Note: Descriptions are shown in the official language in which they were submitted.
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LIGHT-ADJUSTABLE HYDROGEL AND BIOANALOGIC INTRAOCULAR LENS
CROSS-REFERENCE TO RELATED APPLICATIONS
This application claims priority to U.S. Application No. 15/190,715, filed
June 23,
2016, which is a Continuation-in-Part of U.S. Application Serial No.
14/760,868, filed on
July 14, 2015, which is the U.S. National Phase of International Application
No.
PCT/M2013/060869, filed on December 12, 2013, which claims the benefit of
priority of
U.S. Provisional Application No. 61/752,685, filed on January 15, 2013. The
entire contents
of all of the above applications are incorporated herein by reference.
FIELD OF THE INVENTION
This invention relates to a hydrogel implantable ophthalmic lens whose optical
parameters can be optimized and/or customized by a controlled absorption of
electromagnetic
radiation, such as laser radiation in the visible and/or near-IR region, and
more particularly
radiation emitted in pulses shorter than one nanosecond (so called Femtosecond
Laser, FSL)
resulting in a change of the refractive index of the irradiated hydrogel.
BACKGROUND OF THE INVENTION
Intra ocular lenses (IOLs) are surgically implantable lenses which replace or
supplement optical function of the NCL. So called "posterior chamber
intraocular lens", or
PC IOLs, replace the NCL in the case of cataract or, more recently, in the
case of presbyopia
by so called "clear lens exchange", or CLE. Other implantable lenses are
placed into the
anterior chamber of the eye (AC IOLs), into the cornea (corneal or
intrastromal implants) or
between the NCL and iris (so called "implantable contact lens" or ICL). So
far, most of these
IOLs were designed to replace or to supplement the basic optical function of
the NCL only.
It should be appreciated that an NCL in a human eye, depicted in the Fig. 1,
is a complicated
structure with several functions. The main eye parts include the cornea 101;
the iris 102; the
NCL 103; the posterior capsule 104; the cilliary muscle 105; the zonules 106;
the vitreous
body 107; and the retina 108.
The basic optical function of the NCL 103 consists in helping the cornea 101
to focus
the incoming light so that a distant object can be projected on the retina
108. The other
important optical function is accommodation ¨ adjustment of optical power of
the lens in
such a way that objects at various distances can be projected onto the retina
108. There are
several theories explaining the accommodation mechanism. See for example L.
Werner et al,
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Physiology of Accommodation and Presbyopia, ARQ. BRAS. OFTALMOL. 63(6),
DEZEMBRO/2000-503.
The most firmly established theory is von Helmholtz theory explaining that,
referring
to the Fig. 1, relaxed cilliary muscle 105 causes tension in the zonules 106
that pull the lens
103 periphery outward to keep the NCL 103 in its deformed (flattened) shape
that provides a
lower refractive power suitable for distant vision. Focusing on a near object
is caused by
tension in the cilliary muscle 105 that relaxes the zonules 106 and allows the
NCL 103 to
obtain its "natural" configuration with a smaller diameter, larger central
thickness and smaller
radii of curvature on both anterior and posterior surfaces. This increases the
NCL's refractive
power and allows for projection of the image of near objects on the retina
108.
Most of the common intraocular lenses have spherical surfaces that can be
manufactured rather readily. It has been assumed for some time that the NCL
103 is
essentially spherical. However, a spherical lens is not exactly monofocal,
instead it
demonstrates so called "spherical aberration" wherein rays incoming through
the center are
bent into a focal point that is slightly further from the lens than rays
incoming through the
lens periphery. Therefore, a spherical lens is somewhat more refractive in its
periphery than
in its center. This change is continuous: such a lens does not have a single
focal point, but
many focal points in a short interval of distances (focal range) between the
longest and
shortest focal distance. In other words, a spherical lens is negatively
polyfocal (its focal
distance decreases from the center to the periphery). Lenses with elliptical
rather than
spherical surfaces (such as surfaces created by solidification of a static
liquid meniscus) have
even more distinct spherical aberration and are, therefore, even more
negatively polyfocal
than spherical lens.
Some artificial intraocular lenses include hyperbolic surfaces alongside with
other
surfaces of second order, such as spheric or even elliptic surfaces that have
negative
polyfocality and very opposite optical effect. More importantly, the prior art
generally
combines second order (or conic section) surfaces with meniscoid surfaces that
are poorly
defined and merely approximate second order surfaces with positive spherical
aberration
(although never surfaces with hyperbolic aberration).
For example, Wichterle in US Pat. No. 4,971,732 claims the meniscoid surfaces
to
approximate a flat ellipsoid while Stoy in US Pat. No. 5,674,283 considers
meniscoid
surfaces an approximation of a spherical surface, both having negative
polyfocality. A
combination of surfaces with positive and negative polyfocality diminishes or
negates
advantages of the former.
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Furthermore, Wichterle '732 describes a manufacturing method of the
intraocular lens
where a monomer solidifies in an open mold, one (posterior) side of the lens
having the shape
of the mold cavity while the anterior side has a shape of a solidified liquid
meniscus
(presumably approximating a flat ellipsoid shape with negative polyfocality,
being
somewhere between purely spherical and purely ellipsoid surface). The mold
cavity has the
shape of a second order surface that may include a hyperbolic surface. One can
note that
each of the optical surfaces is created differently ¨ one by solidification of
a polymer
precursor against a solid surface while the other by solidification on the
liquid-gas interface.
It is known to those skilled in the art that the surface quality of the two
optical surfaces
formed under such different circumstances may differ profoundly in both
optical and
biological respects.
Wichterle in US Pat. No. 4,846,832 describes another manufacturing method of
the
intraocular lens where the posterior side of the lens has the shape of the
solidified liquid
meniscus (presumably approximating a flat ellipsoid shape with negative
polyfocality) while
the anterior side is formed as an imprint of the solid mold shaped as a second
order surface
that may implicitly include also a hyperbolic surface. Again, we can note that
each of the
optical surfaces is created differently ¨ one by solidification of a polymer
precursor against a
solid surface while the other by solidification on the liquid-gas interface.
Stoy '283 discloses modifying the method described by Wichterle '732 using a
two
part mold, one part being similar to the Wichterle's mold while the other
being used to form a
modified meniscoid of a smaller diameter on the anterior lens surface. The
meniscoid optical
surface is of the same character as the meniscoid resulting from Wichterle
'732, albeit of a
smaller diameter and, therefore, probably closer to a spherical surface than
an ellipsoid
surface. In any case, such a surface has negative polyfocality. The posterior
side is formed as
an imprint of the solid mold shaped as a second order surface that may include
a hyperbolic
surface while the other optical surface is formed by solidification of the
liquid polymer
precursor on the liquid-gas interface.
Michalek and Vacik in PCT/CZ2005/000093 describe an IOL manufacturing method
using a spin-casting method in open molds. Molds filled with monomer mixture
spin along
their vertical axis while polymerization proceeds. One of the optical surfaces
is created as the
imprint of a solid mold surface while the other is formed by the mold
rotation. The imprinted
surface has the shape formed by rotation of the conic section along the
vertical axis (which
may include hyperboloid shape). The other surface is shaped as a meniscoid
modified by the
centrifugal force that will transfer some of the liquid precursor from the
center toward the
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periphery. In the case of the convex meniscus, the centrifugal force will
flatten the center and
create a steeper curvature in the periphery, i.e. increase the spherical
aberration of the surface.
In the case of a convex meniscus, the centrifugal force will create a meniscus
with smaller
central radius and modify the surface to approximate something between spheric
and
parabolic shape. In any case, the hyperbolic aberration cannot be achieved for
either a convex
or concave meniscoid surface.
Sulc et al. in US Pat. Nos. 4,994,083 and 4,955,903 discloses an intraocular
lens with
its anterior face protruding forward in order to be in permanent contact with
the iris that will
center the lens. Both posterior and anterior surfaces may have the shape
obtained by rotation
of a conical section around of the optical axis (sphere, parabola, hyperbole,
ellipse). The iris-
contacting part of the lens is a hydrogel with very high water content (at
least 70% and
advantageously over 90% of water) that is inherently soft and deformable.
Therefore, the
optical surface deformed by the contact with iris cannot be exactly a conic
section surface,
but a surface with a variable shape that will depend on the pupil diameter,
probably close to a
sphere with a somewhat smaller central radius. Namely, this situation is
similar to the lens
from another reference that achieves decrease in the central diameter by
pressing a
deformable gel-filled lens against a pupil-like aperture in an iris-like
artificial element (Nun
in US Pat. No. 7,220,279). Nun '279 does not mention or imply use of
hyperboloid optical
surfaces. Cummings in US Pat. Publ. Nos. 2007/0129800 and 2008/0269887
discloses a
hydraulic accommodating IOL in which a liquid is forced into the internal IOL
chamber by
action of cilliary apparatus causing thus change of the optical surface and
accommodation.
Hong et al. in US Pat. No. 7,350,916 and US Pat. Publ. No. 2006/0244904
disclose an
aspheric intraocular lens with at least one optical surface having a spherical
aberration in
order to compensate the positive spherical aberration of the cornea. The
negative spherical
aberration is achieved by hyperbolic shape of the optical surface.
Hong et al. in US Pat. Publ. No. 2006/0227286 discloses optimal IOL shape
factors
for human eyes and defines the optimum lens by a certain range of "shape
factors" from -0.5
to +4 (the shape factor being defined by Hong as the ratio of sum of anterior
and posterior
curvatures to their difference), and at least one of the optical surfaces is
advantageously
aspherical with conic constant between -76 and -27.
Hong et al. in US Pat. No. 7,350,916 describes an IOL with at least one of the
optical
surfaces having a negative spherical aberration in a range of about -0.202
microns to about
-0.190 microns across the power range.
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SUMMARY OF THE INVENTION
In at least one aspect, the present invention provides an artificial lens
implantable into
the posterior chamber of human eye for replacement of the natural crystalline
lens, the lens
(referring to the Fig. 3) having a main optical axis 1A; the central optical
part 2 and the
peripheral supporting part 3; the overall shape of the implant being defined
by its anterior
surface 4, posterior surface 5 and the transition surface 6 between the upper
boundaries of the
anterior and posterior surfaces 7A and 7B of the implant; having the central
anterior optical
surface 8A with boundary 9A and anterior apex 10A; the central posterior
optical surface 8B
with boundary 9B and posterior apex 10B; and anterior peripheral supporting
surface 11A
and posterior peripheral supporting surface 11B.
Artificial lens implantable into the posterior chamber of human eye for
replacement of
the natural crystalline lens that simulates as closely as practicable the
shape, size, optical
properties and material properties of an NCL while respecting the need for
surgical
implantation through a small incision.
The artificial lens according to at least one embodiment of the invention has
at least
the posterior surface approximating the shape and size of the posterior
surface of the natural
lens in order to achieve substantially complete contact with the posterior
capsule of the eye.
As defined in this context, the term"substantially" can mean that either 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 (that has diameter larger that the lens) is in contact
with the posterior
surface of the lens. At least the part of the artificial lens according to the
invention that is
contacting the posterior capsule is made from a transparent flexible hydrogel
material
approximating the optical, hydrophilic and electrochemical character of tissue
forming the
natural lens. The anterior side is designed to avoid a permanent contact with
iris.
In at least one embodiment, the anterior surface is shaped to avoid a
permanent
contact with the iris with the anterior peripheral supporting surface 11A
being concave.
In at least one embodiment, the artificial lens according to the invention has
at least the major
parts of its anterior and posterior surfaces, including both optical surfaces,
defined by rotation
of one or more conic sections along the optical axis and formed by
solidification of a liquid
polymer precursor in contact with a solid wall of a mold, preferably a
hydrophobic plastic
mold.
One aspect of the invention is directed to a hydrogel comprising UV-absorbing
dopant
moieties and activator moieties that are negatively charged at physiological
pH, where
exposure of the fully hydrated hydrogel to electromagnetic radiation results
in two-photon
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absorption which causes one or more structural changes in the hydrogel and a
change in the
refractive index. In one embodiment the hydrogel is a covalently crosslinked
hydrogel. In
one embodiment, the structural change is achieved without substantially
changing the volume
of the treated hydrogel segment after it gets into equilibrium with the liquid
medium around
the lens. One convenient method to determine the volume change can consist in
the following
procedure: several equal zones in the hydrogel (e.g., 50 x 50 microns) are
treated with various
laser settings to achieve different phase shift in different zones. After
allowing time sufficient
to achieve equilibrium, the linear dimension of the treated zones does not
change by more
than 20% from the original zone dimensions, the dimension change being less
than 10% for
most conditions. Since the depth of the treated zone cannot be readily or
directly measured, it
is assumed that the change of the volume is isotropic and the change of
treated volume is due
to the same relative expansion or contraction for all dimensions. One
embodiment of the
hydrogel comprises a polymer comprising monomer units of (meth)acrylic acid
derivatives
and/or (meth)acrylic acid. In one embodiment of the hydrogel the dopant
moieties and
activator moieties are pendant groups on a polyacrylate 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 rhodamines, benzophenones, coumarins,
fluoresceins,
benzotriazoles, and derivatives thereof In one embodiment, the UV absorbent
moiety
contains a carbonyl group conjugated with an aromatic system, a phenolic
hydroxyl group
conjugated with an aromatic system, or advantageously both carbonyl and
phenolic hydroxyl
groups conjugated with an aromatic system. In one embodiment the activator
moiety is a
compound comprising a carboxylate group, sulfonate group, sulfate group,
phenolate group
or phosphate group. In one embodiment of the hydrogel the one or more
structural changes
comprise partial depolymerization of the hydrogel. In one embodiment the
partial
depolymerization forms an aqueous-filled void in the hydrogel. In one
embodiment of the
hydrogel the change in the refractive index is a negative change. In one
embodiment the
depth of the partial depolymerization of the hydrogel depends on the
cumulative energy
absorbed in a given location 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 aspect of the invention is directed to an ophthalmic implant
comprising the
above hydrogel.
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Yet another aspect of the invention is directed to an in situ-adjustable
hydrogel
ophthalmic implant comprising an acrylate or methacrylate copolymer hydrogel
where the
copolymer comprises at least four co-monomers: a) an acrylate or methacrylate
ester
containing at least one pendant hydroxyl group; b) a polyol acrylate ester or
polyol
methacrylate ester or amide with at least 2 acrylate or methacrylate groups
per polyol ester or
amide; c) a derivative of acrylic or methacrylic acid having at least one
pendant carboxyl
group; and d) a vinyl, acrylic or methacrylic monomer having a pendant UV-
absorbing
group; where the refractive properties of the implant are adjusted by a
controlled absorption
of targeted electromagnetic radiation by the hydrogel resulting in a change of
refractive index
in selected locations of the implant. In one embodiment the ester of component
a) carries at
least one pendant hydroxyl group on the alcohol portion of the ester. In one
embodiment the
copolymer is covalently crosslinked. In one embodiment the implant has
anterior and
posterior refractive surfaces forming a lens with positive or negative
refractive power. In one
embodiment the lens is an intrastromal lens. In another embodiment the lens is
an anterior
chamber lens. In yet another embodiment the lens is a phakic lens for location
between the
iris and the natural crystalline lens. In a further embodiment the lens is a
posterior chamber
lens for at least partial replacement of the natural crystalline lens. In one
embodiment of the
hydrogel ophthalmic implant at least one of the refractive surfaces is an
aspheric surface with
negative spherical aberration. In another embodiment of the hydrogel
ophthalmic implant the
monomer containing the pendant carboxyl group is a neutralized or partially
neutralized
methacrylic acid. In one embodiment the monomer containing the pendant
carboxyl group is
present in a concentration between 0.1 molar % and 5 molar % based on all
monomer units of
the co-polymer. In another embodiment the monomer containing the pendant
carboxyl group
is present in a concentration between 0.5 molar % and 2 molar % based on all
monomer units
of the co-polymer. In one embodiment the monomer containing the pendant UV-
absorbing
group is present in a concentration between 0.1 molar % and 5 molar % based on
all
monomer units of the co-polymer. In another embodiment the monomer containing
the
pendant UV-absorbing group is present in a concentration between 0.2 molar %
and 2.5
molar % based on all monomer units of the co-polymer. In one embodiment of the
hydrogel
ophthalmic implant the pendant UV-absorbing group contains a carbonyl group
conjugated to
an aromatic group; in one embodiment the monomer containing the pendant UV-
absorbing
group is present in a concentration between 0.1 molar % and 5 molar % based on
all
monomer units of the co-polymer. In one embodiment of the hydrogel ophthalmic
implant at
least one of the UV-absorbing pendant groups is selected from the group
consisting of
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derivatives of benzophenone, derivatives of benzotriazole, derivatives of
coumarin and
derivatives of fluorescein. In one embodiment the pendant carboxyl groups and
pendant UV-
absorbing groups are present in a molar ratio between about 0.25 and 5; in
another
embodiment the pendant UV-absorbing groups are present in a molar ratio
between about 0.5
and about 3.5. In one embodiment of the hydrogel ophthalmic implant the
copolymer
contains at least two different comonomers containing different UV-absorbing
groups; in one
embodiment at least one of the UV-absorbing groups is a benzophenone.
In one embodiment of the ophthalmic lens the pendant carboxyl group is
ionized, and
the molar ratio of ionized pendant carboxyl groups to UV-absorbing pendant
groups is from
about 0.5 to about 3.5. In another embodiment of the ophthalmic lens at least
the major
portion of the polymer of the hydrogel is a derivative of methacrylic acid; in
one embodiment
at least the major portion 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
crosslinked hydrogel contains more than 30% by weight of liquid under
equilibrium
physiological conditions. In one embodiment the covalently crosslinked
hydrogel contains
less than 55% by weight of liquid under equilibrium physiological conditions.
In another
embodiment the covalently crosslinked hydrogel contains between 35% and 47.5%
by weight
of liquid under equilibrium physiological conditions. In one embodiment of the
ophthalmic
lens at least the posterior optical surface contributes refraction with
negative spherical
aberration. In another embodiment the posterior optical surface contributes
refraction with
negative spherical aberration between -0.1 microns and -2 microns. In a
further embodiment
the negative spherical aberration is between -0.5 microns and -1.5 microns.
Alternatively the
negative spherical aberration is between -0.75 microns and -1.25 microns. One
embodiment
of ophthalmic lens comprises both UV-absorbing pendant groups containing
carbonyl groups
conjugated to an aromatic system as well as UV-absorbing groups containing a
benzotriazole
structure. In one embodiment the UV-absorbing pendant groups containing
carbonyl groups
conjugated to an aromatic system and the UV-absorbing groups containing a
benzotriazole
structure are located in separate layers of the ophthalmic lens. In one
embodiment the
ophthalmic lens is implanted into a cornea. In another embodiment the lens is
implanted into
the anterior chamber of an eye between the cornea and iris. In yet another
embodiment the
lens is a phakic lens implanted between the iris and the natural crystalline
lens. In still
another embodiment the lens is implanted into the posterior chamber of eye, at
least partially
replacing the natural crystalline lens.
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Another aspect of the invention is directed to a method of adjusting the
refractive
properties of a fully hydrated hydrogel of the invention, where the method
comprises the step
of focused irradiation of the hydrogel with electromagnetic radiation such
that two-photon
absorption occurs, and the polymer component of the hydrogel undergoes partial
depolymerization and/or decomposition, with effective removal of part of the
polymer
scaffold to create a void.
An additional aspect of the invention is directed to a method of in situ
adjusting the
optical parameters of a hydrogel ophthalmic implant, the method comprising the
steps of: a)
providing an eye containing a hydrogel ophthalmic implant according to claim
13; and b)
irradiating a portion of the hydrogel ophthalmic implant with electromagnetic
radiation using
a femtosecond laser, whereby part of the copolymer of the hydrogel is
depolymerized and/or
ablated; where the optical parameters of the implant are adjusted. In one
embodiment of the
method the optical parameters include the refractive index. In one embodiment
of the method
the irradiation produces elongated cavities or voxels inside the hydrogel
ophthalmic implant.
In some embodiments the voxel depth is up to 20-30 microns or more. In one
embodiment
increasing the depth of the voxels increases the phase shift while the
refractive index remains
approximately constant; in one embodiment the refractive index is .3335. In
some
embodiments of the method the phase shift is up to 3 wavelengths of green
light. In other
embodiments of the method the depolymerized matter ablated by irradiation
comprises
soluble, easily diffusible compounds of low toxicity. Since the procedure
releases only very
low concentrations of depolymerized compounds into the intraocular space, the
low toxicity
of the method is also evident. In one embodiment of the method the modified
optical
properties are provided by forming in the hydrogel a pattern of elongated
voxels of varying
depth while keeping the modified refractive index approximately constant. In
one
embodiment of the method the phase shift is controlled by varying voxel depth
rather than by
varying their refractive index.
BRIEF DESCRIPTION OF THE DRAWINGS
The accompanying drawings, which are incorporated herein and constitute part
of this
specification, illustrate the presently preferred embodiments of the
invention, and, together
with the general description given above and the detailed description given
below, serve to
explain the features of the invention. In the drawings:
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Fig. 1 illustrates the internal arrangement of the eye with main structures
including the
cornea, sclera, iris, NCL, vitreous body, retina and the suspensory apparatus
of the lens
(capsule, zonules and cilliary muscle)
Fig. 2 illustrates distribution of refractive power in a lens with one
hyperbolic surface.
Fig. 3A is a cross-sectional view of a bioanalogic intraocular lens according
to an
exemplary embodiment of the 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 with a
circular
optical part and elliptical support part.
Fig. 4B is atop view of another exemplary embodiment of a lens with a circular
support part truncated by a single straight cut.
Fig. 4C is a top view of another exemplary embodiment of a lens with a
circular
support part truncated by two symmetric crescent cuts.
Fig. 4D is a top view of another exemplary embodiment of a lens with a
circular
support part truncated by one straight and two crescent cuts.
Fig. 4E is a top view of another exemplary embodiment of a lens with a
circular
support part truncated by four symmetric crescent cuts.
Fig. 4F is a top view of another exemplary embodiment of a lens with a
circular
support part truncated by two straight parallel cuts and the cylindrical lens
with cylinder axis
1B in the angle a with regard to the cuts direction.
Figs. 5A, 5B and 5C illustrate top views of exemplary lenses with the optical
surfaces
divided into two or more optical zones.
Figs. 6A, 6B and 6C are cross-sectional views of alterative lens in accordance
with
the invention composed from two or more materials.
Figs. 7A, 7B and 7C are expanded views illustrating alternative profiles of
the
supporting peripheral part of the exemplary lenses.
Fig. 8 illustrates the schematic arrangement of the mold for production of a
lens in
accordance with an exemplary embodiment of the invention.
Fig. 9A shows a comparison of the Raman spectra of a representative hydrogel
of the
invention before and after two photon absorption (TPA) using a femtosecond
laser.
Figs. 9B and 9C show graphs of relevant parameters of the Raman spectrum.
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DETAILED DESCRIPTION OF THE INVENTION
For the purposes of the present application, the terms "(meth)acrylic" and
"(meth)acrylate" denotes either acrylic/acrylate or methacrylic/methacrylate
moieties. In
some embodiments the polymers comprise acrylic/acrylate moieties. In other
embodiments
the polymers comprise methacrylic/methacrylate moieties. In still other
embodiments the
polymers comprise both acrylic/acrylate and methacrylic/methacrylate moieties.
In a
preferred embodiment the polymers comprise methacrylic/methacrylate moieties.
Also, for the purposes of the present application, the term "ablation" refers
to removal
of a section of the polymeric structural support of a hydrogel, preferably by
decomposition of
the polymer to diffusible, low-molecular fragments. The term
"depolymerization" constitutes
a special type of ablation where the fragments are monomers.
There are numerous types of implantable ophthalmic lenses used in various
locations
in the eye, from corneal stroma through anterior chamber to posterior chamber.
One of the
problems with implantable lenses is the complicated selection of the correct
refractive
properties (so called biometry) and difficulty to replace them or correct them
if the biometry
turned out to be wrong or if the optical requirements of eye change over time.
This inspires
the effort of the industry to develop implantable lenses whose optical
parameters could be
adjusted non-invasively and post-operatively. Change of the optical properties
of an artificial
lens, either pre-operatively or post-operatively, can be achieved by changing
refractive index
of the lens material.
The change of the refractive index by action of Two Photon Absorption (TPA) or
Multi-Photon Absorption (MPA) for implantable lenses is described in prior art
as positive
for hydrophilic polymers and negative for hydrophobic polymers. This is
consistent with
assumed mechanism of increased crosslinking density and consequent decrease of
water
content in hydrophilic acrylates and hydrogels, and conversely, with increased
hydrophilicity
in the hydrophobic acrylates. The assumed cause of these changes is a local
increase of
temperature.
Change of the optical properties of the natural (e.g., human cornea) or an
artificial
lens by changing refractive index of their material is also described in
numerous papers,
patents and patent applications, such as the following: Phillips, A.J., System
and Method for
Treatment of Hyperopia and Myopia, US Patent No. 6,102,906, Bille J.F.: System
For
Forming And Modifying Lenses And Lenses Formed Thereby, US Patent No.
8,292,952;
USP 8,920,690; 9,192,292; Sahler; Ruth et al.: "Hydrophilicity alteration
system and
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method" USP 9,023,257; 9,186,242 and 9,107,746; Sahler; Ruth et al.:
"intraocular lens
(TOL) fabrication system and method" US Patent Application Publication No.
20160074967;
Smith, T. et al.: Optical Hydrogel Material With Photosensitizer And Method
For Modifying
The Refractive Index, US Patent Application Publication Nos. 20130268072;
20090287306
and US Patent No. 8,901,190; Knox, Wayne H. et al.: Optical Material And
Method For
Modifying The Refractive Index, US Patents Nos. 8,932,352; 8,337,553;
7,789,910 B2;
Knox, Wayne H. et al.: Optical Material And Method For Modifying The
Refractive Index,
US Patent Applications Publication Nos. 20130138093; 20130178934; 20100298933;
20080001320; 20090143858; 20090143858; Knox, Wayne H. et al.: Method For
Modifying
The Refractive Index Of An Optical Material And Resulting Optical Vision
Component, US
Patent Application Publication No. 20120310340, Knox, Wayne H. et al.: Method
For
Modifying Refractive Index Of Ocular Tissues, US Patent Nos. 8,486,055;
8,512,320;
8,617,147 and US Patent Application Publication Nos. 20110071509; 20130226162
and
20140107632; Knox, Wayne H. et al.: Method For Modifying Refractive Index Of
Ocular
Tissues And Applications Thereof, US Patent Application Publication No.
20120310223,
each of which is incorporated herein by reference. None of these prior art
references suggests
an ablation or depolymerization as the mechanism of the optical adjustment in
hydrogels.
Applicant's co-pending international application, PCT/M2016/052487, Method and
Device for Optimizing Vision Via Customization of Spherical Aberration of Eye,
filed on
May 02, 2016, contains related disclosure.
The wavelength of the laser beam is usually in the range of near infrared
radiation,
about 800 nm to 1300 nm, or more typically in the range of visible and near
infrared radiation
from about 660 nm to about 1100 nm. The use of higher wavelengths is usually
preferred
because of safety concerns. One method of changing the refractive index of the
lens material,
by means of a femtosecond laser ("FSL procedure" for short), can, in
principle, achieve many
changes of refractive properties, such as spherical refractive power,
cylindrical refractive
power, spherical aberration, etc. In principle, the procedure can selectively
change any of the
coefficients in the Zernike polynomial ("Zernike Coefficients") repeatedly
even in the
implanted lens. This has been demonstrated by Gustavo A. Gandara-Montano et al
"Femtosecond laser writing of freeform gradient index microlenses in hydrogel-
based contact
lenses", 1 Oct 20151 Vol. 5, No. 10 DOI:10.1364/0ME.5.0022571OPTICAL
MATERIALS EXPRESS 2257. The authors noted a negative phase shift in the
ETAFILCON contact lens hydrogel when treated by femtosecond laser at 800 nm.
However, none of these systems is in clinical use because, among other
reasons, the change
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of the refractive index using an FSL procedure is rather small for the
currently available
ophthalmic lens materials.
The FSL procedure can be performed both on hydrophilic (where the refractive
index
RI is usually increased) and hydrophobic IOL materials (where the RI is
usually decreased).
The FSL procedure can be more advantageously performed on hydrogel ophthalmic
lenses, particularly on implantable lenses of various types, since the
refractive change even in
the currently available hydrogels is higher than in the hydrophobic materials.
In addition,
FSL procedure in hydrophobic acrylates could lead, at least in theory, to so-
called
"glistening" or other problems related to formation of hydrophilic "osmotic
cells" within a
hydrophobic material. Therefore, various hydrogels were tested as substrates
for
micromachining by femtosecond lasers. The sensitivity of hydrogels can be
increased by
various "dopants" designed to respond to various wavelengths of
electromagnetic radiation.
A number of dopants have been described so far in the patent and scientific
literature. All
presently known dopants are single compounds capable of UV absorption, but no
known
dopants are capable of single-photon absorption at the wavelength used for the
FSL
procedure.
We have found that the effect of dopants for TPA and MPA in the acrylic and
methacrylic hydrogels can be improved, modified and strengthened by adding
certain TPA
activators, e.g. co-monomers containing negatively charged pendant groups,
particularly
organic carboxylic acid salts. Both the dopant and its activator can be
advantageously
covalently bound to polymer chains, more advantageously to the chains forming
the polymer
network of the hydrogel. Both the dopant and its activator may be bound to the
same
polymer chain or to different polymer chains.
More particularly, the inventive hydrogels contain a combination dopant in the
form
of a minor part of a pendant UV-absorbing structure that does not
significantly absorb visible
light, with a minor part of the activator in the form of pendant groups
comprising an ionized
salt of an acid, with a major part of methacrylate neutral hydrophilic
derivative, particularly
glycol or glycerol esters of methacrylic acid and a minor part of a polyol
with at least two of
the hydroxyl groups esterified by acrylic or methacrylic acid. These hydrogels
are materials
particularly suitable for adjustment of refractive index by absorption of
femtosecond pulses
of visible or near infrared radiation (femtosecond laser (FSL) treatment). The
absorption of
electromagnetic energy causes controlled degradation and depolymerization of
the polymeric
component of such hydrogels thereby forming domains with lower refractive
index. Thanks
to the activator, such domains can be formed even if the hydrogel is
irradiated by a
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femtosecond laser with pulses of relatively low energy and at a very high
"writing speed" or
"scanning velocity".
Disclosed herein are bioanalogic covalently crosslinked acrylic and
methacrylic
hydrogels containing negatively charged groups, particularly carboxylate
groups, and
containing also monomers with pendant UV-absorbent groups, such as, for
example,
methacryloyloxybenzophenone (MOBP). Also disclosed herein is a method to
depolymerize
such polymers by absorption of electromagnetic radiation in order to adjust
optical properties
of the implantable hydrogel using femtosecond laser to depolymerize parts of
the hydrogel
and to form internal cavities that would form new refractive members, such as
toric lenses.
Implicitly but certainly, such newly formed cavities in the hydrogel matrix
would be
necessarily filled with water or aqueous fluid (either alone or also
containing the residual of
the degraded polymer component of the hydrogel) and would have, therefore,
different
(specifically lower) refractive index than the parent hydrogel material. Also
disclosed is the
possibility to perform this lens adjustment in situ on the implant since the
products of such a
decomposition are water-soluble compounds of low toxicity and capable of a
slow diffusion
through the hydrogel.
We have confirmed that the UV-absorbent group used in the hydrogel acts as a
"dopant" increasing the depolymerization rate, while the negatively charged
group acts as an
"activator" for the dopant, increasing the dopant's efficacy still further. In
addition, the
activator acting as a quencher protects the material from undesirable
"charring" or "burning"
that may happen over a certain amount of absorbed energy per time. This allows
one to
lower the overall amount of energy needed to achieve a certain refractive
change, and
therefore to use safely the more energetic visible light rather than less
energetic near infrared
(NIR) radiation.
Without wishing to limit the scope of the invention by any particular
hypothesis or
theory, the proposed explanation of the phenomena of the hydrogel
depolymerization by
electromagnetic radiation is as follows. Focused femtosecond lasers are known
to facilitate
two-photon-absorption (TPA) (or even Multi-Photon Absorption (MPA) in their
focal
volume. The volume of the material affected by the TPA is usually called
"volume pixel", or
"voxel". The size of the voxel depends on various parameters, and generally
increases with
the absorbed energy. The smallest possible voxel size corresponds to the
"focal volume", or
an ellipsoid with volume roughly the cube of the wavelength of the laser light
absorbed. The
diameter of a voxel is much smaller than its depth thanks to the self-focusing
of the laser
beam in the case of TPA or MPA, and is typically around 500 nm. The voxel
depth is much
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larger and grows with deposited energy. The reported maximum voxel depth is
about 6
microns in currently known materials.
If the radiation of a certain wavelength is absorbed via TPA or MPA, then the
dopant
accumulates excitation energy equivalent to a single-photon absorption of the
wavelength of
incident light divided by the number of the absorbed photons. For example,
laser light at 400
nm absorbed via TPA by the dopant will correspond to 200 nm light absorbed via
SPA. Light
of 200 nm wavelength is a hard UV light (UVC Band) that is sufficiently
energetic to cause a
breakage of chemical bonds and a deep structural rearrangement. Of course, 3
photon
absorption or 4 photon absorption, although much less probable than TPA, would
accumulate
even higher energy concentration. The excited state with high accumulated
energy is short-
lived since the energy dissipates quickly via one of several possible
pathways, the most usual
being conversion into thermal energy. The usual effect of femtosecond laser
treatment of
tissues or synthetic hydrogels involves local increase of temperature that may
cause ¨
depending on the amount of the energy absorbed ¨ effects starting with
conversion of the
matter into a plasma (e.g., via laser ablation) through heating sufficient for
charring, to more
subtle additional crosslinking via various mechanisms such as re-
esterification,
disproportionation or dehydration with formation of ether links.
However, the presence of the activator (pendant negatively charged groups,
particularly carboxylate groups), changes the mechanism of this process. We
believe that this
is caused by certain cooperation of the activator groups with dopants. In this
cooperation,
activation of dopants further facilitates TPA absorption and increases its
effect. The activator,
such as the carboxylic acid group of methacrylic acid or a salt thereof, is
apparently in an
interaction with the dopant because its presence causes a change of the
dopant's UV/Vis
spectrum (namely, a subtle shift from the UV toward the visible region). This
may increase
.. the dopant's "TPA cross-section" and increase the absorption efficacy of
the dopant.
However, the main role of the activator may be in channeling the absorbed
energy of
the two (or more for MPA) absorbed photons from the dopant to the main polymer
chain to
cause breakage of the covalent bond in the main chain. It is understood from
the general
character of this process, and from the known depolymerization kinetics of
methacrylate
polymers, that this cleavage is homopolar and produces free radicals that
start the
depolymerization process in the vicinity of the negatively charged activator
(e.g.,
carboxylate). This free radical mechanism may vastly increase the quantum
yield of the
photodegradation and helps to explain why even a relatively low concentration
(on molar
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basis) of the dopant can cause profound changes in the local composition and
structure, and
consequently a substantial change in the local refractive index.
Furthermore, the activator groups help to divert the absorbed energy from the
dopant
to quench the excited state and "consume" this energy for the breakage of
covalent bonds.
This "energy conduit" function of the activator prevents excessive heat
accumulation in the
vicinity of the dopant moiety, helping thus to "recycle" the dopant molecule
and preserve it
for the future TPA cycle. It also helps to protect the polymer structure from
burning and
charring and helps to increase the efficacy of the whole absorption process.
In addition, this proposed mechanism can explain the fact that in the absence
of the
activator, TPA in hydrogels with only the dopant proceeds via a different
mechanism and
achieves the opposite result: while in presence of the activator the
refractive index decreases,
in its absence the refractive index increases. One possible reason for the
increased refractive
index seems to be an additional crosslinking leading to decreased water
content and,
consequently, increased refractive index since water has the lowest refractive
index of all
hydrogel constituents.
Another consequence of the stipulated mechanism involving the activator is the
reduced (or even eliminated) change of the volume due the TPA process. Namely,
in the case
of the additional crosslinking the water content is reduced while the amount
of the polymer
components stays approximately the same. Therefore, the mass and the volume of
the treated
hydrogel volume has to be reduced, with various adverse consequences on the
product
(generation of internal stress, change of geometry and change of mechanical
properties, for
instance). Conversely, if the polymer becomes more hydrophilic and attracts
more water, the
treated volume of the polymer has to expand (again, with adverse effect on the
geometry,
optics and stresses within the material).
The mechanism stipulated in the present invention involves exchange of a
certain part
of the polymer mass for water, with a low net volume change, if any. The
depolymerization
of the hydrogels covered by the present invention yields low toxicity monomers
and/or their
fragments, primarily 2-hydroxyethyl methacrylate, methacrylic acid and
ethylene glycol. All
these decomposition products are well soluble in water and capable of
diffusing through the
surrounding intact hydrogel network, to be dissipated from the implant over
some time in
very low concentrations.
The proposed mechanism may explain very large phase shift achievable in
hydrogels
according to this invention. The phase shift is determined by the change in
the refractive
index and the length of light path through material with changed refractive
index, in other
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words, the voxel depth. Refractive index in hydrogels cannot be practically
lower than
refractive index of water, or about 1.3335 since the lowest conceivable index
can be achieved
in a cavity filled with aqueous liquid. The cavity in a hydrophilic polymer
matrix cannot be
filled with a gas for certain basic thermodynamic reasons, and there is no
known or readily
conceivable mechanism of creating a hydrophobic cavity within the hydrogel.
Therefore, large phase shifts (more than one wavelength) will require
formation of
voxels with large depth. It is usually assumed that the voxel depths can reach
at most about 5
to 6 microns. However, if the refractive index in the focal volume decreases,
then the
elongated voxel would act as a light-guide that channels additional light
pulses to be absorbed
in the voxel bottom. Consequently, the depth of the voxel can gradually
increase with
increasing number of pulses, and so does then the phase shift even though the
refractive index
remains approximately constant and equal to or higher than 1.3335. This
mechanism is then
different from mechanisms described to date where voxel size remains
approximately
constant but refractive index change increases with increased absorbed energy
(achieved e.g.,
number of pulses). It is believed that the presently proposed mechanism may
provide a voxel
depth larger than 10 microns, and phase shifts corresponding to voxel depth as
high as 25 or
30 microns have been observed. This is mentioned not to set a limit on the
invention, but
demonstrate the fundamental difference of the invention from prior art. One of
the practical
consequences of this difference is the following: materials used in the prior
art allow for
creation of refractive or diffractive structures by forming a refractive index
gradient (GRIN)
while hydrogels according to the present invention allow achieving similar
refractive or
diffractive effects by forming a pattern of varying voxel depth while keeping
the modified
refractive index approximately constant. In addition, it is believed that this
refractive index
value first approaches and ultimately approximates the refractive index of
isotonic saline
solution.
As a consequence of this novel mechanism, the newly created refractive or
diffractive
structures form a system of parallel waveguides (i.e., elongated voxels with
refractive index
lower than the surrounding hydrogel) of different length, whereby the phase
shift is
controlled by the varying voxel depth rather than their varying refractive
index.
An additional feature of the invention is creation of a gradient of refractive
index in
the vicinity of individual voxels. The monomers and other fragments released
by the
depolymerization inside the voxel migrate radially by diffusion through the
surrounding
hydrogel. When the temperature decreases below the ceiling temperature (around
200 C in
the case of methacrylates), at least part of the released monomers and/or
fragments may re-
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polymerize and create a denser network structure with a higher refractive
index than the
parent hydrogel. This mechanism has two beneficial consequences: first, it
reduces the
amount of compounds that diffuse outside of the implant to be metabolized, and
second, the
gradient of refractive index thus formed improves the light-guiding properties
of the voxel.
Thus, monomers released by depolymerization may not all diffuse out of the
hydrogel, but may partly re-polymerize in the voxel vicinity. In that sense
the present method
is somewhat similar to FS laser ablation in that that both remove some polymer
mass and
retain water instead, although the term "ablation" frequently denotes a
process of
decomposition where the polymer and water are converted into a plasma (gas).
The presently
.. disclosed process is much more gentle so that polymer decomposition does
not create an
exploding bubble of gas.
Although depolymerization of the hydrogel copolymer is believed to be the
primary
mechanism involved in the presently disclosed process, it may be supplemented
by other
decomposition reactions that form small, water-soluble fragments, such as
hydrolysis of the
(meth)acrylate pendant groups, or oxidation.
Dopant concentration in hydrogels according to the invention varies from about
0.05
%-mol to about 5%-mol, advantageously from about 0.1 %-mol to about 2.5 %-mol.
Alternatively the different optimum dopant concentration can be used for
various dopants,
e,g, 0.25 to 0.55 molar % for benzophenone derivatives, or 0.1 to 0.2 molar %
for
benzotriazole derivatives. The preferred dopants are UV absorbers with low
absorbance for
visible light, i.e. above a wavelength of about 390 nm. Examples of suitable
dopants are
vinyl, acrylate or methacrylate derivatives of benzophenone, benzotriazole and
coumarin,
although those skilled in the art can certainly identify other suitable UV
absorbers that work
as dopants in the sense of the invention.
The activating groups are present in concentration from about 0.25 %-mol to
about 5
%-mol, advantageously from about 0.75 to about 3.5 %-mol. In some embodiments
the
preferred molar concentration of methacrylic acid is between 1 and 1.25 %
molar. The
preferred activator groups are derivatives of acrylic or methacrylic acid
containing a pendant
acidic group including carboxylate group, sulphate group, sulfonate group and
phosphate
group. Such acidic groups are preferably neutralized by suitable organic or
inorganic cations.
The activator to dopant group molar ratio should be about 0.75 to 10,
advantageously from
about 1 to about 5 . Alternatively the different optimum molar ratio
activator/dopant are
different for different dopant. For instance, the optimum ratio for
benzophenone derivatives is
from about 2 and 4 and for benzotriazol derivatives between about 5 and about
7.
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The preferred composition of a hydrogel according to the invention comprises a
major
fraction of a methacrylate monomer with pendant neutral hydrophilic groups. By
"major
fraction" is meant at least 50%-mol of all monomer units in the hydrogel. In
some
embodiments the molar fraction of the hydrophilic methacrylate monomer is
between 90%-
mol and 99.5%-mol, and in many cases between 97.5%-mol and 99%-mol. Such
hydrophilic
methacrylate monomers can be esters of polyol aliphatic compounds, such as
glycols, glycol
ethers, glycerol and sugars. The most familiar in this group is 2-
hydroxyethylmethacrylate
(2-HEMA). Alternative monomers to the above mentioned esters comprise amides
of
methacrylic acid, such as methacrylamide, N-isopropyl methacrylamide or N-(2-
hydroxyethyl) methacrylamide. This major fraction of the hydrogel polymer may
also
comprise a mixture of such hydrophilic monomers. Alternatively, a minor
portion of the
methacrylate monomers (but not more than 25%-mol) can be replaced by analogous
acrylic
acid derivatives. In some embodiments 0.5 %-mol to 5%-mol can be replaced by
acrylate
monomers.
A minor part of the monomer units is formed by the above mentioned activator
monomers with a pendant negatively charged group, and still another minor part
is formed by
the above mentioned dopant monomers.
The polymer is advantageously covalently crosslinked. The crosslinking can be
achieved by any of the methods known to those skilled in the art, such as
radiation
crosslinking, crosslinking by formation of ether links between pendant OH
groups, etc.The
preferred crosslinking method is copolymerization with a minor fraction of
methacrylate or
acrylate crosslinking diesters or triesters of polyols, such as, for example,
triethylenglycol
dimethacrylate.
Refractive index of the hydrogel according to the invention is between about
1.38 and
about 1.48, preferably between about 1.40 and about 1.45. In some embodiments
the 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.
The preferred hydrogel according to the invention contains between about 25%-
wt
and about 85%-wt liquid in equilibrium with live intraocular environment. For
intraocular
lenses supplementing or replacing the natural crystalline lens, the more
advantageous
equilibrium liquid concentration is between 35 %-wt and 50 %-wt, and more
particularly
between 40% and 47%-wt. In some embodiments the equilibrium water content is
41%-wt
0.75%, or 42.5 1%-wt, or 44.5 1% by weight. It is understood by those
skilled in the art
that equilibrium liquid content in hydrogels is subject to many variables,
such as body
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temperature, composition of body fluids or pressure exerted by surrounding
tissues or body
structures. Also, measurement of liquid content in small hydrogel implants may
be burdened
by a certain measurement error, so that these values are illustrative. Since
the negatively
charged activator groups tend to increase equilibrium liquid content and thus
to decrease the
refractive index of the hydrogel, the higher concentration of the activator
may be
compensated with addition of hydrophobic acrylates or methacrylates, such as
methylmethacrylate, ethylmethacrylate, benzylmethacrylate,
isobornylmethacrylate or
ethoxyethyl methacrylate that will reduce the water content and increase
refractive index into
a desirable range for the ophthalmic implant design. Typical concentration of
such added
methacrylates or acrylates is up to about 40% molar. In some embodiments the
concentration
of added hydrophobic monomers is between about 5%-molar and 25%-molar.
The dopant group and activating group may be located on the same polymer
chain, or
on different polymer chains that are in an intimate contact e.g. in a polymer
blend, or being
on two different segments of polymer network. One or other may be on a graft,
or one can be
on a graft and the other on the basic chain. The dopant group and activating
group may be on
a single molecule in a suitable steric relationship allowing their mutual
interaction. Such a
"dopant-activator complex" may be then covalently bound to a polymer chain, or
admixed
into the polymer, or become part of the main polymer chain.
Negatively charged activator groups have some additional advantages for
intraocular
implants, such as improved biocompatibility, resistance to adsorption of
proteins, resistance
to formation of biofilms, resistance to calcification, and resistance to
adhesion and spreading
of cells (which translates into resistance to sclerotization of posterior
capsule and resistance
to posterior capsule opacification).
The hydrogel implant according to the invention can be placed in various
locations
within the eye along the optical path. It may have the form of a "blank" in
which a refractive
or diffractive lens is created, or it may have the form of a refractive or
diffractive lens which
optical properties are modified by the refractive index change within selected
locations of the
implant.
Intraocular lenses may replace partly or fully the natural crystalline lens.
Intraocular
hydrogel lenses have some advantages over other IOL types, and various types
and related
manufacturing methods are described, e. .g., in the following patents and
patent applications:
Stoy, V. et al.: Bioanalogic Intraocular Lens, International Patent
Application
W02014111769; Wichterle, 0.: Method Of Molding An Intraocular Lens, US Patent
No.
4846832; Wichterle, 0.: Soft And Elastic Intracameral Lens And Method For
Manufacturing
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Thereof, US Patent No. 4,846,832; Stoy, V.: Implantable Ophthalmic Lens, A
Method Of
Manufacturing Same And A Mold For Carrying Out Said Method, US Patent No.
5,674,283;
Sulc, J., et al.: Soft Intracameral Lens, US Patent Nos. 4,994,083 and US
Patent 4,955,903;
and Michalek, J. et al.: Method Of Manufacturing An Implantable Intraocular
Planar/Convex,
Biconvex, Planar/Concave Or Convex/Concave Lens And A Lens Made Using This
Method,
US Patent No. 8409481; each of which is incorporated herein by reference.
Another type of implantable ophthalmic lens is so called "Implantable Contact
Lens"
(or ICL). ICL is a phakic lens placed between iris and the natural crystalline
lens. It is
described in several patents, e.g., Fedorov, , et al. Intraocular lens for
correcting moderate to
severe hypermetropia, US 5,766,245; Feingold V., Intraocular contact lens and
method of
implantation, UP 5,913,898; Intraocular refractive correction lens, US
6,106,553; each of
which is incorporated herein by reference.
These implantable lenses are made from hydrogels that incorporate a biological
component, usually collagen. These so called "Collamers" are described in
various patents,
such as Feingold, et al., Biocompatible, optically transparent, ultraviolet
light absorbing,
polymeric material based upon collagen and method of making, US 5,910,537;
Feingold, et
al. Biocompatible optically transparent polymeric material based upon collagen
and method
of making, US 5,654,349, US 5,654,388 and US 5,661,218; Fedorov, et al.,
Biocompatible
polymeric materials, methods of preparing such materials and uses thereof, US
5,993,796;
Method of preparing a biological material for use in ophthalmology; each of
which is
incorporated herein by reference.
Another approach to refractive correction is intrastromal or intracorneal
implants or
inlays, described e.g. by Miller in Aspherical corneal implant, US 7,776,086;
Lang, Alan in
Design of Inlays With Intrinsic Diopter Power, US 2007/0255401; Dishler; Jon
et al. in Small
Diameter Inlays, US 2007/0203577; Lang, Alan et al. in Intracorneal Inlays, US
2007/0129797; and Dishler, et al. in Method of using small diameter
intracorneal inlays to
treat visual impairment, US 8,057,541; each of which is incorporated herein by
reference.
Some intrastromal implants are designed for post-operative adjustment of
optical
power by laser, as described by Peyman, Gholam A. in Intrastromal corneal
modification via
laser, US 2001/0027314; in Adjustable ablatable inlay, US 2002/0138069 and
2002/0138070;
Ablatable intracorneal inlay with predetermined refractive properties in US
2003/0093066;
and in Bifocal implant and method for altering the refractive properties of
the eye, US
2005/0222679; each of which is incorporated herein by reference.
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There are also intraocular implants using two lenses in tandem in order to
achieve
improved accommodation capability, modularity of the design or reduced size of
implantation incision when the implant is in situ assembled from individual
parts.
All such ophthalmic lens types can be manufactured from the hydrogels
according to
the present disclosure and then modified by using a femtosecond laser
treatment.
Again, without wishing to be limited by any particular theory, the partial
depolymerization of the presently disclosed method appears to be the mechanism
of the
profound negative phase shift observed in the inventive hydrogels when exposed
to focused
femtosecond lasers (FSLs). As noted above, this process appears to generate a
system of
parallel longitudinal "voxels" that get more elongated with increasing number
of absorbed
FSP pulses. The voxel length controls the phase shift achievable for a given
refractive index,
that cannot decrease below the RI value for water, which is 1.3335. Further,
it is actually the
phase shift, not the refractive index per se, that is affected in the
presently disclosed
hydrogels and methods. This mechanism is different from anything disclosed by
the prior art
(presumed crosslinking leading to decrease in the water content that leads to
the refractive
index increase and, therefore, a corresponding positive phase shift).
A further difference from the prior art is that the presently disclosed
hydrogels and
methods replace part of the hydrogel-forming copolymer with water or aqueous
fluid, rather
than decreasing the water content of the hydrogel by modifying the copolymer
properties
(such as by polymer crosslinking).
In the drawings, like numerals indicate like elements throughout. Certain
terminology
is used herein for convenience only and is not to be taken as a limitation on
the present
invention. The following describes preferred embodiments of the present
invention.
However, it should be understood, based on this disclosure, that the invention
is not limited
by the preferred embodiments described herein.
The NCL has a very complicated structure that develops over time. One of the
structural features is asphericity of posterior and anterior surfaces of the
NCL 103. As
established in recent years E.L.Markwell et al, MRI study of the change in
crystalline lens
shape with accommodation and aging in humans, Journal of Vision (20110
11(3);19, 1-16;
M.Dubbelman et al, Change in shape of the aging human crystalline lens with
accommodation, Vision Research 45 (2005), 117-132;F. Manns et al, Radius of
curvature and
asphericity of the anterior and posterior surface of human cadaver crystalline
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
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(2001) 1867-1877, both anterior and posterior surfaces of a young human lens
are hyperbolic
and can be characterized by the equation:
Y - Yo = XA2/1 Ro*(1+1-h*(X!Ro)^2)^0.51 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 apex position on the main
optical axis 1A,
Ro is the central radius of curvature and h is the conic constant (or the
shape parameter). The
Eq. 1 describes any conic section curve depending on the shape parameter h
value: it is a
parabola for h= 0, a circle for h = 1, hyperbole for h<0, prolate ellipse for
0<h<1 and oblate
ellipse for h > 1.
It has been found that for a typical young human NCL, the anterior surface is
more
hyperbolic than the posterior surface, that hyperbolicity increases
significantly with
accommodation, and that the human lens grows with age and its hyperbolicity
decreases so
that an old NCL may become approximately spherical.
The referenced studies mapped dimensions of typical NCL for selected
population
samples. According to these references, a typical human lens anterior central
radius ranges
from about 5 to 13 mm and the average anterior conic parameter is about -4
(ranging from
about -22 to +6). The posterior central radius ranges from about 4 to 8 mm and
the average
posterior conic parameter is about -3 (ranging from about -14 to +3).
The central thickness of a young, relaxed NCL ranges typically from about 3.2
mm to
about 4.2 mm, increasing with age and/or with the near-focus adjustment to a
thickness from
about 3.5 mm to about 5.4 mm. The posterior part depth of the NCL is typically
the same as,
or larger than the anterior part depth. Therefore, the sagittal depth of the
posterior lens
surface is typically from about 1.75 mm to about 2.75 mm on equatorial
diameter from about
8.4 mm to about 10 mm. This defines the basic dimensions of the posterior
capsule in its
"natural" state.
Although the above references do not state any particular connection between
the
geometry and optical properties of the NCL, we have found by mathematical
modeling that
the hyperbolic surfaces turn a lens polyfocal, with the refractive power
maximum at its center
and gradually decreasing toward the periphery. One direct consequence expected
from such a
polyfocality is a large focal depth of the lens so that a near object can be
projected on the
retina even without any particular lens shape change. Another implication of
the modeling is
that the average refractive power of the lens increases with decreasing
aperture. Therefore, it
is concluded that the near focus can be improved by pupil constriction (this
so called
"pupillary reflex" or "near myosis" that can be actually clinically observed
at near focus).
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Another consequence of the natural lens hyperbolicity is the capability of the
human (and
particularly young) brain to naturally neuro-adapt to, and correctly interpret
images formed
by projection through a hyperbolic lens onto retina.
This accommodative mechanism utilizing certain type of the polyfocality
deserves
further explanation as follows.
Lenses with at least one hyperbolic surface demonstrate a "hyperbolic
aberration" that
is opposite of the spherical aberration: rays incoming through the center are
bent into a focal
point that is closest to the lens, and the focal point becomes progressively
further from the
lens for rays incoming in increasing distance from the lens center toward the
lens periphery.
Therefore, the lens with hyperbolic surface is positively polyfocal: it has
the shortest
focal distance (i.e., highest refractive power) at its center, and the focal
distance increases
(i.e., refractive power decreases) from the center toward the lens periphery.
The focal range
of a hyperbolic lens can be rather large and is controllable by so called
conic constant or
shape parameter defining the hyperbolic surface shape.
Examples of the distribution of refractive power in a lens with hyperbolic
surface is
shown in Fig. 2 where local refractive power in Diopters are plotted
against the distance
from the optical axis in mm. It is understood that such an optical profile
with refractive power
decreasing with the distance from the optical axis, or with the aperture of
the imaging system,
or pupil diameter of the eye, can be present in the originally implanted lens,
or it may be
created by the hydrogel modification by a laser after lens implantation.
Based on the present studies it is understood that the positive polyfocality
and its
changes in the natural lens assist the eye to accommodate in several ways:
It projects on the retina simultaneously images of all objects in the field of
view in all
distances covered by the focal range of the lens. This significantly (by more
than 1 Diopter)
increases the depth of the focus of the eye since all objects create a well-
focused image
(accompanied by many dis-focused images that the brain learns to suppress).
The natural lens increases its hyperbolicity due to the accommodation, which
further
increases the focal range of the lens and, therefore, the depth of the focus
still further.
The eye helps to focus on near objects by narrowing the pupil. This so called
"pupillary reflex" or "near myosis" has two consequences: first, it decreases
the aperture and
thus increases depth of the focus of the eye as the optical system (narrowing
aperture blocks
rays that are far from the axis and coming in at sharp angles with respect to
the axis); and it
increases the average refractive power of the lens by using only its central
portion with the
highest refractive power.
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It is obvious from our studies that near myosis can assist the near focus only
for lenses
with hyperbolic aberration, i.e. with positive polyfocality. It has little
effect in monofocal
parabolic lenses, and it is counterproductive in lenses with negative
polyfocality: spherical or
elliptical (e.g., meniscoid) lens becomes weaker lens with lower refractive
power by the near
myosis rather than a stronger lens that is needed for near focus.
An artificial lens according to the invention is a hydrogel device implantable
into the
posterior chamber of human eye for replacement of the natural crystalline
lens. It is designed
to mimic or replicate essential physiological and optical functions of natural
lens without
creating problems that earlier attempts could cause in some situations. It is
important to
recognize that this is achieved by a novel thoughtful combination of features
that might have
been individually, or in different combinations, applied previously with a
lesser success. The
natural lens also achieves its function due to its balanced combination of
features rather than
to a single feature.
The features contributing to the overall function and combined according to
the
invention include size and shape of the implant; material properties; surface
properties;
optical properties; implantation method; and manufacturing method. We will
describe the
various features below and provide exemplary configurations of how individual
features
mutually interact to provide beneficial effect. It is important to recognize
that the implant
may combine several of the described features to achieve desirable effects,
however, the
invention is not limited to the exemplary configurations described below and
includes various
combinations of features.
Referring to Figs. 3A and 3B, the implant has a main optical axis 1A with a
central
optical part 2 and a peripheral supporting part 3. The overall shape of the
implant is defined
by its anterior surface 4, posterior surface 5 and the transition surface 6
between the upper
boundaries 7A and 7B of the anterior and posterior faces, respectively. Each
face is
composed of two or more surfaces. The anterior central optical surface 8A has
boundary 9A
and central posterior optical surface 8B has boundary 9B. Each of the surfaces
may be
divided into two or more zones with the boundary between them (denoted 13A and
13B in
Figs. 5A to 5C) being circles, straight lines or otherwise defined shapes. The
apexes of the
central anterior optical surface 10A and central posterior optical surface 10B
are positioned
on the main optical axis 1A. The anterior peripheral supporting surface is 11A
and the
posterior peripheral supporting surface is 11B.
Referring to the Figures 6A, 6B and 6C describing lenses comprising several
different
materials or layers, it is understood that any layer or structure in the
optical path may be
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formed by the hydrogel according to the present invention. The preferred layer
may be the
layer closest to the cornea, i.e. the one forming the anterior optical surface
of the lens. Optical
structures formed by the hydrogel optical modification may be refractive
structures,
diffractive structures, Fresnel lenses shown in the Fig. 6B, refractive index
gradient lenses or
similarly. Such structures can be formed within the hydrogel, or on its
surface, preferably on
the anterior optical surface.
The boundaries 7A and 7B are distinguishable as a discontinuity on the top of
the
anterior and posterior surfaces 4 and 5, respectively. Such a discontinuity
lay in the inflexion
point of the surface in the direction of the optical axis, or a in a point of
discontinuity of the
second derivative of the surface in the direction of the optical axis. The
boundary can be
rounded and continuous, but advantageously it is formed by a sharp rim or
edge. The
advantage of the sharp edge is in forming the obstacle to migration of cells
such as fibroblasts
along the capsule surface (the usual reason for posterior capsule
opacification).
The overall lens diameter is defined as the larger diameter of the boundaries
7A and
7B. The lens optical zone diameter is defined as the smallest diameter of the
boundaries 9A
and 9B. The posterior sagittal depth is the vertical distance between the
posterior apex 10B
and the plane defining the posterior boundary 7B. Central thickness is the
distance between
apexes 10A and 10B. Anterior depth is the vertical distance between the
anterior apex 10A
and the plane defining the anterior boundary 7A.
The main optical axis 1A may be the axis of symmetry in the case that
boundaries 7A
and 7B, as well as boundaries 9A and 9B, are defined by circles in the plane
perpendicular to
the optical axis, and if the central optical part 2 is symmetrical and e.g.,
does not have any
cylindrical component. Such implant with symmetric circular footprint is shown
in Fig. 3B.
However, the rims and boundaries may have other than circular footprint, e.g.
elliptical as
shown in Fig. 4A, or may have the footprint shaped as a truncated circle in
Figs. 4B to 4E
with single, double, triple or quadruple truncating cuts 12A to 12D. These
truncated footprint
shapes serve several purposes:
They provide better access into the space behind the lens during the
implantation. It is
important to clean this space well in order to remove any viscoelastic
polymers or lubricants
or other auxiliary agents before the surgical incision is closed.
They prevent rotation of the lens after the capsule shrinks around the IOL.
This is
particularly important for toric lenses.
They facilitate folding and insertion through a small incision.
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In the case that the optics has a cylindrical component, then the cylinder
axis 1B will
be positioned in a defined way with respect to the asymmetry of the outside
rim, e.g. be in the
angle a to the truncating cuts 12A and 12B as shown in the Fig. 4F. Needless
to say that the
truncating cuts 12A to 12D may not be necessarily straight cuts, but may be
suitably formed
to e.g. a crescent shape, and their number may be even higher than 4. Also,
the truncating
cuts may not be of the same length or positioned symmetrically. It can be
appreciated that the
footprint with truncated rim will facilitate folding of the implant and its
insertion through a
small surgical incision. In addition, the asymmetric rim footprint will
prevent the implant
rotation once the capsule settles around it. This is particularly important
for toric lenses with
a cylindrical component designed to compensate for astigmatism.
The posterior surface 5 is shaped and sized to approximate the shape and size
of the
posterior surface of the natural lens and to achieve contact with at least the
major part of the
posterior capsule of the eye. This is important for several reasons:
The implant will keep the posterior capsule in its natural shape, unwrinkled
and
smooth for the optimum optical performance;
The tight contact between the capsule and the implant will prevent migration
of
fibroblasts that could cause the posterior capsule opacification; this is
particularly effective if
the posterior surface is highly hydrated and carrying fixed negative charge.
The implant will occupy the space vacated by the posterior side of the natural
lens and
keep thus vitreous body from advancing forward and prevent thus the decrease
of the
pressure of vitreous body against retina (which could facilitate retinal
detachment and/or
cystoic macular edema).
It should be noted that the intimate contact between the implant and posterior
capsule
is beneficial particularly if the contacting surface of the implant is
hydrophilic and carrying
fixed negative charge in order to prevent capsular fibrosis and its consequent
stiffening,
opacification and contraction that would interfere with the implant function
(or could even
dislocate it), as will be described hereinafter.
In the preferred embodiment of the invention, at least the major part of the
posterior
surface 5 is formed by a generally smooth convex surface formed by rotation of
conic
sections around the optical axis, or a combination of such surfaces. The
peripheral part is
preferably formed by a conic surface or a hyperboloid surface, while the
central optical
surface is preferably hyperboloid, paraboloid or spherical surface (or a
combination thereof).
The sagittal depth of the posterior surface (i.e. the vertical distance
between the posterior
central optical surface apex 10B and the boundary of the posterior surface 7B,
measured on
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the main optical axis 1A) should be larger than 1.1 mm in order for lens to
perform its
function well. To perform well in the whole refractive range, the posterior
sagittal depth
should be larger than 1.25 mm, advantageously larger than 1.75 mm and
preferably larger
than 2 mm, but in any case less than about 2.75 mm.
The overall outer diameter of the implant (LOD) is important for its
centricity,
position stability and capsule-filling capability. The outer diameter of the
posterior surface 5,
i.e. the largest dimension of the posterior outer boundary 7B (in the plane
perpendicular to
the main axis 1A) should be larger than 8.4 mm, desirably at least 8.9 mm and
preferably at
least 9.2 mm. The largest outer diameter permissible is about 11 mm, but
desirably should be
lower than 10.75 mm and preferably at smaller than 10.5 mm. The considerable
flexibility in
the outer dimensions is allowed by several factors ¨ flexibility of the lens,
and particularly
flexibility of the outer peripheral supporting part 3 that can accommodate
various capsule
sizes and capsule contraction without deforming the central optical part 2.
The central optical surfaces may consist of one or more zones with different
geometry. The zones may be concentric, in which case the posterior boundary
13B between
them in the Fig. 5A will be circular. Zones may also be divided by straight
boundaries, in
which case the zones may have crescent or wedge footprint. Various examples
are shown in
Figs. 5A to 5C. The zones may be on the anterior or posterior optical surface.
Fig. 5A shows
the posterior optical surface is divided by the boundary 13B into two
concentric optical zones
¨ the central optical zone 8B1 and the outer optical zone 8B2. For instance,
the posterior
optical surface of the central optical zone 8B1 may be a spherical or
parabolic zone used for
the sharp near vision, while the hyperbolic outer zone serves for intermediate
and far distance
vision. Alternatively, both zones may have hyperboloid surfaces with different
central radii
Ro and/or different conic constants. Each optical surface may be also divided
into more than
two zones. The example in Fig. 5B shows the top view of the lens which
anterior optical
surface 8A is divided by a straight boundary 13A into two optical zones of
equal area 8A1
and 8A2. Each of those zones has different shape with different optical
parameters. The
example in Fig. 5C shows a top view of a lens with anterior optical surface 8A
divided by
two straight boundaries 13A and 13B into four paired optical zones 8A1 and
8A2, each
having a different area and different optical parameters. For instance, 8A1
may have higher
refractive power that 8A2 and serve for near focus. One of the zones may have
a cylindrical
component.
Both optical surfaces (or their zones or segments) are surfaces formed by
rotation of a
conical section along the optical axis, or by a combination thereof One or
both optical
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surfaces may contain one or more spherical optical zones. Advantageously, at
least one of
the optical surfaces comprises at least one hyperbolic surface, preferably in
the outer optical
zone. Preferably, both optical surfaces comprise at least one hyperbolic zone
each. Such
hyperbolic surface resembles the surfaces of the NCL and mimics some of its
beneficial
optical properties. Even more preferably, both posterior and anterior optical
surfaces are
hyperbolic surfaces or a combination of two or more concentric hyperbolic
zones. Lenses
with at least one hyperbolic surface have so called hyperbolic aberration, the
very opposite of
spherical aberration of lenses with spherical, ellipsoid or meniscoid
surfaces. The lenses with
hyperbolic aberration have highest refraction in the center and gradually
decreasing with
distance from the optical axis. (In lenses with spherical aberration the
refractive power
increases with distance from the optical axis.) The hyperbolic aberration
helps the eye to
accommodate through several mechanisms described above. It should be
understood that
hyperbolic aberration can be created not only by the exactly hyperbolic
surfaces in the
geometry sense, but also by similar surfaces where surface steepness generally
decreases with
the distance from the optical axis. Therefore, by "hyperbolic surfaces" are
meant also other
hyperbole-like surfaces approximating this property.
As the spherical aberration is measured with increasing aperture (or pupil
diameter)
the negative spherical aberration of hyperboloid-like surfaces increases in
absolute value (i.e.,
gets more negative). The spherical aberration can be expressed in various
alternative ways,
such as a deviation of wave-front in microns, or as steepness of decrease of
local refractive
power from the optical axis, or by decrease of the refractive power with
increasing aperture,
or by the value of conic constant or shape parameter corresponding to such
optical profiles.
Those skilled in the art can readily convert one of such values into the
corresponding value
expressed in a different way. The implant according to the present invention
can have
originally any spherical aberration, since its value can be adjusted post-
operatively using the
method of this invention.
The spherical aberration in the final implanted state (i.e., after the
hydrogel
modification by a laser) will be generally between about -0.1 microns and -2
microns on
aperture 4.5 mm. Preferably, the final spherical aberration will be between
about -0.5
microns and -1.5 microns on aperture 4.5 mm, and even more advantageously the
spherical
aberration will be between about -0.75 microns and -1.25 microns on the
aperture 4.5 mm.
In order to mimic the optical properties of the NCL, conical constants of the
anterior
and posterior optical surfaces are selected so that the refractive power of
the central optical
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part 2 generally decreases from the highest value at the optical axis to the
lowest value at the
periphery of the central optical part 2.
The steepness of the refractive power decrease with the distance from optical
axis is
dependent on the shape parameter (conic constant) of the hyperbolic surface.
The conic
parameter should be selected that the average decrease of the refractive power
is between -
0.25 Dpt/mm and -3 Dpt/mm, advantageously between -0.5 Dpt/mm and -2.5 Dpt/mm
and
preferably between about -1 Dpt/mm and -2 Dpt/mm.
The posterior central radius of curvature (at the point where the optical axis
intersects
the posterior apex) is advantageously from 2.5 to 8 mm, and preferably from
about 3.0 to 5
mm. The conic constant of the posterior surface is advantageously selected
from the range of
about +3 to about -14 reported for NCL, preferably from about -1 to -8.
The central radius Ro of the anterior optical surface 8A is selected to be
either larger
than about +3 mm or smaller than about -3 mm, and preferably larger than from
about +5 mm
or smaller than about -5 mm.
The conical constant of the anterior optical surface 8A is selected from the
range from
+6 to -22 reported from human NCL, preferably from the range between about -1
to -8 mm.
The anterior optical surface 8A may be formed partly or fully by a spherical
surface
or a parabolic surface. In that case the central posterior optical surface 8B
should be
preferably hyperbolic with the conic parameter selected in such a range so
that the whole lens
has hyperbolic aberration.
Preferably though, at least the major part of the anterior optical surface 8A
is a
hyperboloid surface, particularly the outer optical zone. The central optical
zone of the
anterior optical surface having diameter between about 1.5 to 4 mm,
advantageously between
about 2 and 3.5 mm, can be formed by parabolic or spherical surface in order
to further
improve the near focus resolution.
Fig. 2 shows schematically one example of the preferred optical profile of the
lens
according to the invention. It should be appreciated that different eyes
require different
refractive power of the implanted lens.
Most of the current IOLs are not bioanalogic since they are designed to
simulate just
the basic optical function of NCL, i.e. to provide the basic refractive power
needed to focus a
distant object on retina. Depending on the specific eye, the basic refractive
power is usually
between 15 and 30 Dpt, with some deviations on either side. This requirement
can be met by
an approximately monofocal (usually spherical) rigid lens located somewhere
near of the
principal plane of the NCL. Since most detailed images are projected onto a
relatively small
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part of retina (macula) located on the optical axis, and since many of our
activities are
performed at small eye aperture (constricted pupil), most IOLs are
significantly smaller than
the NCL (4.5 to 6 mm for most IOLs as opposed to 9.5 to 10.5 mm for the NCL).
The small
size of optics is preferred by some IOL manufacturers for easier adaptation of
such IOL for
implantation through a small incision. For the same reason, most IOLs are made
from a soft,
elastic material that allows implantation through a small incision in a
deformed (folded,
rolled, etc.) shape. This deformability has no relation to the optical
function, however.
Small size of optics has its disadvantages, however. IOL edges may reflect
light at
large pupil opening (e.g., during night driving) and cause glare, halos and
other adverse
effects. Besides, a small optic cannot project all peripheral and off-axis
rays that NCL does,
particularly at a large pupil opening. Lastly, a small size optics interferes
with clear visibility
of retinal periphery that is sometimes needed for diagnostics and treatment.
For those
reasons, the large optics similar in size to NCL is preferable over a smaller
one that is used in
most of the current IOLs. Importantly, the whole large optical zone has to
have well defined
geometry to be optically useful. Lenses with meniscoid optical surfaces have
poorly defined
shape particularly in the peripheral region. This may cause unexpected and
disturbing optical
phenomena.
Some modern IOLs are designed to simulate to some extent the accommodation or
pseudoaccomodation of the NCL (i.e. allowing the eye to focus on both far and
near objects).
Various IOLs use different means to achieve this goal: some are using bifocal,
multifocal or
polyfocal optics; others are using designs allowing anterior-posterior shift
of the IOL optics
with respect to the eye; or allow change of optical power by changing mutual
position
between two lenses. Some lenses even change the refractive power due to liquid
transfer
within the lens driven by pressure of cilliary muscles and/or vitreous body,
change of head
position or by a miniature pump.
These designs are sometimes rather intricate contraptions, very different in
size, shape
and material properties from the NCL. This makes them susceptible to various
problems,
such as fibrosis of the capsule or cell ingrowth or protein deposits on their
surfaces that
interfere with their function. In addition, their increased bulk and
complicated design
interferes with the need of all modern IOLs to be implantable through a small
incision. This
requires designs with small-diameter optics and use of materials with high
refractive index
that are more reflective than the NCL, increasing thus the glare and halo
problems.
In most cases, these lenses are using optics of a small diameter, typically
4.5 to 6 mm,
with slender, flexible "haptics" to position the optics in the center of the
optical path. In
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addition, deformable materials are used to allow folding or rolling for
implantation through a
small incision. The surface properties of such IOLs are sometimes modified to
achieve better
biocompatibility (e.g., A. M Domschke in the US Pat. Publ. No. 2012/0147323,
J. Salamone
et al in the US Pat. Publ. No. 2008/0003259).
This common design allows folding the IOL for the implantation through a
relatively
small incision (usually 2 to 3 mm). However, the small IOL size has its own
drawbacks:
The small optics with diameter 6 mm or less may not fully replace the
crystalline lens
of diameter 9 to 10.5 mm if eye aperture is large due to poor light conditions
(causing night
glare, halos, limited peripheral vision etc.) or if the IOL becomes decentered
(causing the
"sunset syndrome" or other problems);
Small optics cannot project all peripheral and off-axis rays the NCL does,
reducing
thus the imaging performance particularly at large pupil openings (needed for
e.g., night
peripheral vision);
Small optics may complicate or even prevent retinal examination and treatment
(which may be important particularly in the case of diabetics).
In addition, a small IOL size leaves essentially vacant the space that was
originally
occupied by the much larger NCL. Consequently, the vitreous body is allowed to
advance
and its pressure against the retina is partly relieved. This may cause an
increased probability
of retinal detachment after the cataract surgery as reported by J.A. Rowe,
J.C. Erie, K.H.
Baratz et al. (1999). "Retinal detachment in Olmsted County, Minnesota, 1976
through
1995". Ophthalmology 106 (1): 154-159. The same effect may also cause or
facilitate
Cystoid Macular Edema (CME). See Steven R. Virata, The Retina Center,
Lafayette, Indiana:
Cystoid Macular Edema, WEB page.
There is another disadvantage of a small optics and the conventional IOL
design with
haptics: The IOL with optics suspended in the relatively vacant space by means
of relatively
fragile haptics may be sensitive to damage and/or dislocation in case of an
accidental impact
(fall on a slippery surface, car collision, a punch, etc.).
Some problems derived from a small bulk of IOLs and small-diameter optics are
being addressed by IOL designs that fill the space vacated by NCL to a smaller
or larger
extent. There are several approaches to this, each with its own advantages and
disadvantages:
Capsule-filling by a liquid that can solidify into a clear, flexible solid
such as a
silicone rubber. As long as the filler material has similar deformability as
the NCL, it was
expected that this approach would restore the natural lens accommodation
(e.g., Gasser et al.
in US Pat. No. 5,224,957). However, the materials used so far often cause
fibrosis and
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opacification of the capsule. Besides, it is difficult to control the shape
and optical parameters
of the in situ formed IOL
Implantation of a large, bulky IOL in a highly deformed shape that allow
implantation
through a reasonably small incision and fills a significant part of the
capsule. This approach
was tried with hydrophobic memory polymers that can be "frozen" in a highly
deformed
shape for implantation, and returns into the original functional shape upon
heating to body
temperature (Gupta in US Pat. No. 4,834,750 and US Pat. No. RE 36,150).
However, the
hydrophobic memory polymer is very foreign material and causes similar
problems like the
materials used to fill the capsule.
Similar approach was also tried with hydrogels. Very large IOLs, mimicking
size and
shape of the natural lens, have been implanted into the vacated capsule (e.g.,
Wichterle '732
and Stoy '283). The problem of these particular IOLs was their peculiar
optics. These lenses
had meniscoid anterior optical surfaces that deviated strongly from the
geometry of an NCL.
The meniscoid shape was formed by solidification of a free surface of the
monomer mixture,
and there was a problem with control of the optical properties of such IOLs.
In addition,
these lenses were often too bulky for implantation through a small incision.
Moreover, some
of the hydrogels used in these lenses lacked the fixed negative charge, and
such hydrogels
have tendency to calcify sometime after their implantation. Some other capsule-
filling lenses
(Sulc et al. '083 and '903) had anterior protrusions touching the iris and
stabilizing thus the
lens in the approximately central position but causing various problems such
as blockage of
the liquid flow, deformation of lens optics and iris erosion.
Another approach was implantation of a hollow lens (or a lens shell) that was
filled
after implantation by a liquid solidifying in situ (e.g., Nakada, et al. in US
Pat. Nos.
5,091,121 and 5,035,710).
Another approach was implantation of dual-optics IOLs with two lenses, one
being in
contact with anterior and the other with the posterior capsule, both lenses
being kept apart by
flexible members or connectors (US Pat. No. 4,946,469; US Pat. No. 4,963,148;
US Pat. No.
5,275,623; US Pat. No. 6,423,094; US Pat. No. 6,488,708; US Pat. No.
6,761,737; US Pat.
No. 6,764,511; US Pat. No. 6,767,363; US Pat. No. 6,786,934; US Pat. No.
6,818,158; US
Pat. No. 6,846,326; US Pat. No. 6,858,040; US Pat. No. 6,884,261).
Such implants filling essentially the whole capsule of the original
crystalline lens have
also some problems:
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Unless made from extremely biocompatible materials with similar hydration and
negative charge as NCL, the anterior face of the IOL may touch the iris and
cause its erosion,
depigmentation, bleeding or inflammation.
Some materials are made more biocompatible by having high equilibrium water
content. However, that decreases their refractive index far below the optimum
value (the
value for young NCL).
Important for the IOL is not only the shape and optics type, but also its
material. An
NCL is composed of an intricate natural hydrogel structure comprising water,
salts, and
polymeric component containing collagenous proteins, polysaccharides and
proteoglycans.
Importantly, the polymeric components contain a considerable concentration of
acidic
ionizable groups, such as carboxylates or sulfates. These groups provide the
lens material
with a fixed negative charge. The hydration and the negative charge influence
the interaction
between the NCL and proteins in the intraocular fluids. Furthermore, its
surface properties
affect the interaction between the lens and cells. It is known that synthetic
hydrogels
containing surface with a fixed negative charge do not attract the proteins
and cells and make
hydrogel more resistant to calcification (Karel Smetana Jr. et al,
"Intraocular biocompatibility
of Hydroxyethylmethacrylate and Methacrylic Acid Copolymer/ Partially
Hydrolyzed
Poly(2-Hydroxyethyl Methacrylate)," Journal of Biomedical Materials Research
(1987) vol.
21 pp.1247-1253), and are not recognized as a foreign body by immune system
(Karel
Smetana Jr. et al, "The Influence of Hydrogel Functional Groups on Cell
Behavior", Journal
of Biomedical Materials Research (1990) vol. 24 pp. 463-470). Although many
IOL
manufacturers avoid materials with carboxylate groups based on the assumption
that
carboxylates attract calcium ions and thereby cause calcifications, there are
several references
to hydrogel IOLs containing carboxylate groups (Wichterle '732, Sulc et al.
'083 and '903,
Stoy in in US Pat. No. 5,939,208, Michalek and Vacik in '093).
Carboxylate groups may be uniformly dispersed in the hydrogel, or concentrated
mainly on the surface forming a gradient of swelling and charge density, as
described e.g. in
Stoy '208 and Sulc et al. US Pat. No. 5,158,832. Typically, the NCL material
contains, on
average, about 66% by weight of water. However, the NCL is structured with
denser core and
more hydrated jacket and the NCL hydration changes with age and from
individual to
individual. Therefore, one cannot assign a single water content value to the
NCL other than
average.
Similarly, various layers of the NCL have different refractive indices. The
refractive
index of the lens varies from approximately 1.406 in the central layers down
to 1.386 in less
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dense layers of the lens. See e.g. Hecht, Eugene. Optics, 2nd ed. (1987),
Addison Wesley,
ISBN 0-201-11609-X. p. 178. Therefore, the optically meaningful equivalent
refractive
index, or ERI, is given as the characteristic of the NCL. Both refractive
index and water
content change with the lens age. Average ERI = 1.441 -3.9x10^-4xAGE,
decreasing thus
from about 1.441 at birth to about 1.414 at 70 years. See 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, FIG. 9.
In addition, the ERI increases with accommodation by about 0.0013 ¨ 0.0015 per
Diopter. See M.Dubbelman et al, "Change In Shape Of The Aging Human
Crystalline Lens
With Accommodation", Vision Research 45 (2005), 117-132Ref pp. 127-128. One
can
speculate that this change of refractive index is related to a change
(decrease) of water
content due to the lens deformation during the accommodation. Disregarding
these
complications, we will use the average ERI = 1.42 unless stated differently.
Interesting to see that it is very difficult, if not impossible to find a
synthetic hydrogel
with same water content and ¨ at the same time ¨ refractive index as the NCL
material.
Specifically, a synthetic hydrogel containing 66% by wt of water would
typically have a
refractive index of about 1.395 rather than 1.42 that would be expected with
hydrogel
containing closer to 50% of water.
The average liquid contents for ERI=1.441 (very young average NCL) would be
40%
of water while for ERI = 1.414 (old average NCL) would need a hydrogel with
water content
about 55% by weight. Since we believe that for bioanalogic IOL material it is
more
important to simulate refractive index than water content of NCL, we have
selected the
desirable average water content range of the IOL according to an exemplary
embodiment of
the invention, to be between 40% and 55% by weight. Of course, this is the
average water
content ¨ similarly as with the NCL, the lens may have various layers with
different water
contents, e.g. inner parts with higher refractive index and outer layers with
lower refractive
index.
A number of prior art references mention IOLs from hydrogels with high water
content, however, they do not recognize the relation between the water content
and the
refractive index value. For instance, Wichterle '732 specifies the desirable
refractive index
value around 1.4 (broadly from 1.37 to 1.45, which is clearly impossible for
known synthetic
hydrogels with the specified water content: at least 60% and preferably 65 to
70% translates
into the refractive index range from 1.39 to 1.405). The examples show
formulations with a
low content of carboxylate groups.
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Sulc et al. '083 and '903 disclose water content at least 70% and
advantageously at
least 90% on the surface or its part, and mentions 55-70% water content in
prior art IOLs. A
core with higher and a casing with lower refractive index are mentioned, and
the core may
have the form of a Fresnel lens. The gradient of both hydration and refractive
index is
optionally obtained by NaOH treatment that achieves reorganization of the
hydrogel covalent
network. Example 1 of this reference shows an IOL with water content 88.5 %,
Example 2
shows the IOL with water content 81%, and Example 4 shows the lens with water
content
91%. No water content is given for Example 3.
Charles Freeman in US Pat. Pub!. No. 2009/0023835 describes a hydrogel
material
with water content lower than 55% and refractive index higher than 1.41 and
the sodium ion
flux in the range of about 16 to about 20 micro.eq-mm/hr/cm2, useful
particularly for phakic
posterior chamber IOLs. No carboxyl or acidic groups are mentioned, although
their presence
is known to increase the ion diffusion flux through the hydrogel.
Hydrogel character of the NCL material has some possible, less obvious but
.. potentially important consequences: its water content is dependent on the
pressure against
the lens. Consequently, the NCL adjusted to the far distance may have a
different water
content, and therefore a different refractive index, than the relaxed lens
adjusted to the near
objects. Since the stress in the NCL adjusted for far distance is not
distributed evenly, a
gradient of swelling and gradient of refractive index may result. This will
create subtle
changes in the optical properties, in addition to the polyfocality of the NCL
surfaces. These
subtle changes may be important for vision, and it will be difficult to
replicate them otherwise
rather than by using a hydrogel of similar physical-chemical and optical
properties, as well as
geometry similar to that of an NCL. In particular, the hydrogel of the NCL
substitute should
have a similar refractive index and capability to change water content by an
external stress
that can be reasonably expected to act on an NCL. Therefore, the hydrogel used
in a
bioanalogic IOL should have a hydraulic flow capability for water.
Therefore, at least the part of the implant contacting the posterior capsule
is made
from a transparent flexible hydrogel material approximating the optical,
hydrophilic and
electrochemical character of tissue forming the natural lens.
The anterior part of the IOL may interfere with, or even block the flow of the
vitreous
humor causing thus increase of TOP and ultimately glaucoma. This design often
requires a
preemptive iridectomy.
Unless made from extremely biocompatible materials with similar hydration and
negative charge as an NCL, the large-area contact between the capsule and
artificial materials
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used in current IOLs sometimes cause the capsule opacifications, fibrosis,
etc.. These
problems are now being solved by the bioanalogic intraocular lens according to
this
invention.
The central optical part 2 is made of a deformable, elastic material, such as
a hydrogel
with equilibrium water content between about 35 and 65%, advantageously
between about
38% and 55% and preferably between about 40% and 50% (all % are weight percent
and
equilibrium water content is water content in equilibrium with intraocular
fluid, unless stated
otherwise).
Deformability of the optical part is important both for the implantation
through a
small incision and for its accommodation function. The optical part may be
constructed as a
hydrogel shell with a core composed from a liquid or a soft gel, as shown in
the Fig. 6A. Fig.
6A shows a cross-sectional view of a lens with the posterior hydrogel jacket
14, the softer
core 15 and the anterior shell 16. The posterior hydrogel jacket 14 is
advantageously integral
with the peripheral supporting part 3 of the lens and contains the fixed
negative charge at
least on its posterior surface. The core 15 can be advantageously made from a
hydrophobic
liquid, such as mineral oil or silicone oil, or from a soft silicone or
acrylic slightly cross
linked gel that can be easily designed and created by those skilled in the
art. Alternatively, the
core can be made or a hydrophilic fluid or a soft hydrogel. The anterior shell
16 can be made
from the same or different material as the posterior hydrogel jacket 14.
In one embodiment, the hydrogel jacket and the soft core 15 have essentially
the same
refractive index so that the major part of the refraction takes place on the
outer optical
surfaces of the lens rather than on its internal interfaces. This can be
achieved e.g. by making
the core from a silicone liquid or a silicone gel having refractive index
around 1.42, and
making the jacket from a hydrogel with water content between about 41 and 45%
of water.
By formulating the hydrogel correctly one skilled in the art can adjust the
water content in the
hydrogel to achieve the substantial match of the refractive indices.
Alternatively, the core and
the jacket can have different refractive indices so that part of the
refraction takes place on the
internal interfaces between materials.
Fig. 6B a cross-sectional view of a lens with an internal interface between
the core 15
and adjacent optical medium 16 that is shaped to form a compound lens, e.g. a
Fresnel lens.
The materials of core 15 and the optical medium 16 have different refractive
indices, and one
of them is advantageously a fluid that can improve both deformability and
refraction. The
zone 15 or 16 (the one with the lower refractive index) can be created by the
modification of
a hydrogel according to the present invention using a laser. The hydrogel
modification can be
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carried out either preoperatively or postoperatively. Advantage of this
arrangement is the
possibility to use hydrogels with high water content and low refractive index
as the basic
construction material, and yet achieve relatively low central thickness of the
lens that allows
implantation through a small incision.
Fig. 6C shows an alternative design of the lens comprising two different
materials.
Material on the posterior side 14 is a hydrogel with high hydration rate and
containing
negatively charged groups. It is the same for the optical and supporting part.
The anterior side
material of core 15 is a material with lower water content and higher
refractive index. The
interface between the two materials is refractive.
Both central optical anterior surface 8A and central posterior optical surface
8B have
a diameter larger than about 5.6 mm, advantageously larger than about 6.5 mm
and
preferably larger than about 7.2 mm. Optimum diameter of the larger of the two
optical
surfaces is larger than about 7.5 mm, advantageously about 8 mm to approximate
the size of
the NCL optics. Such a large optic is usually suitable for convex-concave or
plano-convex
central optical part 2. For a biconvex optical part, the anterior optical
diameter is usually
selected smaller in order to minimize the central thickness of the optical
part. In any case, the
diameter of the anterior optical surface 8A is advantageously not larger than
the diameter of
the central posterior optical surface 8B.
The central optical surfaces 8A and 8B are surrounded by boundaries 9A and 9B
that
are not necessarily circular. The boundaries 9A and/or 9B may be also
elliptical or have a
shape of a truncated circle, in order to facilitate the lens folding and
implantation through a
small incision. Non-circular optical surfaces are particularly suitable for
lenses with a
cylindrical component.
The posterior peripheral supporting surface 11B is formed by a convex surface,
advantageously a hyperbolic or conical surface with the axis identical with
the main optical
axis 1A. This surface is highly hydrophilic and carrying a fixed negative
charge due to a
content of acidic groups such as carboxylate, sulfo, sulphate or phosphate
groups. This
combination of hydration and negative charge prevents a permanent adhesion to
the capsule,
prevents migration of cells, particularly fibroblasts, along the interface
between the lens and
the capsule, decreases irreversible protein adsorption, and discourages
capsular fibrosis and
opacification. The posterior peripheral surface is advantageously limited by a
sharp edge 7B
that further discourages cell migration toward the optical zone.
The anterior peripheral supporting surface 11A is a concave surface with its
apex
located on the optical axis and it is preferably symmetrical along the axis
1A.
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Advantageously it is a conical or hyperbolic surface with its axis coinciding
with the main
optical axis 1A. The surface is advantageously highly hydrophilic and carrying
fixed
negative charge in order to discourage cell adhesion and migration and
anterior capsular
fibrosis. The anterior peripheral surface is advantageously limited by a sharp
edge 7A that
.. further discourages cell migration.
The anterior and posterior peripheral supporting surfaces 11A and 11B together
with
the connecting surface 6 define the shape of the peripheral supporting part 3.
The peripheral
supporting part is convex on the posterior side and concave on the anterior
side, the average
distance between the two surfaces ranging from about 0.05 to 1 mm,
advantageously from
about 0.1 to 0.6 mm and preferably from about 0.15 to 0.35 mm. The optimum
distance
depends on the stiffness of the material that is dependent on water content,
negative charge
density, crosslinking density and other parameters.
If the posterior and anterior surfaces are formed by surfaces of similar
geometry, such
as hyperbolic surfaces, then the peripheral supporting part 3 will have even
thickness. The
arrangement shown in Fig. 7A has the advantage to be readily deformable and
adjustable to
various sizes of the capsule, and two sharp edges 7A and 7B preventing
migration of
fibroblasts toward the optical zone.
The peripheral supporting part 3 can be also made less or more deformable by
increasing or decreasing its thickness from the rim toward the center, as
shown in Figs. 7B
.. and 7C, respectively. These figures also show various alternative
arrangements of edges 7A
and 7B.
The anterior surface 4 of the implant is shaped to avoid any permanent contact
with
iris that could cause iris erosion, pupilar block, iris pigment transfer to
the implant and other
problems. Such a contact could also interfere with the flow of the intraocular
fluid causing
.. thus adverse changes of the intraocular pressure. It could also interfere
with the contraction of
the pupil as to prevent so called near myosis that helps the near focus both
by the natural lens
and by the implant according to the invention. Therefore, the anterior central
optical surface
8A part is partially sunk due to the anterior peripheral supporting surface
11A concavity and
due to positioning the boundary 9A under the plane defined by the anterior
boundary 7A. The
.. central anterior surface 8A is a plane, a convex surface or a concave
surface with its anterior
apex 10A not exceeding the uppermost point of the lens (the higher of 7A and
7B) by more
than about 0.25 mm, advantageously not exceeding the upper rim at all and
preferably having
the anterior apex 10A bellow the uppermost point 7A by at least 0.1 mm.
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At least the major part (including the central optical surfaces 8A and 8B) of
both
anterior and posterior surfaces 4 and 5 are defined by rotation of one or more
conic sections
around the main optical axis 1A. wherein the term "conic section" includes a
segment of a
line for purpose of this application. The surfaces defined by the rotation
will include a plane
perpendicular to the axis and conical surface symmetrical by the main optical
axis 1A. The
peripheral supporting part is convex on the posterior side and concave on the
anterior side,
the average distance between the two surfaces ranging from about 0.05 to 1 mm,
advantageously from about 0.1 to 0.6 mm and preferably from about 0.15 to 0.35
mm.
In at least one embodiment, the lens according to the invention is
manufactured by
solidification of liquid polymer precursors. In the preferred embodiment, the
solidification
takes place in contact with a solid mold, particularly a mold made of a
hydrophobic plastic.
It can be appreciated that the surface microstructure of a polymer depends on
the
environment in which its solidification took place. The surface microstructure
will be
different if the solidification occurs on the solid liquid interface that if
it takes place on the
liquid-liquid or liquid-gas interface. Preferably, at least all optical
surfaces are created by
solidification of the precursor on a solid interface. Even more preferably,
whole surface of the
implant is formed by solidification of a liquid precursor against a solid
surface, particularly a
hydrophobic plastic surface. Preferred plastic for the mold is a polyolefin,
and particularly
preferred plastic is polypropylene. The polyolefin has low polarity and has
low interaction
with highly polar monomers that are used as hydrogel precursors. Likewise, the
hydrogel
formed by the liquid precursor solidification has very low adhesion to the
mold surface and
can be cleanly detached without even a microscopic surface damage. This is
important for
both optical properties and for long-term biocompatibility of the implant.
Manufacturing a relatively large lens of a precise shape by molding is
difficult. It is
recognized by those skilled in the art that any solidification of the liquid
precursor is
accompanied by the volume shrinkage that may even exceed 20 percent. In a
closed mold of
a constant volume, such a shrinkage will prevent copying of the internal mold
surface and
cause formation of vacuoles, bubbles, surface deformities and other
imperfections. This is the
main reason why the meniscus casting methods described above were used for IOL
molding.
Other inventors have described a method and a mold design allowing the excess
of monomers
to be transported from adjacent spaces by the suction created by the volume
contraction
(Shepherd T., US Pat. No. 4,815,690). However, this method cannot be used in
cases where
the liquid precursor gellifies at a low conversion (e.g., 5 to 10 percent) due
to the crosslinking
polymerization.
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We have discovered a different method for the volume shrinkage compensation,
namely, decrease of the internal mold cavity volume due to the deformation of
certain mold
parts. The mold depicted in Fig. 8 is composed from two parts 18A and 18B, the
part 18A
being used for molding the anterior surface 4 and the part 18B for molding the
posterior
surface 5.
The shaping surface 19B of the part 18B has a shape needed to form the
posterior
optical surface 8B of the lens. The peripheral part 22B of the molding surface
has a diameter
larger than the diameter of the lens and advantageously a hyperboloid or
conical shape
The part 18A has the shaping surface 19A that is divided into the central part
21A
shaping the anterior optical surface 8A of the lens, and the peripheral part
22A of the
diameter larger than the diameter of the lens. The peripheral part 22A has
advantageously a
hyperboloid or conical shape. The peripheral surface 22A is substantially
parallel to the
corresponding surface 22B of the part 18B.
The diameter of the molding the mold parts 18A and 18B are larger than
diameter of
the lens and advantageously they are the same. One of the surfaces for 22A or
22B is
equipped with a relatively thin and deformable barrier 20 with inner surface
corresponding to
the geometry of the surface 6 of the lens. The height of the part 20 is
typically between about
0.05 mm and 1.3 mm, and its thickness is lesser than its height. The profile
of the part 20 is
advantageously wedge-like or triangular. At least one of its surfaces is
advantageously
parallel to the optical axis 1A. The barrier 20 may be separate from the parts
18A and 18B,
but advantageously it is an integral part of one of them. Advantageously, this
part 20 is
located on the concave surface 22B. In a preferred mode of the operation, the
liquid precursor
is filled into the concave mold part 18B in a slight excess to reach over the
barrier 20, and
then it is covered with the part 18A. The mold is constructed in such a way
that the only
contact between parts 18A and 18B is via the part 20. The solidification of
the precursor
generates its contraction and the consequent decrease of the pressure in the
mold cavity. At a
low conversion, the additional liquid precursor is pulled into the mold
cavity. Once the gel-
point is reached due to the crosslinking, the precursor cannot flow anymore.
The decreased
pressure will cause deformation of the part 20 and decrease of the distance
between parts 18A
and 18B and the consequent decrease of the molding cavity volume. The two-part
mold for
the IOL according to the invention is preferably made by injection molding
from a
polyolefin, advantageously from polypropylene.
The preferred liquid precursor for the invention is a mixture of acrylic
and/or
methacrylic monomers with crosslinkers, initiators and other components known
well to
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those skilled in the art. The preferred precursor composition comprises a
mixture of acrylic
and/or methacrylic monoesters and diesters of glycols where monoesters are
hydrophilic
components and diesters are crosslinkers. The preferred precursor also
comprises acrylic
and/or methacrylic acid or its salts. It advantageously comprises also a UV
absorbing
molecule with a polymerizable double bond, 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 are ionizable groups bearing a negative
charge,
such as carboxylate, sulfate, phosphate or sulfonate pendant groups. They may
be introduced
by copolymerization with appropriate monomers bearing such groups, such as
methacrylic or
acrylic acid. In this case, the ionogenic functionality will be uniformly
dispersed in the
hydrogel. Particularly advantageous are hydrogels with ionogenic groups
concentrated
mainly on the surface with the consequent gradient of swelling and charge
density. Such
gradients can be created by after-treatment of molded lenses, e.g. by methods
described in
Stoy '208 and Sulc et al. US Pat. Nos. 5,080683 and 5,158,832.
Other methods include, e.g. grafting of monomers comprising ionogenic groups
on
the lens surface. It is understood that only a part of the lens surface may be
treated to contain
high concentration of ionogenic groups, or that different parts of the surface
may be treated
by different methods.
The lens according to the invention can be implanted in the deformed and
partly
dehydrated state. The controlled partial dehydration can be achieved by
contacting lens with a
suitably hypertonic aqueous solution of physiologically acceptable salts, such
as chlorides,
sulfates or phosphates magnesium or monovalent ions, such as sodium or
potassium. Salt
concentration can be adjusted to achieve hydration between about 15% and 25%
by weight of
the liquid. The lens in the hypertonic solution can be advantageously
sterilized by
autoclaving.
Another method for preparing the hydrogel lens for implantation through an
incision
with reduced size is plastification of the hydrogel by a non-toxic organic
water-miscible
solvent, such as glycerol or dimethylsulfoxide, in such a way that the
plasticized hydrogel has
softening temperature above ambient but lower than eye temperature. Such
composition and
process is described e.g. in Sulc et al, US Pat. No. 4,834,753 that is hereby
incorporated by
this reference.
The lens according to at least one embodiment of the invention is
advantageously
implanted in the state of the osmotic non-equilibrium to adhere to the tissue
temporarily. The
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osmotic non-equilibrium allows the lens centering by adhering it against the
posterior capsule
while the capsule shrinks around it. Once the lens is enveloped by the
capsule, its position is
stabilized. The osmotic non-equilibrium can be achieved in various ways:
soaking the lens
prior to the implantation in a hypertonic salt solution, e.g. in a solution of
10% to 22% by wt.
NaCl, advantageously 15% to 19% by wt.; replacing water prior to the
implantation by a
smaller concentration of a water-miscible solvent, such as glycerol or
dimethylsulfoxide; or
implanting the lens in the state in which the iogenic groups are not fully
ionized, i.e. in the
acidic state prior to the neutralization, and letting the neutralization
proceed spontaneously in
situ by positive ions from the body fluids. The lens achieves its osmotic
equilibrium
spontaneously in hours to days after the implantation.
The lens shape is being formed preferably by crosslinking copolymerization of
methacrylic and/or acrylic esters and salts in the closed two-part mold.
The shape of the lens can be adjusted after the molding by removing some part
of the
lens, e.g. by cutting off part of the supporting part, by drilling the lens
outside the optical
zone etc. The shape adjustment can be made in the hydrogel or the xerogel
(i.e. non-hydrated)
state. We have found that the negatively charged hydrogel material even allows
use of
methods developed primarily for living tissues (incl. NCR), such as ultrasonic
phacoemulsification, cauterization or femtosecond laser treatment. These
methods allow
shape adjustment even in the fully hydrated hydrogel state. The femtosecond
laser may be
used even for formation of cavities inside the hydrogel lens that can be used
to form new
refractive members in the lens, for instance as a refractive cylindrical lens
for astigmatism
compensation. In the case that the matter removed by the shape adjustment
(e.g., by a laser
treatment) is water-soluble and substantially non-toxic, such an optical
adjustment can be
conceivably achieved even post-operatively in situ. The composition of the
hydrogel in at
least the treated part of the lens should be advantageously based on esters of
polymethacrylic
acid. It is known that such polymers are capable of depolymerization to their
parent
monomers (such as 2-hydroxyethyl methacrylate or methacrylic acid) that are
well soluble,
easily diffusible compounds of low toxicity. Other polymers, such as
polyacrylates, polyvinyl
compounds or polyurethanes do not have this advantage.
The invention is further illustrated by the following Examples that are meant
to
provide additional information without limiting the scope of the invention.
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EXAMPLE 1:
The following monomer mixture was prepared: 98 weight parts of 2-hydroxyethyl
methacrylate (HEMA), 0.5 wt% of triethyleneglycol dimethacrylate (TEGDMA), 1
wt% of
methacryloyloxybenzophenone (MOBP), 1 wt% of methacrylic acid, 0.25 wt% of
camphorcquinone (CQ) and 0.05 wt% of trieathanolamine (TEA). The mixture was
de-aired
using by carbon dioxide and filled into two-part plastic molds shown
schematically in Fig. 8
where 18B is the part of the mold for molding the a posterior lens surface,
18A is the part of
the mold to shape the anterior part of the surface of the lens. Both parts are
injection molded
from polypropylene (PP). The shaping surface 19B of the part 18B has shape
formed by two
concentric hyperboloids. The central part of the surface has the diameter 3
mm, central radius
of 3.25 mm and conic constant -3.76 while the peripheral is hyperboloid with
central radius
of 3.25 mm and conic constant -6.26. The molding surface is equipped with a
protruding
circular barrier 20 on diameter 8.5 mm that has asymmetric triangular profile,
height 0.2 mm.
This lip is designed to shape the connecting surface 6 in Fig. 3A.
The part 18A has the shaping surface 19A that is divided into the central part
21 of
diameter 6.8 mm and the peripheral part 22A of the diameter 13 mm. The
peripheral part is
formed by a hyperboloid with the central radius 3.25 mm and the conic constant
-6.26. The
peripheral hyperbolic surface is parallel to the corresponding surface of the
part 18B. The
central portion of the part 18A has the central radius of curvature -20 mm and
conic constant
h=1.
About 0.1 ml of the monomer mixture is pipetted into the part 18B, then it is
covered
by the part 18A that is carefully centered and pressed gently against it by a
small weight. The
only direct contact between the parts is the circular contact between the
barrier 20 and the
peripheral part of 22A. The mold is then illuminated for 10 minutes by a blue
light at the
wavelength 471 nm. The light initiates polymerization of the monomers
accompanied by
gelling at a relatively low conversion and by volume contraction that is
roughly proportionate
to conversion. The contraction of the soft gel creates a mild vacuum that
pulls both parts of
the mold together. The conical peripheral part 22A of the mold 18A presses
against the
barrier 20, deforms it slightly and closes to the part 18B to reduce the
volume of the molding
cavity. This compensates for the volume shrinkage due to the polymerization.
The described
mold design is particularly suitable for production of relatively bulky IOLs
from materials
with high polymerization contraction that achieves gel-point at a relatively
low conversion.
The mold parts are separated and the xerogel lens, the exact copy of the mold
cavity,
is neutralized by solution of sodium bicarbonate and extracted with isotonic
solution. The
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linear expansion factor between the xerogel and hydrogel lens is 1.17. After
evaluation of
optical properties the lens was immersed in the 18% by weight aqueous solution
of NaCl in a
sealed blister package and sterilized by autoclaving.
EXAMPLE 2:
The following monomer mixture was prepared: 94 weight parts of 2-hydroxyethyl
methacrylate (HEMA), 0.5 wt% of triethyleneglycol dimethacrylate (TEGDMA), 4.5
wt% of
methacryloyloxybenzophenone (MOBP), 1 wt% of methacrylic acid and 0.25 wt% of
dibenzoylperoxide. The mixture was de-aired using nitrogen carbon and filled
into two-part
plastic molds shown schematically in Fig. 8. The shaping surface 19B of the
part 18B has a
shape formed by two concentric surfaces. The central part of the surface has
the diameter 3
mm, central radius of 3.00 mm and conic constant 1 while the peripheral
section is a
hyperboloid with central radius of 3.25 mm and conic constant -6.26. The
molding surface is
equipped with a protruding circular barrier 20 on diameter 8.8 mm that has
asymmetric
triangular profile, height 0.15 mm. The inner side of the barrier 20 is
designed to shape the
connecting surface 6 in Fig. 3A.
The part 18A has the shaping surface 19A that is divided into the central part
21 of
diameter 7.1 mm and the peripheral part 22A of the diameter 13 mm. The
peripheral part is
formed by a hyperboloid with the central radius 3.25 mm and the conic constant
-6.26. The
peripheral hyperbolic surface is parallel to the corresponding surface of the
part 18B. The
central portion of the part 18A is a plane perpendicular to the optical axis
1A.
About 0.1 ml of the monomer mixture is pipetted into the part 18B, then it is
covered
by the part 18A that is carefully centered and pressed gently against it by a
small weight. The
only direct contact between the parts is the 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 parts are separated and the xerogel lens, the exact copy of the mold
cavity,
is neutralized by solution of sodium bicarbonate and extracted 3 times with
ethyl alcohol and
5 times with isotonic solution. The lens was yellow with complete absorption
of UV light and
part of the blue visible light. The linear expansion factor between the
xerogel and hydrogel
lens is 1.13. After evaluation of optical properties the lens was immersed in
the 15% by
weight aqueous solution of NaCl in a sealed blister package and sterilized by
autoclaving.
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EXAMPLE 3:
The following monomer mixture was prepared: 94.5 weight parts of 2-
hydroxyethyl
methacrylate (HEMA), 0.5 wt% of triethyleneglycol dimethacrylate (TEGDMA), 5
wt% of
methacryloyloxybenzophenone (MOBP) and 0.25 wt% of dibenzoylperoxide. The
mixture
was de-aired using nitrogen carbon and filled into two-part plastic molds
shown
schematically in Fig. 8. The shaping surface 19B of the part 18B has a shape
formed by two
concentric surfaces. The central part of the surface has the diameter 6.5 mm,
central radius of
4.5 mm and conic constant 0 while the peripheral section is a hyperboloid with
central radius
of 4.25 mm and conic constant -8. The molding surface is equipped with a
protruding
circular barrier 20 on diameter 9.3 mm that has asymmetric triangular profile,
height 0.35
mm. The inner side of the barrier 20 is designed to shape the connecting
surface 6 in Fig. 3A.
The part 18A has the shaping surface 19A that is divided into the central part
21 of
diameter 6.4 mm and the peripheral part 22A of the diameter 13 mm. The
peripheral part is
formed by a hyperboloid with the central radius 4.25 mm and the conic constant
-8. The
peripheral hyperbolic surface is parallel to the corresponding surface of the
part 18B. The
central portion of the part 18A is a surface of diameter 6.4 mm, central
radius -3.75 mm and
conic constant -6.
About 0.1 ml of the monomer mixture is pipetted into the part 18B, then it is
covered
by the part 18A that is carefully centered and pressed gently against it by a
small weight. The
only direct contact between the parts is the 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 parts are separated and the xerogel lens, the exact copy of the mold
cavity,
is extracted. The lens is then treated by a quaternary base as described in
the reference Stoy
'208.
The z lens from the clear, electroneutral crosslinked hydrophilic polymer has
a
surface created by a gradiented layer with high hydration and negative charge
density. The
lens was neutralized by solution of sodium bicarbonate and extracted 3 times
with ethyl
alcohol and 5 times with isotonic solution. The lens was clear with complete
absorption of
UV light. The linear expansion factor between the xerogel and hydrogel lens is
about 1.12.
.. After evaluation of optical properties the lens was immersed in the
isotonic aqueous solution
of NaCl in a sealed blister package and sterilized by autoclaving.
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EXAMPLE 4:
Raman spectra of the hydrogels of the invention show a significant difference
between the original hydrogel and the same hydrogel subjected to TPA by
exposure to the
focused beam of a femtosecond laser. Namely, there is a significant difference
in the ratio of
the signal at 3420 cm' corresponding to water, and at 2945 cm' corresponding
to the CH2
group from the polymer backbone (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 modified and unmodified material (see Fig 9B). The Raman scan
across the
modified strip within the hydrogel showed increased water content in the
treated area (see Fig
9C). On the other hand, the Raman spectrum of the region below 2000 cm' showed
no
indication of new chemical groups. This is consistent with replacement of part
of the
polymer mass for an aqueous liquid through a mechanism such as
depolymerization.
These and other advantages of the present invention will be apparent to those
skilled
in the art from the foregoing specification. Accordingly, it will be
recognized by those
skilled in the art that changes or modifications may be made to the above-
described
embodiments without departing from the broad inventive concepts of the
invention. It should
therefore be understood that this invention is not limited to the particular
embodiments
described herein, but is intended to include all changes and modifications
that are within the
scope and spirit of the invention as defined in the claims.
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