepared: 94.5 parts by weight of 2-hydroxyethyl methacrylate (HEMA), 0.5% by weight of triethylene glycol dimethacrylate (TEGDMA), 5% by weight of methacryloyloxybenzophenone (MOBP) and 0 .25% by weight of dibenzoyl peroxide. The mixture was degassed using nitrogen and carbon and filled into a two-part plastic mold as schematically shown in FIG. The shaped surface 19B of the portion 18B has a shape formed by two concentric planes. The central part of this surface is hyperbolic with a diameter of 6.5 mm, a central radius of 4.5 mm and a conic constant of 0, and the peripheral part with a central radius of 4.25 mm and a conic constant of -8. The mold surface is provided with a projecting circular barrier 20, with a diameter of 9.3 mm and a height of 0.35 mm, with asymmetric triangular sides. The inner surface of the barrier 20 is designed to form the connecting surface 6 in FIG. 3A.

  The portion 18A has a shaped surface 19A divided into a central portion 21 of diameter 6.4 mm and a peripheral portion 22A of diameter 13 mm. The peripheral part is formed by a hyperboloid having a central radius of 4.25 mm and a conic constant of -8. The peripheral hyperboloid is parallel to the corresponding surface of the portion 18B. The central portion of the portion 18A is a surface having a diameter of 6.4 mm, a central radius of -3.75 mm and a conic constant of -6.

  About 0.1 ml of the monomer mixture is pipetted into portion 18B and then carefully centered and covered with a lightly pressed portion 18A of small weight. The only direct contact between the parts is a circular contact between the barrier 20 and the peripheral part of 22A. The mold is then heated to 75 ° C. for 6 hours.

  Separate the mold parts and withdraw the xerogel lens, which is an exact duplicate of the mold cavity. The lens is then treated with the quaternary base described in reference Stoy, ‘208.

  The z-lens with a transparent electrically neutral cross-linked hydrophilic polymer has a surface created by a gradient layer with high hydration and negative charge density. The lens is neutralized with sodium bicarbonate solution and extracted three times with ethyl alcohol and five times with isotonic solution. The lens was clear and completely absorbed UV light. The linear expansion coefficient between the xerogel lens and the hydrogel lens is about 1.12. After evaluation of the optical properties, the lenses were immersed in an isotonic aqueous NaCl solution in a sealed blister package and sterilized by autoclave.

The Raman spectra of the hydrogels of the invention show a marked difference between the original hydrogel and the same hydrogel that has been subjected to TPA by exposure to a focusing beam of a femtosecond laser. That is, the ratio of the signal at 2945cm ^-1 corresponding to the CH ₂ group of the derived signals and the polymer backbone of at 3420cm ^-1 corresponding to the water there is a significant difference (see FIG. 9A). The ratio between the intensities of the two peaks is proportional to the phase shift (in number of wavelengths) of the test laser beam between the correction material and the non-correction material (see FIG. 9B). Raman scans on the modified strips in the hydrogel showed increased water content in the treated area (see FIG. 9C). On the other hand, Raman spectra in the region below 2000 cm ^-1 did not show any indication of new chemical groups. This is consistent with the partial replacement of the polymer mass by the aqueous liquid through mechanisms such as depolymerization.

  These and other advantages of the present invention should be apparent to those skilled in the art from the foregoing specification. Therefore, it should be recognized by those skilled in the art that changes or modifications can be made to the above embodiments without departing from the broad inventive concept of the present invention. Thus, the present invention is not intended to be limited to the particular embodiments described herein, but is intended to encompass all changes and modifications that are included in the scope and spirit of the present invention as set forth in the claims. It goes without saying that it is intended.

Claims (55)

  Covalently crosslinked comprising a polymer comprising a derivative of (meth) acrylic acid and / or monomeric units of (meth) acrylic acid and comprising a UV absorbing dopant moiety and an activator moiety which is negatively charged at physiological pH Hydrogel, wherein exposing the fully hydrated hydrogel to electromagnetic radiation results in two-photon absorption which induces one or more structural changes and a negative change in refractive index in the hydrogel , Hydrogel.   The hydrogel according to claim 1, wherein the dopant moiety and the activator moiety are pendant groups of a polyacrylate polymer or polymethacrylate polymer.   The hydrogel of claim 1 wherein the dopant moiety is a UV absorbing compound that does not strongly absorb light of about 400 nm wavelength.   The hydrogel according to claim 1, wherein the dopant moiety is a compound selected from the group consisting of rhodamine, benzophenone, coumarin, fluorescein, benzotriazole and derivatives thereof.   The hydrogel according to claim 1, wherein the activator moiety is a compound containing a carboxylate group, a sulfonate group, a sulfate group or a phosphate group.   The hydrogel of claim 1, wherein the one or more structural changes comprise partial depolymerization of the hydrogel.   7. The hydrogel of claim 6, wherein the partial depolymerization forms a water-filled void in the hydrogel.   7. The hydrogel of claim 6, wherein the depth of the partial depolymerization of the hydrogel is dependent on the accumulated energy absorbed at the predetermined position of the hydrogel.   An ophthalmic implant comprising the hydrogel according to claim 1. An in situ controlled cross-linked hydrogel ophthalmic implant comprising a hydrogel of an acrylate or methacrylate copolymer, wherein said copolymer comprises at least four comonomers:
a) acrylic or methacrylic esters comprising at least one pendant hydroxyl group;
b) polyol acrylic esters or amides or polyol methacrylic esters or amides with at least two acrylate or methacrylate groups per polyol ester or amide;
c) a derivative of acrylic acid or methacrylic acid having at least one pendant carboxyl group; and d) a vinyl-based monomer having a pendant UV-absorbing group, an acrylic monomer or a methacrylic monomer,
The refractive properties of the implant are adjusted by controlling the absorption of target electromagnetic radiation by the hydrogel, resulting in a negative change in refractive index at selected locations of the implant.
Hydrogel ophthalmic implant.   11. The hydrogel ophthalmic implant according to claim 10, wherein the implant has anterior and posterior refractive surfaces forming a lens with positive or negative refractive power.   The hydrogel ophthalmic implant according to claim 11, wherein the lens is a matrix inner lens.   The hydrogel ophthalmic implant according to claim 11, wherein the lens is an anterior chamber lens.   The hydrogel ophthalmic implant according to claim 11, wherein the lens is a non-quartz lens for placing between an iris and a natural lens.   The hydrogel ophthalmic implant according to claim 11, wherein the lens is a posterior chamber lens for at least partially replacing a native lens.   The hydrogel ophthalmic implant according to claim 11, wherein at least one of the refractive surfaces is a non-spherical surface having negative spherical aberration.   11. The hydrogel ophthalmic implant according to claim 10, wherein the monomer comprising pendent carboxyl groups is neutralized or partially neutralized methacrylic acid.   11. The hydrogel ophthalmic implant of claim 10, wherein the monomer comprising the pendant carboxyl group is present at a concentration of 0.1 mole% to 5 mole% based on total monomer units of the copolymer.   19. The hydrogel ophthalmic implant of claim 18, wherein the monomer comprising the pendant carboxyl group is present at a concentration of 0.5 mol% to 2 mol% based on total monomer units of the copolymer.   11. The hydrogel ophthalmic implant of claim 10, wherein the monomer comprising the pendant UV absorbing group is present at a concentration of 0.1 mole% to 5 mole% based on total monomer units of the copolymer.   21. The hydrogel ophthalmic implant of claim 20, wherein the monomer comprising the pendant UV absorbing group is present at a concentration of 0.2 mol% to 2.5 mol% based on total monomer units of the copolymer.   11. The hydrogel ophthalmic implant of claim 10, wherein the pendant UV absorbing group comprises a phenolic hydroxyl group conjugated to an aromatic group.   23. The hydrogel ophthalmic implant of claim 22, wherein the monomer comprising the pendant UV absorbing group is present at a concentration of 0.1 mole% to 5 mole% based on total monomer units of the copolymer.   11. The hydrogel ophthalmic implant according to claim 10, wherein at least one of the UV absorbing pendant groups is selected from the group consisting of derivatives of benzophenone, derivatives of benzotriazole, derivatives of coumarin and derivatives of fluorescein.   11. The hydrogel ophthalmic implant of claim 10, wherein the pendent carboxyl group and the pendent UV absorbing group are present in a molar ratio of about 0.25 to about 5.   26. The hydrogel ophthalmic implant of claim 25, wherein the pendant carboxyl groups and pendant UV absorbing groups are present in a molar ratio of about 0.5 to about 3.5.   The hydrogel ophthalmic implant according to claim 10, wherein the copolymer comprises at least two different comonomers comprising different UV absorbing groups.   The hydrogel ophthalmic implant according to claim 27, wherein at least one of the UV absorbing groups is benzophenone or a derivative thereof.   The ophthalmic lens according to claim 11, wherein the pendant carboxyl groups are ionized and the molar ratio of ionized pendant carboxyl groups to UV absorbing pendant groups is about 0.5 to about 3.5.   The ophthalmic lens according to claim 11, wherein at least a major portion of the polymer of the hydrogel is a hydrophilic derivative of methacrylic acid.   31. The ophthalmic lens according to claim 30, wherein at least a major portion of the hydrophilic methacrylic acid derivative is a glycol ester of methacrylic acid.   The ophthalmic lens according to claim 11, wherein the covalently cross-linked hydrogel comprises a liquid in a proportion greater than 30% by weight under equilibrium physiological conditions.   The ophthalmic lens according to claim 11, wherein the covalently crosslinked hydrogel comprises liquid in a proportion of less than 55% by weight under equilibrium physiological conditions.   The ophthalmic lens according to claim 11, wherein the covalently crosslinked hydrogel comprises a liquid in a proportion of 35% to 47.5% by weight under equilibrium physiological conditions.   The ophthalmic lens according to claim 11, wherein at least the posterior optical surface provides refraction at negative spherical aberration.   The ophthalmic lens according to claim 11, wherein the posterior optical surface provides refraction at negative spherical aberration of −0.1 microns to −2 microns.   The ophthalmic lens according to claim 36, wherein the negative spherical aberration is between -0.5 microns and -1.5 microns.   40. The ophthalmic lens of claim 37, wherein the negative spherical aberration is -0.75 microns to -1.25 microns at an aperture greater than 4.5 mm.   The ophthalmic lens according to claim 11, comprising both UV-absorbing pendant groups comprising phenolic hydroxyl groups conjugated with aromatic systems and UV-absorbing groups comprising benzotriazole structures.   40. The ophthalmic lens according to claim 39, wherein the UV absorbing pendant group comprising a phenolic hydroxyl group conjugated to the aromatic system and the UV absorbing group comprising the benzotriazole structure are located in separate layers of an ophthalmic lens. lens.   The ophthalmic lens according to claim 11, wherein the lens is embedded in the cornea.   The ophthalmic lens according to claim 11, wherein the lens is implanted in the anterior chamber of the eye between the cornea and the iris.   The ophthalmic lens according to claim 11, wherein the lens is a non-quartz lens embedded between an iris and a natural crystalline lens.   The ophthalmic lens according to claim 11, wherein the lens is embedded in the posterior chamber and at least partially substitutes for the native lens.   A method of adjusting the refractive properties of a fully hydrated hydrogel according to claim 1, comprising focusing irradiation of the hydrogel with electromagnetic radiation such that two-photon absorption occurs, wherein the polymer component of the hydrogel is The way in which depolymerization and / or decomposition take place. A method for the in situ adjustment of optical parameters of hydrogel ophthalmic implants, comprising
a) including in the eye the hydrogel ophthalmic implant according to claim 9; and b) irradiating a portion of the hydrogel ophthalmic implant with electromagnetic radiation using a femtosecond laser, Part of the copolymer of said hydrogel is depolymerized and / or degraded by
The method wherein the optical parameters of the implant are adjusted.   47. The method of claim 46, wherein the optical parameter comprises refractive index.   47. The method of claim 46, wherein the irradiation results in an elongated cavity or voxel inside a hydrogel ophthalmic implant.   49. The method of claim 48, wherein the voxel depth is greater than 10 microns.   49. The method of claim 48, wherein increasing the depth of the voxel increases the phase shift while the refractive index remains approximately constant.   51. The method of claim 50, wherein the refractive index is greater than or equal to 1.3335.   51. The method of claim 50, wherein the phase shift is 30 microns or less.   47. The method of claim 46, wherein the radiation-ablated depolymer comprises a soluble, easily diffusing, low toxic compound.   47. The method of claim 46, wherein the modified optical properties are provided by forming a pattern of elongated voxels of various depths in the hydrogel while keeping the modified refractive index substantially constant.   51. The method of claim 50, wherein the phase shift is controlled by changing voxel depth rather than by changing refractive index.