Ординатура / Офтальмология / Английские материалы / Biomaterials and regenerative medicine in ophthalmology_Chirila_2010

implanted in primate eyes, reported no difference in the optical outcomes among the materials tested. The curvature change achieved when hydrogels were tested in primates (which have a Bowman’s layer present in the anterior stroma) were encouraging over considerable periods from months to years (Mester et al., 1972; Binder et al., 1981/82; Werblin et al., 1983; Beekhuis et al., 1986; Beekhuis et al., 1987; McCarey et al., 1990; McDonald et al., 1993; Parks et al., 1993). In another study, refractive changes that had been observed in rabbits (without Bowman’s layer) did not occur in the first human patient (with Bowman’s layer present) and it was suggested that moving the lens more anteriorly might help (Sendele et al., 1983). The depth at which the lenses of differing water contents were implanted was also considered and tested. Etafilcon A lenses with a range of water contents (58%, 68% and 72%) were implanted at different depths in the monkey stroma and data showed that the higher water content materials did not necessarily give the best refractive data when implanted at 60% depth (Beekhuis et al., 1986). Deeper implants of Lidofilcon A placed at >79% corneal depth, were found to give little, or less predictable, refractive outcomes than those mid-depth or shallower in the cornea (Koenig et al., 1984; McCarey et al., 1990). However, deeper implants were generally well tolerated by corneal tissue, as demonstrated with lenses made of Lidofilcon A and Perfilcon A (McCarey,

1991; Parks et al., 1993) and Surfilcon A (Koenig et al., 1984). It is likely that a contributing factor to this outcome was the availability of glucose to the tissue forward of the implant being increased if the lens was deep (McCarey and Schmidt, 1990), rather than the type of hydrogel material that was used. Often these deep stromal implants were associated with bulging of the posterior stroma and endothelium into the anterior chamber (Koenig et al., 1984; McCarey, 1991; Parks et al., 1993). In addition, the stability of the refractive outcomes with deeper hydrogel implants was not always assured and the contraction of stromal tissue that occurs during wound healing was implicated in at least one study (McCarey et al., 1990).

Factors such as the design and size of the implant, inconsistent microkeratome cuts and decentration caused by migration of the implants were reported to have caused poor outcomes in some cases (Beekhuis et al., 1987; McCarey et al., 1990; McDonald et al., 1993), as did contaminants associated with materials used for implants (Beekhuis et al., 1987; Werblin et al., 1992b). Overall, clinical complications that developed with hydrogel intracorneal lenses included oedema and inflammation, crystalline deposits, accumulation of extracellular matrix around the implant, epithelial ingrowth, fibrosis around implant and ulceration of tissue forward of the implant (McCarey and Andrews, 1981; Sendele et al., 1983; Samples et al., 1984; Yamaguchi et al., 1984b; Beekhuis et al., 1987; McCarey et al., 1990; Parks et al., 1993). Variation in the biological response to the same material was also evident (Parks et al., 1993). While many implants were well tolerated, thinning

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of the epithelium forward of the implant and the presence of keratocytes and debris at the stromal–lens interface were commonly reported in cases where histology was conducted (Sendele et al., 1983; Samples et al., 1984; Yamaguchi et al., 1984b; Beekhuis et al., 1987; McCarey, 1991; Werblin et al., 1992b; McDonald et al., 1993). Data arising from these studies with hydrogels showed notable similarities to those that tested the biocompatibility of acellular biological tissue for intracorneal implants.

Guiding principles arose from this work with hydrogels, with these studies flagging the significance of high permeability and biocompatibility as essential to the success of synthetic materials as intracorneal implants. The possibility of slow stromal alterations occurring in response to hydrogel intracorneal implants, either by remodelling or by encapsulation of non-biological hydrogel, material, was raised (McCarey and Andrews, 1981). Later, it was noted that the need for biocompatibility of hydrogels depended on the toxicity, stability and solute–solvent permeability of materials and, based on this, it was recommended that materials be free of residual monomers, cleaning agents and debris that could stimulate inflammation (McCarey et al., 1988). Further work lead to the identification of requirements for the use of hydrogels for intracorneal implants to correct refractive error that included: minimising the thickness of the implant to reduce hypoplasia of the overlying epithelium; provision of the refractive power correction should come from the optics of the implant; simplification of the surgical procedure; the permeability of the implant should be retained in the long term to provide for nutritional biocompatibility (McCarey, 1991). Implant thickness was acknowledged to be an issue for glucose transport and recommendations were made that lens thickness be minimised (McDonald et al., 1993). Given that hydrogel thickness could be a limiting factor in intracorneal lens design, it was clear that developments in surgical techniques, materials and lens designs were required to address these issues.

New microkeratome technology allowed the creation of hinged corneal flaps for use in LASIK surgery. Concurrently, new materials were developed for use as intracorneal implants. The combination offered the opportunity to place the lens under the corneal flap in a sutureless synthetic keratoplasty procedure which was potentially reversible (corneal inlay). Hyperopic correction was the target of the new technology since the optical designs were relatively simple and myopia was being treated using LASIK as it was easier to flatten the cornea effectively than to steepen it at that time (Esquenazi et al., 2006). Nutrapore, described as a microporous hydrogel material with 70% water content and a refractive index close to corneal tissue, was made into a thin hyperopic design lens to correct up to +6D (PermaVision from Anamed

Inc., Lake Forest, California, USA). Confocal microscopy on PermaVision implants placed under a 150 μm corneal flap in rabbits showed a minimal tissue response over 6 months to the presence of implant in an outcome that

was attributed to the thin design of the inlay made from what was regarded to be a highly permeable material (Ismail, 2002). Refractive data on PermaVision implants came from another study (Michieletto et al., 2004), where hyperopic implants made from a 78% water content version of the hydrogel with a central thickness ranging from 25 to 60 μm were maintained in the corneas of patients with promising optical outcomes and relatively few complications over the 6 month period. A third study (Alio et al., 2004) on PermaVision Nutrapore implants in 11 eyes reported early, and serious, complications during the 6-month period including patient discomfort and loss of best- corrected visual acuity (BCVA), combined with edge opacities that did not respond to steroids and were attributed to the presence of epithelial cells under the inlay. In another study (Knorz, 2005), stable refractive outcomes were shown after ‘first-generation’ PermaVision inlays were implanted in 21 eyes of 12 patients, but the majority developed corneal haze which started around the implant edge at 3 months and gradually covered the lens in up to 3 years. A follow-up study with PermaVision lenses implanted in 23 eyes of 20 patients over 2 years reported relative refractive predictability with 70% ±0.5D but with many complications such as decentration, induced astigmatism, stromal opacification, night halos and glare (Ismail, 2006).

Confocal microscopy was used to compare hyperopic PermaVision inlays with LASIK performed using a femtosecond laser (IntraLase) and identified interface particles and activated keratocytes in both groups with chronic central corneal epithelial thinning in the inlay group (Petroll et al., 2006). It is possible that the permeability of hydrogel materials such as these is not maintained in the long term due to fouling or other reasons. Only recently, Larrea et al. (2007) developed a three-dimensional model of solute transport in the cornea which simultaneously addressed the axisymmetric oxygen and glucose diffusion for different lens permeabilities and positions of the implant in the cornea. Simulations using inlay with 3 mm diameter and 20 μm thick with a range of permeabilities (supplied by BioVision AG) showed the mid-posterior stroma (75% depth) to be the optimal position for the inlay to be implanted to enhance the supply of oxygen and glucose to the cornea. This publication (Larrea et al., 2007) also reminds us that transport of other solutes through the cornea, such as lactate, is significant and may be affected by the presence of an implanted lens. Lactic acid is a product of cell metabolism that increases in concentration under hypoxic conditions and can lead to accumulation of lactate which causes stromal oedema.

4.4.5Current corneal implant technologies

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implant, increasing the permeability of the material or decreasing the thickness of the implant, or a combination of all these approaches. Additionally, placing the implant superficially is less damaging to the corneal tissue and offers reversibility. The currently available technologies in the synthetic intracorneal implant area reflect a good understanding of the knowledge gained from this research. A range of materials and designs have been tested as corneal inlays and onlays and these are well documented in review articles (Hughes and Chan, 2001; Xie et al., 2001; Sweeney et al., 2008). Currently, the majority of technologies emerging in this area are corneal inlays aimed at presbyopia to provide an alternative to reading glasses for the growing market arising from the ageing ‘baby-boomer’ generation (Pieper, 2006). These presbyopic implants make use of synthetic materials in small-diameter, thin designs and are generally being implanted intracorneally in the non-dominant eye of subjects with no other refractive error (emmetropes), some of whom may have had previous or concurrent refractive surgery (Alio, 2007; Dalton,

2008). In each case, the inlays and accompanying surgical procedures are directed toward a solution that is convenient, minimally invasive, preserving of distance visual acuity and reversible.

ReVision Optics has manufactured the PermaVision hydrogel into small, thin, plano-presbyopic lenses to treat presbyopia (Presbylens from ReVision Optics, Lake Forest, California, USA). The material is iso-refractive with corneal tissue and works by changing the anterior curvature of the cornea with a central near-add and an additional draping effect that improves intermediate vision. This ‘refractively neutral’ approach is expected to reduce the incidence of halos and glare (Bethke, 2007). The small 1.5–2  mm diameter inlays are implanted into the central cornea using either a microkeratome flap or delivered using an applicator into a tunnel pocket made with a femtosecond laser. General reports suggest that these inlays show predictable and stable visual outcomes in a procedure that is reversible and exchangeable (Slade, 2006; Dalton, 2008; Lang et al., 2008). ReVision Optics is working towards a combination ‘one-visit’ therapy where LASIK and an inlay are used concurrently in a presby-LASIK approach (Bethke, 2007).

BioVision AG has also designed a small, thin lens to improve near vision for emmetropic presbyopes using a monovision approach (Invue from BioVision AG, Brüggs, Switzerland). The 3 mm diameter inlay is made of a permeable acrylic hydrogel material that is implanted in the patient’s nondominant eye. The donut-shaped bifocal lens has a central neutral zone of 1.8 mm and a peripheral zone with refractive power to correct near vision. The implant has a small hole centrally that provides for corneal nutrition without interfering with the optics. The lens is implanted mid-depth in the central cornea in a tunnel made using a microkeratome dissector (Visitome from BioVision AG) or femtosecond laser. The Invue inlay potentially offers a multifocal cornea to emmetropic presbyopes with no astigmatism, allowing

them to read. Some reports suggest that this can be achieved in the majority of cases (Sanchez Leon, 2006; Kymionis et al., 2007a; Sanchez Leon, 2008), but there are several disadvantages reported with this technology, including slow visual rehabilitation, a small loss of contrast sensitivity or quality of distance vision, some night-vision problems such as halo, and complications requiring the removal of the inlay in a small number of patients (Lindstrom,

2005; Bethke, 2007; Dalton, 2008).

AcuFocus has addressed the correction of presbyopia using a small-diameter device (ACI-7000 intracorneal inlay from AcuFocus, Inc., Irvine, California, USA) with a central pinhole 1.6 mm in diameter to increase the depth of field based on the principle of small-aperture optics (Yilmaz et al., 2008). The current version of this inlay is a made from poly(vinylidene fluoride) (PVDF; refractive index 1.42) with carbon nanoparticles to make it opaque. The lens is very thin (approximately 10 μm thick) with a 3.6 mm area around the central hole that incorporates a random arrangement of 160 pores (25 μm) to provide for nutritional flow (Yilmaz et al., 2008). The inlay is placed on the non-dominant eye beneath a conventional corneal flap (approximately 160 μm deep) cut with a microkeratome or a femtosecond laser. ACI-7000 inlays are being implanted in naturally emmetropic patients and previously ametropic patients who have had LASIK surgery (Yilmaz et al., 2008). The depth of the implant appears to be significant in minimising flap complications and preventing a change in shape of the corneal curvature (Yilmaz et al., 2008). Centration of the inlay over the pupil is essential to a good outcome as the opaque ACI-7000 inlay is placed over the pupil to create a fixed aperture setting of 1.6 mm to increase the depth of field, improving intermediate and near vision in presbyopic patients without significantly reducing distance vision (Yilmaz et al., 2008). Clinical testing has shown improved near vision (of about +1.5D) without marked loss of distance vision over a period of 1 year (Dalton, 2008; Yilmaz et al., 2008). In some patients there were problems with glare caused by high light transmission from the pinholes, which also reduced contrast sensitivity. The manufacturers have addressed this by design modifications that resulted in an even thinner implant (5 μm thick), with lower light transmission through the pinholes (Bethke, 2007). Disadvantages of this technology include the appearance of the inlay (black) and some loss of night vision because of reduced light entering the eye (Dalton, 2008). Data from a

Phase III trial of 200 eyes implanted with the ACI-7000 were presented to the US Food and Drug Administration (FDA) in 2007. In the future, the company plans to implant inlays in ametropic patients who are undergoing LASIK with the intention of correcting refractive error and treating presbyopia in a single visit (Bethke, 2007).

Vision CRC (Vision Cooperative Research Centre, Sydney, Australia) has developed a non-hydrogel material with broad potential in the additive refractive keratoplasty area. The higher permeability nature of this material

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does not restrict its use to the correction of presbyopia but offers the potential for broader corrections – including myopia, hyperopia, presbyopia and astigmatism – since thickness is not a limiting factor in the implant design. The group has drawn on the considerable experience gained from their own research and that of others in the field to produce a non-hydrogel material family with characteristics defined as ideal for an intracorneal application. Implants made of the perfluoropolyether (PFPE) -based material create a stable change in corneal curvature when implanted and could be used to treat presbyopia, myopia or hyperopia. This material offers many advantages for corneal augmentation as it is iso-refractive (refractive index of 1.34 (Rice and

Ihlenfeld, 1989)), inert, transparent, flexible and chemically and thermally stable. Biostability of the PFPE material has been demonstrated in samples implanted subcutaneously in sheep over 2 years (G. F. Meijs, K. Schindhelm, R. Odell and H. Chen, unpublished data). PFPE is inherently hydrophobic and some hydrophilicity has been introduced by co-polymerisation with monomers containing a zwitterion, which has also enhanced its anti-fouling properties. Permeability has been chemically induced into the PFPE material (Chaouk et al., 2001) to meet requirements established by implanting model polymer membranes with a calculated range of permeabilities into the feline cornea and monitoring clinical outcomes (Sweeney et al., 1998). Porous PFPE membranes have demonstrated permeability to appropriate corneal metabolites such as glucose and model proteins such as albumin and inulin (Evans et al., 2000; Hughes and Chan, 2001).

The Vision CRC polymer has been tested as an inlay in both animal and human trials. PFPE inlays, 4.3 mm diameter and 80 μm thick centrally, in a hyperopic-shaped design with tapered sides, were well tolerated by corneal tissue over the long term when placed under a microkeratome corneal flap in rabbit corneas (Xie et al., 2006) and in unsighted human subjects (Sweeney et al., 2005; Sweeney et al., 2006; Prakasam et al., 2007a; Sweeney et al., 2008; Sweeney et al., 2009). In both cases, clinical data have shown that corneas remained clear without inflammation, neo-vascularisation or increased redness compared with accompanying sham-operated corneas for 2 years in rabbits (Xie et al., 2006) and over 4 years in humans (Sweeney et al., 2009), as seen in Fig. 4.5. Confocal microscopy of the implanted human eyes showed the presence of activated keratocytes immediately around the lens at

6 months, which was accompanied by reflective deposits on inlay surfaces in a low-grade response that had stabilised by 12 months (Vaddavalli et al., 2007; Sweeney et al., 2009). Histology and electron microscopy on the implanted rabbit corneas showed a similar overall response with activated and degenerative keratocytes, with some accompanying cell debris along the surfaces of the inlay noted at 6 months, which showed no increase at 12 or 24 months (Xie et al., 2006), as seen in Fig. 4.6. Despite the overall good biocompatibility, PFPE inlays in both human and rabbit eyes have shown a

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in the implanted eyes of average 6.5D (range 4–14D) at 48 months with fluctuations related to the patients’ poor fixation in their unsighted eyes as well as to the patients’ age (Sweeney et al., 2008; Sweeney et al., 2009). Reversibility has also been demonstrated when some inlays were removed and the corneal flap replaced (Sweeney et al., 2008; Sweeney et al., 2009). Interestingly, the changed corneal curvature in implanted human eyes caused the formation of iron rings in the peripheral cornea, which were similar to Fleisher’s ring that occurs in keratoconus (Prakasam et al., 2007b). Iron rings are non-pathological and result from iron deposition in the basal layer of the corneal epithelium, usually in response to sudden changes in corneal contour.

The Vision CRC polymer has also been tested as a synthetic corneal onlay. Onlays are placed superficially under the corneal epithelium without disturbing Bowman’s layer and are incorporated into the cornea by epithelial tissue covering the anterior surface of the lens. Refractive correction is obtained by a change in corneal curvature. This is a challenging application for any synthetic material since, in addition to being inert, biostable, nutrient permeable, transparent and mechanically compliant, it must also allow for the growth and stable adhesion of a stratified epithelium on its anterior surface (Evans et al., 2001a; Hughes and Chan, 2001; Xie et al., 2001). The surface chemistry and topography of the polymer are known to be significant in achieving this outcome and have been carefully examined by this group using in vitro modelling systems (Evans and Steele, 1997; Evans and Steele, 1998; Fitton et al., 1998; Dalton et al., 1999; Evans et al., 1999; Dalton et al., 2001; Evans et al., 2001b; Evans et al., 2003). Initial in vivo work by Vision CRC determined that covalently immobilised collagen type I promoted epithelial growth across the surface of a model synthetic polymer implanted in feline eyes (Sweeney et al., 1997; Sweeney et al., 2003). This coating strategy was applied to PFPE lenses, which supported the growth of epithelial tissue implanted in an open pocket made in feline corneas (Evans et al., 2000). Longer-term trials in feline corneas showed that collagen-coated PFPE lens not only supported epithelial growth but also the formation of recognisable basement membrane and components of cell–matrix junctions

(hemidesmosomes and anchoring fibrils) along the epithelial–lens interface

(Evans et al., 2002). In vivo studies have demonstrated that a PFPE lens with a collagen I coat could be glued on to a debrided feline cornea and be incorporated into the cornea by the migration and adhesion of stratified corneal epithelial tissue, as seen in Fig. 4.7.

4.4.6Intracorneal rings

4.7 Light micrograph of a transverse histological section showing a Vision CRC PFPE onlay that had been fully covered with epithelial tissue for 3 months after placement on a debrided feline cornea (inferior side of onlay at original magnification ∞200).

the treatment of disorders involving corneal ectasia and keratoconus. Intacs (Addition Technology Inc., previously KeraVision, Fremont, California, USA) originally developed a one-piece PMMA intracorneal ring, which was implanted in a tunnel made in the peripheral cornea to flatten the central area and reduce myopia. Subsequent designs produced two half-ring segments made from impermeable PMMA, which were easier to implant. This technology offered the advantage of leaving the central cornea undisturbed (Linebarger et al., 2000; Ruckhofer et al., 2001a) but has failed to provide reliable optical outcomes except in the case of very low myopes (Ruckhofer et al., 2001b). Complications reported with intracorneal rings include corneal dehydration, crystal formation, lipid deposits, opacities and difficulties with implant edges which have caused thinning, fibrosis and ulceration of the stroma, keratitis and corneal vascularisation, with some of these difficulties likely to be associated with the stretching of the corneal tissue.

This would be consistent with recent histological evaluation of failed cases of intracorneal rings used in the treatment of keratoconus and post-LASIK keratectasia that showed an abnormal accumulation of fibrotic extracellular matrix components and proteases near the rings, suggesting ongoing lysis and remodelling of corneal stroma (Maguen et al., 2008). Currently, intracorneal rings are being used in the treatment of keratoconus where they may offer some benefit (Kymionis et al., 2007b). Recent figures show that intracorneal rings have been used to treat 50 000 keratoconus cases worldwide since the first implantation 10 years ago, with stable correction provided for most of those cases (Colin, 2008). Outcomes may be improved with use of the femtosecond laser instead of mechanical methods to create the channels used to implant the rings (Carrasquillo et al., 2007). The KeraRing (Mediphacos,