Purpose
Flanged intrascleral haptic fixation (FISHF) is a useful method for securing intraocular lenses (IOLs) in eyes without capsular support. Biomechanical studies were conducted to support the use of this technique.
Design
Laboratory investigation.
Methods
Haptics of 3-piece IOLs were passed through cadaveric human sclera using 30- and 27-gauge needles. Flanges were created by melting 1.0 mm from the haptic ends using cautery. The forces required to remove the flanged haptic from the sclera and disinsert the haptic from the optic were measured using a mechanical tester and a custom-fabricated mount.
Results
The mean FISHF dislocation force using 30-gauge needles was greatest with the CT Lucia 602 (2.04 ± 0.24 newtons [N]) compared to the LI61AO (0.93 ± 0.41 N; P = .001), ZA9003 (0.70 ± 0.34 N; P = <.001), and MA60AC (0.27 ± 0.19 N; P <.001). Using 27-gauge needles with the CT Lucia resulted in a lower dislocation force (0.56 ± 0.36 N; P <.001). The FISHF dislocation force was correlated with the flange-to-needle diameter ratio (r = 0.975). The FISHF dislocation forces of the CT Lucia and LI61AO using 30-gauge needles were not significantly different from their haptic-optic disinsertion forces ( P = .79 and .27, respectively). There were no differences in flange diameters between 1.0 mm and 2.0 mm haptic melt lengths across the IOLs ( P = .15–.85).
Conclusions
These data strongly support the biomechanical stability of FISHF with the polyvinylidene fluoride haptics of the CT Lucia using small diameter instruments for the creation of an intrascleral tunnel. 1.0 mm of haptic may be the optimal melt length.
S S cleral fixation is an important strategy for secondary intraocular lens (IOL) implantation in eyes without capsular support. This method allows the IOL to be positioned in its physiologic location without the need for consideration of iris or capsular anatomy and may pose a lower risk for corneal decompensation, inflammation, and secondary glaucoma than anterior chamber IOLs. To achieve scleral fixation, suture techniques can be effective, but their use is limited by long-term complications including suture erosion and breakage.
Sutureless techniques in scleral fixation have quickly evolved and have grown in popularity. , Maggi and Maggi reported the first sutureless scleral fixation technique using 8-0 polypropylene sutures attached to a specialized IOL; sutures were subsequently externalized and cauterized to secure the IOL. However, it was not until a decade later that modern sutureless techniques gained traction. Gabor and Pavlidis described transscleral fixation of a standard 3-piece IOL using two 24-gauge ab externo straight sclerotomies with adjacent scleral tunnels.Agarwal and associates also introduced a transscleral fixation technique using scleral flaps and fibrin glue. More recently, Yamane and associates described a transconjunctival flanged intrascleral haptic fixation (FISHF) technique, which has since become widely adopted. The originally described FISHF technique used two 30-gauge TSK (TSK Laboratory, Tochigi, Japan) ultra-thin wall needles to create angled limbal sclerotomies 180-degrees apart, through which the haptics of a 3-piece IOL were externalized using a double-needle technique. Cautery was then applied to the haptic tip creating a flange, which was inserted into the scleral tunnel, preventing the haptic from dislocating back into the eye. Since the initial article by Yamane and associates, there have been dozens of modifications to the technique, including the use of 27-gauge trocars for both vitrectomy and intrascleral haptic fixation. However, there is a lack of data available to guide nuances of the technique such as the amount of haptic to melt, the effect of the intrascleral tunnel diameter on fixation stability, or the IOL models to use for the most favorable long-term stability.
Published outcome data for FISHF of secondary IOLs suggest a dislocation rate ranging from 0% to as high as 29% in a recent series performed in diabetic patients; however, there were differences in the techniques used, and limited data are available for eyes beyond 3 years. To better ascertain the stability of this technique, more long-term data are needed. Furthermore, biomechanical studies can help to assess the stability of the FISHF technique, but there is also currently a lack of data in the medical literature.
To better understand the biomechanical strength of FISHF in secondary IOLs, this study evaluated the force needed to dislocate haptics secured with this technique from human cadaveric sclera and compared it to the haptic-optic disinsertion force of 4 commonly used IOLs in the United States. The diameters of unmanipulated haptics, flanged haptics, and external hypodermic needle diameters were also measured and were correlated with the FISHF dislocation forces of each IOL model.
METHODS
This was an experimental study of the following four 3-piece intraocular lens models that were widely available in the United States: the CT Lucia 602 (Carl Zeiss Meditec, Dublin, California, USA); Tecnis ZA9003 (Johnson & Johnson, Santa Ana, California, USA); AcrySof MA60AC (Alcon Laboratories, Fort Worth, Texas, USA); and SofPort LI61AO (Bausch & Lomb, Rochester, New York, USA).
The unmanipulated cross-sectional haptic diameter of each intraocular lens was measured using a digital microscope with a magnification of 140x (model AM73915MZTL 5MP, Dino-Lite Edge, Torrance, California, USA). A thin-tip marker was used to place markings denoting 0.5, 1.0, and 2.0 mm from the haptic end under the microscope. Flanges measuring 0.5, 1.0, and 2.0 mm were created using a cordless fine-tip cautery device (MediChoice, Owens & Minor, Mechanicsville, Virginia, USA), which was applied head-to-head, parallel with the longitudinal axis of the haptic, without grasping the haptics with forceps. The cross-sectional flange diameters were measured using the same digital microscope. The external and internal diameters of 27-gauge BD needles (Becton, Dickinson, Franklin Lakes, New Jersey, USA) and 30-gauge TSK ultra-thin wall needles were also measured using the digital microscope.
Human cadaveric scleral tissue from 2 donor eyes (CorneaGen, Waltham, Massachusetts, USA), with death-to-preservation times of less than 15 hours, were used within 9 days post mortem. The IOL haptic was inserted into the cadaveric human sclera approximately 1.5 mm posterior to the limbus through a 1.0–1.5-mm intrascleral tunnel using 30-gauge TSK needles (for every IOL model) or 27-gauge BD needles (CT Lucia 602 only). The scleral fixation sites were spaced circumferentially by at least 2 mm. Flanges were created on the external scleral surface by melting 1.0 mm of haptic end using the same technique as described above and were gently positioned to assume an intrascleral position.
Scleral tissue with a preplaced intrascleral fixated haptic flange was positioned on the surface of a custom-fabricated mount that allowed for the fixated IOL to hang freely below ( Figure 1 , A). The optic underneath was grasped using a wedge grip so that the force applied was perpendicular to the sclera. With a motorized single-column tensile testing machine (model ESM 303; Mark-10, Copiague, New York, USA), using a constant cross-head speed of 50 mm/min, the force required to remove the flanged haptic from the sclera was determined. Measurements with each IOL model using 30-gauge TSK needles and the CT Lucia 602 using both 30-gauge TSK and 27-gauge BD needles were repeated 5 times and averaged. To minimize the effect from possible variations in scleral tissue properties, each permutation of haptic and needle type was tested once in a sequential manner circumferentially and then identically repeated.
The force required to cause haptic-optic disinsertion was also measured for each type of IOL using the same mechanical tester. The IOL optic was positioned on the custom-made mount, and the haptic was secured using a wedge grip from below ( Figure 1 , B). Measurements were repeated 5 times for each IOL model and averaged.
Unpaired 2-tailed Student t -tests were used to evaluate differences among variables. A P value of <.05 was considered statistically significant. The flange-to-needle diameter ratio and the flange-to-haptic diameter ratio were calculated by dividing the flange diameter by the external needle diameter or the haptic diameter for each pairing, respectively. The correlation coefficients between the flange-to-needle diameter ratio or the flange-to-haptic diameter ratio and flange dislocation force were computed.
RESULTS
The diameters of the haptics and flanges created after melting 0.5, 1.0, and 2.0 mm of haptic material for each IOL model are reported in Table 1 . The flange-to-haptic diameter ratio ranged from 1.59 to 3.04 across the different IOL models and amount of haptics melted. There were no significant differences in flange diameter between the 1.0- and 2.0-mm flanges for any IOL model ( P = .145–.848). The flanges of the different IOL models exhibited different shapes, for example, the CT Lucia 602 haptic flange resembled a mushroom ( Figure 2 , A), the ZA9003 and LI61AO flanges resembled tapered cones ( Figure 2 , B and C), and the MA60AC flange resembled a bulb ( Figure 2 , D).
| Intraocular Lens Model | Haptic Material | Mean ± SD Haptic (mm) | Mean ± SD 0.5- mm Flange (mm) | Mean ± SD 1.0-mm Flange (mm) | Mean ± SD 2.0-mm Flange (mm) | 1.0-mm vs. 2.0- mm Flange ( P value) |
|---|---|---|---|---|---|---|
| CT Lucia 602 | PVDF | 0.147 ± 0.005 | 0.377 ± 0.008 | 0.445 ± 0.015 | 0.431 ± 0.037 | .458 |
| ZA9003 | PMMA | 0.146 ± 0.004 | 0.333 ± 0.035 | 0.365 ± 0.043 | 0.363 ± 0.023 | .848 |
| LI61AO | PMMA | 0.146 ± 0.003 | 0.232 ± 0.023 | 0.338 ± 0.023 | 0.370 ± 0.037 | .145 |
| MA60AC | PMMA | 0.150 ± 0.005 | 0.291 ± 0.000 | 0.306 ± 0.033 | 0.358 ± 0.095 | .226 |
The external diameters of the 27-gauge BD needles and the 30-gauge TSK needles were 0.424 ± 0.003 mm and 0.321 ± 0.001 mm, respectively. The internal diameters of the 27-gauge BD needles and the 30-gauge TSK needles were 0.224 ± 0.023 mm and 0.192 ± 0.008 mm, respectively.
The mean FISHF dislocation force using 30-gauge TSK needles was greatest with the CT Lucia 602 (2.04 ± 0.24 newtons [N]) compared to the LI61AO (0.93 ± 0.41 N; P = .001), the ZA9003 (0.70 ± 0.34 N; P = .001), and the MA60AC (0.27 ± 0.19 N; P <.001) ( Tables 2 and 3 , Figure 3 ). The FISHF dislocation force for the CT Lucia 602 using 27-gauge BD needles (0.56 ± 0.36 N) was less than that using 30-gauge TSK needles ( P <.001). The haptic-optic disinsertion forces were 1.99 ± 0.29 N (CT Lucia 602), 1.21 ± 0.06 N (ZA9003), 1.09 ± 0.13 N (LI61AO), and 1.03 ± 0.14 N (MA60AC) ( Table 2 ). The FISHF dislocation forces of the CT Lucia 602 and LI61AOusing 30-gauge needles were not significantly different from the haptic-optic disinsertion force of the same IOL ( P = .79 and .27, respectively) ( Table 2 , Figure 3 ). The flange-to-needle diameter ratio was between 0.95 and 1.39 and was greatest for the CT Lucia 602 paired with the 30-gauge TSK needle ( Table 2 ). A greater FISHF dislocation force was correlated with a greater flange-to-needle diameter ratio (r = 0.975) and a greater flange-to-haptic diameter ratio (r = 0.718) ( Figure 4 ).
