Roger F. Steinert, MD
Contents
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Foldable Intraocular Lenses and Small-Incision Surgery |
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Small-Incision Surgery |
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Materials and Optics |
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Aspheric Optics |
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Biocompatibility |
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CHAPTER HIGHLIGHTS |
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Foldable intraocular lenses and small-incision surgery
The soft implant material that makes up a foldable intraocular lens (IOL) enables a 6mm diameter optic intraocular lens (IOL) to be inserted through a 3mm or smaller incision with minimal trauma. For forceps insertion, the surgeon grasps the IOL, bends it in half with folding forceps, grasps the folded IOL with insertion forceps then maneuvers the lens through the incision until the IOL and haptics are appropriately situated. For insertion with an implantation device, the surgeon grasps the IOL with holding forceps, positions the IOL in the device, inserts the tip of the device through the incision, and implants the IOL. Some implantation injector systems allow for the IOL to be shipped in the unfolded state and then folded into the inserter without manual handling, increasing the ease and reliability of the folding process and reducing the potential for contamination.
As the IOL slides into position inside the capsular bag or within the ciliary sulcus, the optic unfurls to resume its original shape. Visual recovery from the aphakic state is almost immediate. Visual acuity after cataract surgery and IOL implantation is nearly always an improvement over vision through a crystalline lens with a cataract, especially with small foldable IOLs and small-incision surgical techniques. While a major advantage of small-incision cataract surgery is that it minimizes corneal shape changes that induce astigmatism and delay visual recovery,[1–3] the small incision also provides increased surgical and postoperative safety.
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Small-incision surgery
Effect of incision size on astigmatism
Surgically induced changes in the structural stability of the cornea may produce or aggravate corneal astigmatism that may not be amenable to correction postoperatively with a spherocylindrical spectacle lens. Surgically induced astigmatism can adversely affect postoperative refraction and the stability of the patient's vision over time. Incision length, location, shape, orientation, and suture technique and material are factors. Numerous clinical studies have consistently concluded that smaller incisions lead to better early postoperative uncorrected and corrected visual acuity. In addition, rates of induced astigmatism decline as the incisions become progressively smaller, down to near neutrality in incisions between 2.5 and 3mm.[1–9]
Effect of intraocular lens implantation on incision size
Surgically induced astigmatism may correlate with small incision sizes, but forcible insertion of the folded IOL through the wound may widen the wound, and the final incision may not be as small as the initial keratome incision. Steinert and Deacon[10] developed a set of incision-size gauges to study the dimensional stability of incisions during cataract extraction and IOL implantation. Steinert-Deacon gauges (Capital Instruments, Ltd., Wan Chai, Hong Kong) are the equivalent thickness of conventional metal keratomes and are manufactured in 0.2mm increments (Figure 37-1). The gauges were initially tested in 51 consecutive patients undergoing phacoemulsification.
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Figure 37-1 Steinert-Deacon incision-size gauges. |
In the 46 cases of temporal clear corneal incisions performed with diamond keratomes, the initial incision was significantly wider than the keratome blade (0.16 ± 0.08mm; P <0.0001). After phacoemulsification and irrigation–aspiration, the incision was again significantly wider (0.09 ± 0.06mm; P <0.0001). If the incision was not widened before insertion of the folded silicone IOL, significant enlargement occurred once more (0.26 ± 0.05mm; P <0.0001). The final insertion size after forceps insertion of a three-piece silicone foldable IOL was not statistically different than injector insertion of a plate haptic silicone IOL (P = 0.56). The investigators concluded that the incision has limited capacity for elastic deformation and may be vulnerable to tearing or irreversible stretching.
Implantation devices for small-incision surgery
Implantation insertion devices provide uniform folding of the IOL and to allow the IOL to be inserted through smaller incisions than is possible with forceps.[11–12] Unlike forceps, inserters isolate the IOL from the external environment, reducing the risk of introducing surface pathogens into the eye at the time of implantation. For most inserters, the surgeon or technician fills the cartridge with viscoelastic, grasps the IOL with holding forceps, loads the IOL in the cartridge, closes the cartridge, and places it in the handpiece. The surgeon enters the incision with the insertion tip of the handpiece and delivers the IOL into the capsular bag or ciliary sulcus. Not all inserters are appropriate for inserting the IOL into the sulcus.
The next generation of insertion systems eliminates the need for manual loading of the IOL into the inserter. In addition to improved ease of use, an automated loading system brings consistency of loading, thus, reducing the potential complication of damage to the IOL during insertion, and the potential for the bacterial contamination of the IOL due to handling.
The challenge of automated pre-loaded insertion systems stems from bio-material constraints. Foldable IOL materials cannot be kept in their folded state for extended periods without permanent damage to the material. Therefore, the IOL cannot be shipped in a pre-folded compressed condition. Second, the materials in the injection cartridges typically are chosen, in part, due to inherent lubricity that facilitates the passage of the compressed IOL. These materials may interact with the different material of the IOL optic producing undesirable secondary chemical changes. For these reasons, automated systems have designs that allow the IOL optic to be shipped uncompressed and in contact with a different plastic than the injector itself, yet still allow it to be transferred into the compression chamber and inserter without manual instrumentation.
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Materials and optics
In the early 1950s, Scales[13] defined an ideal material for implantation into human tissues as chemically inert, stable, not physically modified by contact with tissues, and acceptable to the body, with no inflammation, foreign body response, or tissue chafe. The desired material would be neither carcinogenic nor allergenic, have the capability of being fabricated into the desired form, be able to resist mechanical strains, and be easy to sterilize.
For ophthalmic applications, the material should be optically transparent, and be able to remain so for long periods, be capable of being manufactured into a high resolving power, be able to block ultraviolet (UV) radiation in the 330–400nm wavelengths, and be implantable through a small incision.
Acrylic polymers
The history of acrylic materials used in foldable IOLs begins with polymethylmethacrylate (PMMA), an acrylic material with excellent tissue tolerance that has been used successfully in ophthalmology, oral and dental surgery, orthopedics, and plastic and reconstructive surgery (see Chapter 36). Acrylic materials are polymers synthesized from esters (monomers) of acrylic acid or methacrylic acid. The methyl ester of methacrylic acid, methylmethacrylate, readily polymerizes to PMM. It is a hard, rigid, strong thermoplastic material with excellent optical clarity. It has low water and gas diffusion constants and is highly resistant to the effects of light, oxygen, and hydrolysis. The pure polymer readily transmits UV light but can be manufactured with UV-absorbing chromophores to block ultraviolet energy from reaching the retina.
Hydrophobic acrylic polymers
These polymers are members of the same family as rigid PMMA, but they are tailored for specific optical and mechanical properties by altering the side groups of a standard methacrylate backbone. PMMA and hydrophobic acrylic polymers are similar in their negligible water content and high refractive indices (e.g., 1.55 for AcrySof (Alcon) and 1.47 for Tecnis acrylic (AMO) compared to 1.49 for PMMA). Unlike PMMA, hydrophobic acrylic materials have a relatively long hydrocarbon side chain for increased flexibility at the typical operating room temperatures of 18–22° C. The glass transition temperature (Tg) for PMMA is 105° C, approximately 15.5–21.5° C for AcrySof,[14] and 13° C for Sensar. The Tg is the temperature at which the polymer chains soften from their low-temperature rigid state to their flexible high-temperature form. Thus the lower the Tg, the easier the IOL material is to fold at standard room temperatures.
The AcrySof lens optic is molded from a phenylethyl acrylate and phenylethyl methacrylate polymer (US Patent 5,290,892; 1994). The Sensar lens optic is cut from a sheet of ethyl acrylate, ethyl methacrylate, and trifluoroethyl methacrylate polymer, then cryolathed into its final shape (US Patent 4,834,750; 1989).
Hydrophilic acrylic polymers
This class of IOLs is made with hydrophilic acrylic (hydrogel) polymers. Commonly, the IOL optic is lathe cut and polished from a composite material.
A variant is the Collamer IOL (STAAR). Collamer is a hydrogel-collagen copolymer consisting of a HEMA-based acrylic copolymer into which about 0.01% porcine collagen and a UV-absorbing chromophore have been bonded.[18] Because of its high water content, it is wet packed in a glass vial containing the sterile lens in balanced salt solution (BSS) and placed in a pouch for shipping. The lens is removed from the vial with blunt forceps and loaded into a disposable plastic cartridge. The cartridge is inserted into a disposable plastic injector.
In clinical use, early- and late-onset calcification of hydrophilic IOL materials has been a common source of optical degradation.[19–20] The most common reason for explantation of hydrophilic lenses is calcification, sometimes many years after the original implantation surgery.[21] An animal model developed by Buchen illustrates the appearance of calcium deposits on the surface of a hydrophilic IOL[22] (Figure 37-2).
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Figure 37-2 Scanning electron micrograph of calcification on the surface of a hydrophilic intraocular lens in an animal model |
Silicone elastomers
Silicone is used in medicine for prosthetic devices, estradiol-releasing vaginal rings, intracoronary stents, subdermal and transdermal implants, intrauterine contraceptive devices, catheters, nasolacrimal intubation tubes, finger joints, and a variety of other clinical situations. In ophthalmology, silicone is used for contact lenses, scleral buckles, keratoprostheses, glaucoma shunts, and IOLs. Silicone is inert, stable at high temperatures, flexible and elastic at a wide range of temperatures, and nonadhesive to tissues. Silicone IOLs are optically clear and have refractive indices ranging from 1.42 to 1.46.
Silicone elastomers are safe, stable, and inert over the long term. Silicone IOLs showed no change in optical performance or surface quality in UV and hydrolytic stress tests designed to mimic 20 years of aging.[23,][24] Questions have been raised about the advisability of implanting silicone IOLs in patients at risk for vitreoretinal surgery, because silicone oil used as an intravitreal tamponade may adhere to the solid silicone IOL optic and obstruct the surgeon's view.[25–27] Acrylic IOLs have frequently been the lenses of choice for patients with a history of ocular inflammation, pseudoexfoliation, and diabetic retinopathy or for patients at risk of future vitreoretinal surgery; however, McLoone et al.[28] showed silicone oil adherence to acrylic lenses in vitro. The mean percentage coating of silicone oil on PMMA IOLs was 20.8% and that on foldable hydrophobic acrylic IOLs was from 17.1 to 21.5% (P = NS). One hydrogel IOL (Aqua-Sense, Ophthalmic Innovations International; Ontario, Calif.) had 17.8%, and another (Raysoft, Rayner Intraocular Lenses, UK) had 5.2% coating from silicone oil (P <0.001 vs. the other lenses).
Vacuoles
Vacuoles in foldable IOL materials, particularly hydrophobic acrylics, are ascribed to water vapor that accumulates in microvoids within the material. The number and size of microvoids are affected by the manufacturing process.
Vacuoles within the IOL have been reported in several studies of patients with AcrySof IOL implants.[29–33] Christiansen et al.[30] reported that 42 eyes implanted with AcrySof IOLs all exhibited some degree of vacuole formation that appeared as glistenings under slit-lamp observation. Laboratory studies suggest that temperature elevations increase the level of glistenings[34] as does the addition of serum to AcrySof IOLs in aqueous humor.[35] The presence of vacuoles in the IOL at mild-to-moderate levels does not appear to affect visual function. However, high levels of vacuole formation can decrease visual acuity,[30,][32] decrease contrast sensitivity,[29,][32,][33] and interfere with the ability to target a neodymium:yttrium-aluminum-garnet (Nd:YAG) laser for posterior capsulotomy.[32]
Optic edge designs
Initially, most PMMA and foldable IOLs had rounded edges. Manufacturers now design optics with sharp or squared edges to inhibit the growth of lens epithelial cells over the posterior capsule. Nishi, Nishi, and Sakanishi[36] reported that an AcrySof IOL with a sharp optic edge had a significantly greater effect in inhibiting posterior capsule opacification (PCO) than a PMMA IOL with rounded edges. Later, Nishi, Nishi, and Wickstrom[37] looked at the influence of IOL design and materials, implanting an AcrySof acrylic IOL in one eye and a CeeOn silicone IOL in the other eye of seven rabbits. Both IOLs had a sharp, rectangular optic edge. Miyake-Apple views showed that migrating lens epithelial cells were inhibited at the optic edge of five of the six rabbits available for evaluation. Overall, there was no apparent difference in PCO development between the two IOLs for the first 3–4 weeks. The study suggests that IOL design is an important factor in preventing PCO (Figures 37-3 and 37-4).
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Figure 37-3 Schematic representation of epithelial cell migration under the rounded edge of an intraocular lens. |
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Figure 37-4 Schematic representation of epithelial cell migration inhibited by the sharp posterior edge of an intraocular lens in contact with the posterior capsule. |
With the square-edge hydrophobic acrylic IOL, some patients have reported bothersome edge glare and negative dysphotopsias.[38–41] Holladay, Lang, and Portney[42] conducted ray tracings of biconvex IOLs that were identical except that the edge was either round or square. The sharp-edge design formed an arclike pattern of reflected light on the retina, and the round-edge design formed a diffuse image (Figure 37-5).
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Figure 37-5 Distribution of internally reflected light on the retina. |
Rounding the biconvex lens reduced the peak intensity of the reflected glare image by 90%. The authors concluded that the glare images from the sharp edge appeared like a thin crescent or partial ring in the periphery of the retina opposite the image of the glare source and that rounding the edges significantly reduced the peak intensity of the reflected glare image.
An alternative explanation for the photic phenomena reported by patients with the square-edge AcrySof IOL comes from a study comparing equiconvex silicone (LI16U) and PMMA (P359UV) IOLs with an unequal biconvex acrylic (MA60BM) IOL.[43] The refractive indices for the three IOLs were 1.43, 1.49, and 1.55, respectively. This study suggested that the unequal biconvex design produced greater postoperative glare and external reflections than the equiconvex IOLs. Further, an increase in the refractive index from 1.43 to 1.55 increased the amount of reflected light fivefold.
Another approach to edge design was taken by the Sensar (acrylic) and ClariFlex (silicone) lenses. The Sensar IOL has a standard rounded edge. The Sensar with OptiEdge and ClariFlex have a sharp vertical edge on the posterior surface, a sloped edge along the side of the IOL, and a rounded anterior edge (Figure 37-6). The ray tracings of this new design show less reflection off the posterior surface than a conventional square-edge design and dispersion of light rays through the anterior surface (Figure 37-7). These lessons from early improvements in reducing photic phenomena continue to be evolved in new IOL edge designs.[44] Some studies find variations in rates of complaints about dysphotopsia based on IOL design[45] while others report that some patients perceive similar dysphotopsias with two markedly different IOL designs.[46]
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Figure 37-6 Edge design of the Sensar with OptiEdge and ClariFlex IOLs. |
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Figure 37-7 Ray tracings of the Sensar with OptiEdge. |
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Aspheric optics
The understanding of sophisticated optics and its relevance to clinical ophthalmology is credited to corneal laser refractive surgery. Vision symptoms that were not consistent with excellent high-contrast visual acuity rapidly led to an understanding of the impact of large-aperture optics and aberrations on the quality of vision. This understanding rapidly found its way into cataract and IOL surgery.
The natural corneal curvature has an asphericity that averages +0.27μ RMS. In youth, this is balanced by a lenticular asphericity averaging −0.27μ RMS, leading to minimal total spherical aberration.
The original conventional IOL designs all had optics designed on elementary optics, which resulted in positive spherical aberration in the IOL. This increased the total spherical aberration of the eye after cataract and IOL surgery, with resultant loss of image quality [47–49,][53,][60,][61] and loss of contrast sensitivity.[50]
IOL designs increasingly address the issue of the spherical aberration of the eye. Aspheric optics clinically improve quality of vision and contrast sensitivity compared to spherical optics.[51–54]
The original IOL design with aspheric optics was the lens that is now known as the AMO Tecnis aspheric series of lenses (e.g., Z9000 series). These IOLs are designed with −0.27μ RMS spherical aberration (Figures 37-8 and 37-9). These IOLs have proved to improve the quality of vision with large pupils, night vision, and night vision with glare.[45,][46]
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Figure 37-8 Total ocular aberrations measured by the Hartman-Shack method in a patient with a spherical optic intraocular lens. Note the amount of spherical aberration in particular (term 4,0 at 0.23μ RMS). |
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Figure 37-9 Total ocular aberrations measured by the Hartman-Shack method in the patient from Figure 37-8 after exchange of the intraocular lens (IOL) for an aspheric optic IOL with negative asphericity of −0.27μ RMS. Note the reductions in aberrations, particularly spherical aberration, which is reduced to −0.06μ RMS, closely matching the theoretical result of 0.23–0.27 = −0.04μ RMS). |
Subsequent aspheric designs have targeted different ranges of spherical aberration. The Bausch & Lomb aspheric design (LI61AO) targets zero asphericity in the IOL, not attempting to alter the corneal asphericity. The Alcon aspheric IOLs (e.g., SN60WF series) are designed for −0.17μ RMS of asphericity, striking a middle ground. Aspheric IOLs must be centered within 0.4mm of the pupil center to results in optical benefit.[57] The lower the negative asphericity, the less benefit to a well-centered IOL but, correspondingly, the less degradation from decentration or tilt.[58]
Another factor in the total visual performance of an IOL is chromatic aberration and its balance with spherical aberration. The interaction of these two factors can have an important impact on depth of focus in aspheric IOLs. An analysis of a variety of commercial IOL models showed marked variation in depth of focus, particularly under low-light conditions, with the best depth of focus achieved by the Tecnis aspheric IOL.[59]
One common situation alters corneal asphericity: corneal refractive surgery. Both radial keratotomy and myopic excimer laser refractive surgery generally increase positive spherical aberration. Wavefront-guided and wavefront-optimized corneal myopic refractive surgery also usually increases spherical aberration, but less than in the case of non-“customized” or “guided” treatment. Hyperopic corneal refarctive surgery, on the other hand, tends to decrease positive spherical aberration. As a rule, therefore, a patient with prior myopic corneal refractive surgery will benefit from an IOL with maximal negative aspheicity. A patient with prior hyperopic corneal refractive surgery may do better with one of the original IOL designs that have positive spherical aberration. If possible, patients with prior corneal refractive surgery should have a preoperative measurement of corneal aberrations to guide the optimal IOL selection.
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Biocompatibility
Foreign-body reaction
Clinical measures of biocompatibility include cellular deposits on the capsule, capsule opacification, postoperative blood–aqueous barrier breakdown, cellular reaction at the anterior capsule–IOL interface, and postoperative inflammation. Amon[60] calls biocompatibility one of the most important prerequisites of an intraocular implant and raises concerns about the use of the term in the medical literature. He quotes the definition from the International Dictionary of Medicine and Biology, “Biocompatibility is the capability of a prosthesis implanted in the body to exist in harmony with tissue without causing deleterious changes.”[61]
Ideally, the IOL should remain an inert refractive element within the intraocular structures, but it does react with different tissues in the eye, specifically the uvea and the capsular bag. Capsular biocompatibility, according to Amon, occurs when there is minimal or no lens epithelial cell proliferation over the posterior or anterior capsule. Similarly, uveal biocompatibility exists when the IOL causes only a very mild foreign-body reaction despite the close proximity or direct contact between the IOL and uveal tissue.
Anterior capsule opacification
Posterior capsule opacification and the role of the IOL are covered in detail in Chapter 51. Less widely recognized and discussed is lens epithelial cell migration onto the anterior surface of the IOL. Hollick, Spalton, and Ursell[62] looked at surface cytologic features of PMMA, silicone, and hydrogel IOLs and their effect on the blood–aqueous barrier as an indicator of biocompatibility. They found that patients who had hydrogel IOLs implanted were significantly more likely to have lens epithelial cells on the anterior surface of the IOL for a longer period than those with PMMA or silicone IOLs (P <0.001). The investigators commented that lens epithelial cells grew over the anterior surface to a much greater extent and in a completely different pattern on the hydrogel IOLs than the PMMA or silicone IOLs.
Müllner-Eidenböck et al.[63] conducted a prospective, randomized clinical study of lens epithelial cell migration onto the anterior capsule in 15 eyes, each implanted with one of four IOLs: the HydroView, MemoryLens, AcrySof, or CeeOn 920. The greatest ongrowth was seen with the hydrophilic acrylic IOLs, HydroView and MemoryLens, which had 86.7 and 73.4%, respectively, at 30 days and 86.7 and 22%, respectively, at 180 days. The lowest rates were seen with the hydrophobic acrylic AcrySof IOL (86.7% at 30 days and 0% at 180 days) and CeeOn 920 silicone IOL (26.7% at 30 days and 19.9% at 180 days).
Lenis and Philipson[64] observed 25 cataract patients with HydroView hydrogel IOL implants for up to 12 months. Lens epithelial cells were detected on the anterior IOL surface in 13 patients (52%). Most of these were within 1mm of the capsulorrhexis border, with four cases of cell proliferation centrally. Overall, this migration and proliferation did not affect visual acuity.
House et al.[65] evaluated 41 cases of AcrySof IOL implantation in 31 patients. They found granular deposits of what they classified as lens epithelial cell proliferation on the anterior surface of the IOL in 18 cases (44%) at 3 to 5 weeks after surgery. The deposits did not have an effect on visual acuity.
Koch, Kalicharan, and van der Want[66] reported on 62 (of 196; 33.2%) eyes implanted with a HydroView IOL that developed a layer of lens epithelial cells on the anterior surface of the IOL optic between 8 and 98 weeks after surgery. The presence of the lens epithelial cell membrane produced visual symptoms varying from low vision to hazy vision or light scatter. Removal of the membrane by Nd:YAG laser treatment or surgical membranectomy alleviated the visual symptoms.
Schauersberger et al.[67] found significantly higher levels of anterior ongrowth of lens epithelial cells on the round-edge HydroView hydrophilic acrylic IOL than the square-edge AcrySof or round-edge Sensar hydrophobic acrylic IOLs from 30 days to 1 year after surgery (P =0.015). They concluded that, unlike in PCO, IOL material plays a greater role than edge design in anterior lens epithelial cell ongrowth.
Inflammatory reaction
The trauma of cataract surgery and IOL implantation can trigger a breakdown in the blood–aqueous barrier, with an outpouring of proteins and macrophages to repair tissue and isolate the foreign body from the ocular tissues. The foreign-body reaction depends on the IOL material and the pre-existing status of the ocular tissues, such as diabetic retinopathy. The inflammatory response is measured subjectively by slit-lamp grading of cells and flare, or objectively by a laser flare photometer.
Hollick, Spalton, and Ursell[62] measured the effect of PMMA, silicone (SI30NB), and hydrogel (HydroView) IOL implantation on the blood–aqueous barrier by laser flare and cell meter (Kowa, Osaka, Japan). At 1 month after surgery, small cells were present on 4 of 30 hydrogel, 11 of 30 PMMA, and 15 of 30 silicone IOLs (P = 0.01). The hydrogel lenses showed a significantly shorter duration and lower grades of small cells than the silicone or PMMA IOLs (P <0.001 for both). At the same visit, epithelioid cells were most pronounced on the PMMA (11 of 30), present in only 1 of 30 silicone IOLs, and not present on the hydrogel IOLs (P <0.001). Lens epithelial cells, on the other hand, were present at 1 month on 20 of 30 hydrogel IOLs, 12 of 30 PMMA IOLs, and 0 of 30 silicone IOLs. Lens epithelial cells reached a peak between 1 week and 1 month after surgery and then regressed. The hydrogel IOLs had significantly greater numbers that did not regress, and in half the patients the lens epithelial cells formed a confluent sheet along the entire 360° of the capsulorrhexis rim, probably because of migration from the capsulorrhexis.
In a study by Samuelson, Chu, and Kreiger[68] the levels of postoperative inflammatory giant-cell deposits were measured in patients 6 months after combined cataract and glaucoma surgery. The level of inflammatory giant-cell deposits was slightly higher in patients with an acrylic IOL (AcrySof) than in patients with a silicone IOL (PhacoFlex II); however, the difference was not statistically or clinically significant.
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References
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