Comparison of IOL Materials

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 by Derek W DelMonte, MD on May 27, 2023.


Intraocular lenses (IOLs) are commonly implanted after surgical removal of a cataractous lens. Since the introduction of the first modern IOL in 1949, IOL materials and designs have evolved to encompass medical needs. A variety of IOL materials are currently available, including collamer, hydrophobic acrylic, hydrophilic acrylic, PEG-PEA/HEMA/Styrene copolymer, polymethylmethacrylate (PMMA), and silicone . While PMMA was the first material to be used in IOL implantation, technological advances have brought about silicone and acrylic lenses, which are foldable and therefore, more applicable with the advent of small-incision, phacoemulsification-assisted cataract surgery. These IOL materials contain distinct physical and optical properties, with unique advantages and disadvantages.

IOL Properties

IOL materials can be classified using various properties they possess including abbe number, biocompatibility, hydrophobicity, hygroscopy, glass transition temperature, and refractive index. These properties contribute to visual quality.

Abbe Number

The Abbe number is a measure of an optical material's chromatic dispersion, with higher values representing lower chromatic dispersion, and better optical quality [1]. As the refractive index increases, the Abbe number decreases [2]. The Abbe number of the natural crystalline lens is 47 [1]. The human visual system exhibits spherical refractive errors of 2.5D within the visible spectrum; however, the phakic human eye is well adapted to this longitudinal chromatic aberration (LCA) due to compensatory mechanisms, including blue-light filtering capabilities of the lens and macula [3][4]. Moreover, chromatic dispersion does not vary significantly from one person's eye to another [5]. In pseduophakic eyes, however, various IOL materials determine LCA, with Abbe numbers ranging from 37-55 [5]. Accommodation, contrast sensitivity, emmetropization, and visual acuity are all affected by chromatic aberration [4].


Biocompatibility is defined as the capability of an implanted IOL material to remain optically functional in vivo without causing significant adverse responses such as anaphylaxis, inflammation, necrosis, rejection, thrombosis, or tumorigenesis [6]. IOL materials must also block ultraviolet rays, be optically transparent, and have a high refractive index[7]. Capsular biocompatibility is determined by direct contact of the IOL with remnant lens epithelial cells (LECs), which may lead to adherence and anterior capsule opacification (ACO) and posterior capsule opacification (PCO)[8]. Uveal biocompatibility refers to how well an IOL is tolerated in the eye without causing immunogenic responses of the anterior choroid, ciliary body, and iris [8].


An IOL material surface may be hydrophilic or hydrophobic, measured by photographing a water droplet placed onto the material surface. The contact angle measures the degrees between the material’s surface and water droplet’s surface[7]. Large contact angles define hydrophobic materials, whereas smaller angles define hydrophilic materials[7].


Hygroscopy is the ability of a material to absorb and hold water (also known as a lens material's equilibrium water content) [7]. Increasing a material’s hygroscopy reduces the common postoperative appearance of glistenings, which are small fluid-filled microvaculoes that appear after IOL implantation. This is likely because once water enters the material, it interacts with hydrophilic groups and therefore does not collect in microvacuoles [9].

Glass transition temperature

The fluidity of polymers changes with temperature. A polymer appears solid at lower temperatures because its chains move slowly, whereas it transitions to a viscous material at higher temperatures due to increased energy and molecular mobility [7]. The glass transition temperature of a polymer is the temperature at which the polymer changes from a rigid, solid-like glassy fluid to a softer, rubber-like fluid [7]. IOLs are manufactured with a goal glass transition temperature below the physiologic body temperature of 37°C and normal room temperature because materials with a glass transition temperature above 37°C or normal room temperature would not unfold at physiologic body temperatures or at room temperature [6].

Refractive index

The refractive index is a measure of the bending of a ray of light, calculated by the ratio of the velocity of light within a vacuum to the velocity of light within the material [10]. The refractive index of the natural crystalline lens is 1.4 [1]. The refractive index increases with the addition of halogens, aromatic groups, or sulfur [2]. An inverse relationship exists between the refractive index and IOL thickness [10].

IOL Materials


A collamer is a collagen polymer. Collamer IOLs are widely used foldable IOL, designed for placement in the posterior chamber as a phakic IOL (e.g. Visian implantable collamer lens or ICL). Because of its posterior chamber location, the success of collamer IOL depends upon the central vault, which is defined as the distance between the posterior surface of the IOL and the anterior surface of the crystalline lens [11]. The desired central vault value lies between 0.25 to 0.75 mm [11] [12]. Vault values outside of this range may lead to detrimental post-operative effects. For instance, a low vault value may induce anterior subcapsular cataract formation due to mechanical contact between the IOL and the crystalline lens as well as disrupting the physiologic circulation of aqueous humor; conversely, a high vault value increases the risk of angle closure glaucoma secondary to chronic pigment dispersion from contact of the collamer IOL on the posterior surface of the iris [11] [12][13]. Patient age, anterior chamber depth, sulcus-to-sulcus distance, white-to-white distance, and IOL size are additional factors related to the vault [12]. Benefits of ICL compared to laser corneal refractive surgery, include ability to correct higher degrees of myopia (-3 - -15D), higher degrees of astigmatism (≤2.5D), reducing very high myopia (-15- -20D), and can be implanted in eyes with thinner corneas on pachymetry. Requirements per FDA for implantation of an ICL: minimum anterior chamber depth of ≥2.8 mm for myopia and ≥3.0mm for hyperopia, irido-corneal angle ≥30°, corneal endothelial cell count >2500 cells/mm2 if patient age >21 years or >2000 if patient age > 40 years., and lastly mesopic pupil size <5.0-6.0 [14]. There are different ectasia risk assessment scoring systems to stratify risk of post-corneal refractive surgery ectasias, given specific clinical factors (e.g. corneal topography pattern, residual stromal bed thickness, patient age at time of surgery, corneal thickness, preoperative spherical equivalent manifest refraction) [15]. Adverse effects with phakic IOL implanted in the posterior chamber include increased incident of cataract formation, residual astigmatism, chronic uveitis, corneal endothelial cell loss, endophthalmitis, and pupil ovalization [13].

Hydrophobic Acrylic

Introduced in 1993, hydrophobic acrylic IOLs such as AcrySof are the most widely used IOL material currently. They are composed of crosslinked copolymers of acrylic esters and other acrylic ester co-monomers, with a carbon backbone and ester side groups [6]. Single and three-piece lenses can be folded and implanted through small incisions while maintaining their original shape[7][16]. Long-term studies have demonstrated that patients implanted with hydrophobic acrylic lenses have lower rates of, and less dense, PCO, and were less likely to require Nd-YAG capsulotomy than patients implanted with hydrophilic acrylic IOLs [17]. This may be due to the tendency of hydrophobic acrylic IOLs to adhere to the posterior lens capsule through fibronectin bindings, leading to decreased space for LEC migration to occur in between the IOL and posterior capsule [17] [18]. Unfortunately, glistenings commonly affect hydrophobic acrylic IOLs [19].

Hydrophilic Acrylic

The new generation of foldable IOLs are composed of the methacrylate backbone of PMMA with additional hydroxyl groups introduced in the side chains [7]. The addition of hydroxyethylmethacrylate (HEMA), a material used in the manufacturing of contact lenses, poly(2-HEMA), or poly-HEMA confers flexibility to the IOL [7][10]. With the development of small-incision phacoemulsification assisted cataract surgery, foldable IOLs are the main type of lens used. Since these lenses are highly flexible, they require an incision of approximately 1.8mm, thereby making them highly beneficial in microincision cataract surgery [20].

Hydrogel lenses, as in those with poly-HEMA chains [10], have varying optical properties depending on their water content [7]. Once the manufactured lens, which is dry and opaque, is placed in water, it becomes soft and optically clear [7]. Hydrophilic IOLs have a high water content of 18-34%, leading to superior biocompatibility with lower rates of glare, a lower refractive index of 1.40-1.43 and an increased IOL thickness (Table 1) [7]. In a prospective study of 86 eyes with pseudoexfoliation syndrome (PEX) comparing hydrophilic acrylic, hydrophobic acrylic and silicone IOLs, hydrophilic acrylic IOLs had superior capsular biocompatibility since lens epithelial cell outgrowth was lowest with this group [21]. However, hydrophilic acrylic IOL had poor uveal biocompatibility since this group had the highest deposition of debris on the IOL surface after 12 to 18 months, and the highest rate of PCO [21]. This may be because the swelling that occurs within the lens after manufacturing induces difficulty maintaining a sharp posterior edge [7]. Although older generations of hydrophilic acrylic IOLs were associated with primary internal or surface calcifications due to calcium and phosphoric acid deposition, this does not seem to be a problem in newer generations [7][20].

PEG-PHS/HEMA/Styrene (PHS) Copolymer

The appeal of hydrophilic IOL materials incorporating HEMA derives from its biocompatibility, flexibility for microincisions, and lower rates of glare, but is set back by the increased incidence of PCO and calcification [22].  In contrast, the lower rates of PCO occurrence and surgical complications are more advantageous with hydrophobic acrylic IOL material yet are more likely to form glistenings in the long term [21] [22]. As lenses with low hygroscopy, 0.1% - 0.5%, are more likely to develop glistenings [9]. Companies have sought to maximize the balance in incorporating both hydrophilic and hydrophobic materials to further advance IOL material configuration. For instance, enVista’s MX60 IOL is composed of a copolymer mix of 40% hydrophobic poly(ethylene glycol)-phenyl ether acrylate (PEG-PEA), 30% of hydrophilic HEMA, 26%  of styrene to add bulk and increase the refractive index, and 4% of ethylene glycol dimethacrylate (EG-DMA) for structural integrity [7]. The main composition of enVista MX60L IOL, PEG-PEA, HEMA, and styrene (PHS Copolymer), is a hydrophobic IOL material with significant hydrophilic content that holds a hardness of 1.8MPa and a hygroscopy of 4% [7]. This is partly attributed to the PEG-PEA repeating subunit of ethylene glycol allowing a carbon chain of flexibility and hydrophobicity and HEMA’s terminal hydrophilic hydroxyl group increasing the material’s hygroscopy. Studies have shown that usage of enVista MX60L in cataract surgery demonstrates a good safety profile, stable refractive outcomes, low incidences of PCO and incidences of Nd:YAG capsulotomies, and no incidences of glistenings [23] [24].


PMMA, a non-foldable, rigid, hydrophobic material, is available in one-piece version and three-piece versions that is lathe cut from PMMA buttons or rods [6]. PMMA, the first IOL material implanted into a human by Harold Ridley in 1949, is known for its excellent tissue tolerance, long-term stability, and reduced cost [16]. It has a contact angle of 65° to 71°, hygroscopy percentage of 0.4% to 0.8%, glass transition temperature of 105°C to 113° C, and a high refractive index of 1.49 (Table 1). Furthermore, PMMA IOLs have superior optical clarity, allowing for a broad spectrum of light to be transmitted [16].

Despite these properties, the PMMA IOL has been replaced by modern foldable lenses. PMMA's rigid structure prevents it from being folded to enter the small 2 mm incision made for phacoemulsification; a wider incision of approximately 5.5mm to 6mm is required for implantation, which resulting in delayed healing, poor intraocular pressures, and postoperative astigmatism [10]. A well-known long-term complication of PMMA IOLs is snowflake degeneration, which is an indication of PMMA IOL explantation[25]. PMMA's hydrophobicity may be damaging to the corneal endothelium secondary to adherence to these cells during implantation.[10] PMMA and hydrogel IOLs are associated with higher rates of PCO compared to acrylic and silicone IOLs [26][27]. While this may in part be due to IOL surface roughness providing a scaffold for PCO tissue formation, recent studies suggest that that IOL edges may play a more significant role; long-term PCO and neodymium-doped ytrium aluminum garnet (Nd:YAG) capsulotomy rates between square-edge PMMA and round-edge PMMA IOLs found that a squared posterior optic edge reduced proliferation of lens epithelial cell on the posterior lens capsule, thereby reducing PCO rates [26]. PMMA IOLs are currently used as scleral-sutured IOLs for instance when capsular implantation is not feasible [28].


A synthetic polymer composed of a backbone of repeating silicon-oxygen groups, silicone is used to make foldable IOLs [10]. Silicone lenses are made with plate haptics or modified C-loop haptics. Although plate haptics have lower PC and cystoid macula edema rates, an intact anterior capsule is required to prevent decentering [10].

Silicone has the desirable properties of flexibility, biocompatibility, autoclavability, relative inertness, and chemical stability [29]. Silicone is hydrophobic, with a contact angle of 97° to 120°, hygroscopy percentage of 0.38%, and glass transition temperature of -120°C to -90°C (Table 1). Since silicone lenses have a refractive index of 1.43, which is lower than that of acrylic lenses, they are thicker for the same refractive power - this thickness may require an incision of up to 3.2 mm for implantation [10]. A review article by Kwon et al. demonstrated that silicone IOLs had lower PCO and Nd:YAG capsulotomy rates than hydrophobic acrylic IOLs at long-term use (>6 years) following implantation [30]. This may be because LECs are less likely to adhere to the silicone IOL surface, and because silicone IOLs have the squarest edges [7].

However, silicone is not widely used as an IOL biomaterial due to its susceptibility to adherence of bacteria, cells, and silicone oils [29][31]. This may lead to endophthalmitis, capsular opacification, IOL decentration, and capsular contraction, thereby impacting patients' post-operative visual acuity [31]. Silicone IOLs lead to premature, uncontrolled opening of the IOL within the anterior chamber, which can make IOL implantation more complex; furthermore, silicone IOLs may fog with condensation during subsequent vitrectomies and thus should be used in caution in diabetic eyes [6][16]. Deposition of calcium oxalate on the silicone IOL has been described in eyes with asteroid hyalosis, following YAG capsultomy [32].

Table 1

IOL material IOL Design Implanted Location Sphericity Status Hygroscopy (%) Glass Transition Temperature (°C) Refractive Index (n) Complication
Collamer 1-piece Sulcus Negative 40 40 1.44
Hydrophobic acrylic 1-piece Capsular bag Spheric 0.1-0.5 16 - 55 1.47-1.56 Glistenings
Hydrophilic acrylic 1-piece Capsular bag Neutral 18-38 10 - 20 1.40-1.43
PEG-PEA/HEMA/Styrene copolymer 1-piece Capsular bag Neutral 4-5 28 1.54
PMMA 1-piece Anterior chamber Negative 0.4-0.8 105 - 113 1.49 Snowflake degeneration[25]
Silicone 3-piece Capsular bag Neutral 0.38 -90 - -120 1.43 Deposition of calcium oxalate from asteroid hyalosis following YAG capsulotomy[32]
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