In recent years, lens surgery has become a surgical procedure that seeks to not only replace the cataractous lens, but also to correct any refractive error and presbyopia. Intraocular lenses (IOLs) are no longer limited to monofocal models; on the contrary, they include refractive, diffractive, and extended range design lenses, among others.1–3 These lenses are either toric or have rotational symmetry optimum centration and stability within the capsular bag is required to provide the best possible visual outcomes.4–6
The instability of implanted IOLs has been associated with several factors, including the capsulorhexis size,7 incomplete viscoelastic clearance, IOL diameter versus capsular bag diameter,8,9 IOL material properties,10,11 and haptic designs.12–14
Some studies15–17 have demonstrated that IOLs with C-loop haptics ensure better stability and centration than IOLs with plate haptics. Poyales et al18 compared the rotational stability and centration of two different IOLs, the four-loop IOL (Micro F FineVision model) and the double C-loop IOL (POD model), finding a good rotational stability in both lenses. However, it is not only haptic design that influences stability. Tolerance to compressive forces—to which both the haptics and the optics may be subject, once they have been implanted in the capsular bag—also seem to have an impact on IOL stability.13–15 Several clinical studies19–21 have measured the decentration, tilt, or rotation of IOLs after cataract surgery using different measuring systems, such as Scheimpflug imaging,20 measurements using Purkinje reflections,22 and slit-lamp assessment.21 However, the heterogeneous methods, the accurate marking of the reference axis, and the accurate measurements of preoperative parameters makes comparisons difficult. For this reason, some numerical10 and experimental12,13,15,23 studies have evaluated the biomechanical stability of IOL designs before inserting the lenses in the eye, according to the requirements specified by the International Organization for Standardization (ISO 11979-3).24 In a previous study, Remón et al10 used finite element modeling (FEM) to evaluate the biomechanical stability of four different hydrophobic and hydrophilic IOLs with different haptic designs. The results suggest that FEM can be a powerful tool for increasing the predictability of the biomechanical stability of IOLs before their implantation in the capsular bag.
The goal of the current study was to assess the bio-mechanical stability of three different marketed IOLs with different haptic designs (four-loop IOL [Micro F FineVision model] and double C-loop IOL [POD F and POD FT models] manufactured by PhysIOL), using FEM to reproduce the compression test defined in ISO 11979-3.24 In addition, the stability of these lenses was evaluated in vivo once they had been surgically implanted in a patient to obtain the maximum possible information on the stability of the evaluated lenses. Both studies are complementary.
Materials and Methods
All IOLs evaluated in this study are aspheric trifocal diffractive lenses manufactured by PhysIOL. The optics of all three models are similar, combining two diffractive structures that fit together to offer a +3.50 diopters (D) addition for near vision and a +1.75 D addition for intermediate vision.
More specifically, the Micro F FineVision model is a four-loop IOL with 5° angulation made of 25% hydrophilic material. The optic body diameter is 6.15 mm and the overall diameter is 10.75 mm. The POD F and POD FT models are double C-loop IOLs with 5° angulation made of 26% hydrophilic material. The optic body diameter is 6 mm and the overall diameter is 11.40 mm. The POD FT IOL, aside from its toric design, differs from its non-toric counterpart (POD F) in the hinge section at the haptic–optic junction because the POD FT IOL has been slightly widened.
The differentiating factor between these three models is the haptic design, which is a key element to ensure IOL stability.
In Silico Evaluation
A numerical simulation of the biomechanical behavior of the three IOLs during a compression test was performed by FEM using Abaqus 6.14 software (Abaqus, Inc) following the procedure described in ISO 11979-3,24 which establishes a compression of the IOL haptics up to 10 mm for IOLs intended for capsular bag placement. In this compression test, the IOL is placed between two clamps (with a curvature radius of 5 mm) and compressed to measure its mechanical stability. Initially, the clamps are separated by a distance equal to the overall dimension of the IOL. Then, the right clamp is displaced to different compression diameters: 11, 10.5, and 10 mm (the value specified by the ISO standard) and 9.5 mm while the left clamp remains fixed. In the final or deformed configuration (Figure A, available in the online version of this article), the following variables were measured: the compression force measured at the horizontal plane, the axial shift measured at the vertical plane (ie, axial displacement), the tilt in both x- and y-directions, and the rotation, following the procedure described in a previous article.25 The rotation is the resulting angle between an initial vector, with origin in the center and an extreme point on an edge of the optic, and that vector in a determined time. These variables were evaluated in the reference and the final configurations, as well as in the intermediate positions throughout the test. The axial displacement, the tilt, and the rotation were evaluated by comparing the reference and final configuration of the lens. Additionally, the optic decentration was estimated as a distance between the optic center of the IOL and the geometrical center of the two clamps at the final configuration.
Final configuration for the in silico compression test: (A) Distance between clamps equal to 10 mm, following the procedure described in International Organization for Standardization (ISO) 11979-3 and for diameters of 11, 10.5, and 9.5 mm, and (B) representation for measurement of axial displacement in compression.
The properties of the material and the geometry of the IOLs were introduced as key parameters in the software. A detailed description of the numerical model used for the mechanical characterization of the material of the IOLs can be found in a previous article.10
In Vivo Study
The second part of this study was an in vivo randomized controlled trial comparing the lateral decentration and rotation of the three diffractive trifocal lens models evaluated, two of which were spherical and one of which was a toric lens.
A total of 45 eyes from 45 patients were selected, 15 eyes for each IOL model under assessment. The study was conducted in accordance with the tenets of the Declaration of Helsinki. The study was approved by the Ethical Committee of Hospital Clínico San Carlos, Madrid, Spain, and all participants signed the informed consent.
The eyes included in the study had an axial length within the normal range, to try to exclude potential rotations due to the capsular bag being too big or a slight IOL tilt due to a small capsular bag size. Inclusion criteria were no prior eye surgery and no history of eye trauma. In addition, corneal astigmatism had to be less than 1.00 D for those patients receiving implanted spherical IOLs. Regarding IOL assignment methodology, patients who had corneal astigmatism below 1.00 D were randomly assigned one of the two spherical models (POD F or Micro F) under study, whereas patients with higher astigmatism were always implanted with the toric model (POD FT). The selected patients underwent crystalline lens surgery involving IOL implantation in the capsular bag. In all cases, the IOL power calculation method used was the Barrett or the Toric Barrett formula. The formula's constants were customized: A-constant was set to be 118.78 for the Micro F model and 119.02 for the POD F and POD FT models.
All patients had follow-up visits 1 day, 1 week, 1 month, and 3 months postoperatively.
Lateral decentration along the x- and y-directions and rotation were quantified by the Piolet software.21 Immediately after surgery, photographs were taken with a slit lamp using retroillumination and additional external lighting so both the lens marks and the conjunctival vessels could be seen clearly and sharply (day 0). During the 1-day and 3-month follow-up visits, the pupil was dilated and the same procedure was repeated.
To be able to quantify IOL decentration, the images recorded immediately after IOL implantation were used as baseline (reference) data.
All surgeries were performed by the same experienced surgeon (FP) under topical anesthesia. For the cataract procedure, a continuous curvilinear capsulorhexis measuring 5.5 mm in diameter and lens fragmentation were performed with the Catalys laser (Johnson & Johnson Vision Care, Inc). A 2.2-mm, 45° angled, bevel-up surgical knife (Xstar Safety Slit Knife; Beaver-Visitec International) was used to create a self-sealing clear corneal incision. The selected IOL was then implanted in the capsular bag with a single-use injection system (Accujet; Medicel AG) and subsequently positioned guided by a computer-assisted cataract surgery system (Callisto Eye, Zeiss Cataract Suite Markerless; Carl Zeiss Meditec AG). In all cases, once the IOL insertion was completed, all traces of ophthalmic viscosurgical devices were removed.
For all quantitative variables, summary tables were created containing mean and standard deviation and maximum and minimum values. The R code (R Project) was used to determine differences between groups and between preoperative and postoperative evaluations. Due to the small sample size, the robust non-parametric Wilcoxon ranked-sum test was employed. For all tests, the threshold for statistical significance was assumed to be a P value of less than .05.
In Silico Evaluation of Compression Test
Figure B (available in the online version of this article) shows the evolution of the different variables (compression force, axial shift measured at the vertical plane [ie, axial displacement], tilt, and rotation as a function of the clamp displacement). Table 1 shows the results at the final configuration (compression diameter of 11, 10.5, 10, and 9.5 mm) for each IOL model. For the Micro F IOL, the maximum compression diameter was 9.8 mm; from this value, this lens loses stability. As can be observed in the figure, the compression force has an almost linear behavior for all models. The compression force for the POD F IOL was slightly lower than for the POD FT and Micro F IOLs for all compression diameters. The compression force variation, when the haptic was compressed to 10 mm, ranged from 30.009 mg for the POD F IOL to 120.230 mg for the Micro F IOL. This double C-loop design has a moderate haptic compression force, which contributes to the anteroposterior stability of the lens. The axial displacement was maximum for the POD FT IOL, obtaining a value of 0.703 mm for a compression diameter of 10 mm. For the other models, the axial displacement was insignificant for all compression diameters. The tilt, rotation, and lateral decentration were substantially lower than the tolerance limits considered acceptable in the ISO 11979-2 (Table 1).26 The Micro F IOL had the highest value of optic tilt at 9.8 mm compression (0.785°). In terms of rotation, all models were rotationally stable. The POD F IOL had the maximum magnitude of rotation at all compression diameters, varying from 0.375° to 0.256°. The mean optic decentration for all models at the different compression diameters was 0.084 ± 0.004 mm. In general, the values obtained for all variables for each model at the lowest test diameter of 9.5 mm were slightly higher than those obtained with the largest test diameter. The finite element analysis provides other variables of interest, such as stress, strains, and strain energy.
Evolution of the different variables analyzed as a function of the clamp diameter measured at in silico condition: (A) compression force, (B) axial displacement, (C) optic tilt, and (D) rotation. The different variables were measured until the haptics compressed to a diameter of 9.5 mm, following the procedure described in International Organization for Standardization (ISO) 11979-3. All intraocular lenses are manufactured by PhysIOL.
Compression Force, Axial Displacement, Tilt, Rotation, and Lateral Decentration at Different Compression Diameters for All IOLs Measured at In Silico Condition
Figure 1 shows the internal energy of the lens, mostly due to the stored strain energy. The smaller this magnitude, the better the biomechanical stability of the lens. As can be observed in the figure, the internal energy has an almost linear behavior for the POD F and POD FT IOLs, being higher for the latter. However, the internal energy for the Micro F IOL increases exponentially, which is why this lens has substantial instability.
Evolution of the strain energy (nJ) as a function of the clamp diameter measured at in silico condition. All intraocular lenses are manufactured by PhysIOL.
Figure 2 shows the maximum principal stress for each IOL model at 10 mm compression diameter. The POD FT model showed higher values of stress in the haptic–optic junction than the other models.
Maximum principal stress (MPa) at 10-mm compression diameter measured at in silico condition. All intraocular lenses are manufactured by PhysIOL.
In Vivo Study
The mean axial length values were 23.34 ± 0.97 mm (range: 21.19 to 25.57 mm) for the Micro F IOL group, 23.36 ± 0.90 mm (range: 21.32 to 25.31 mm) for the POD F IOL group, and 24.07 ± 0.72 mm (range: 22.11 to 27.54 mm) for the POD FT IOL group. The mean power of the implanted IOLs was 21.46 ± 2.94 D (range: 16.50 to 27.00 D), 21.46 ± 3.19 D (range: 15.00 to 25.50 D), and 20.91 ± 3.07 D (range: 12.50 to 27.00 D) for the Micro F, POD F, and POD FT IOL groups, respectively.
Table 2 shows the mean lateral decentration in the x- and y-directions and the rotation outcomes measured 1 day and 3 months after surgery for each IOL group. The table includes both actual and absolute values. Actual values represent the directly measured values, whereas in absolute values the measurement's plus or minus sign has not been taken into account.
Lateral Decentration and Rotation Measured at In Vivo Condition 1 Day and 3 Months Postoperatively
Table A (available in the online version of this article) summarizes the significance values between the different lenses. When comparing lenses to each other, statistically significant differences were found between the Micro F and POD F IOLs for absolute rotation at 1 day (P = .01) and for displacement along the x-axis (in absolute value) at 3 months (P = .03). When comparing the Micro F and POD FT IOLs, differences were found for displacement along the x-axis (P = .03) and actual rotation (P = .01) at 1 day, whereas no statistically significant differences emerged for any of the variables at the 3-month follow-up visit. Finally, the POD F versus POD FT comparison revealed statistically significant differences for displacement along the x-axis (P = .02) and absolute rotation (P < .001) at 1 day, whereas no differences were observed for any of these variables at 3 months.
Statistical Significance Values for All Variables Evaluated for the 3 IOLs Measured at In Vivo Condition
Under in silico conditions, a rigid and lower simulated bag size (ie, clamps in this model) generally induced higher values of all evaluated variables because there is no deformation equilibrium between the capsular bag and the IOL material. Furthermore, the bio-mechanical response of the IOLs as a function of the clamp diameters is not a monotonic function due to the contact. This contact between the clamps and the IOL provides nonlinear behavior to the model.
The internal energy was a key output for providing essential information about the haptic design because it objectively determines the stability of the model. For the Micro F model, the internal energy increased exponentially from a diameter of 9.7 mm, presenting an internal energy of 13.06 nJ and provoking instability for lower closure diameters. Related to the internal energy, this IOL model had the highest values of compression force whatever the compression diameter, with a range of compression force between 47.676 mg (10.5 mm compression) and 190.324 mg (9.5 mm compression). High forces could cause damage or breakage of the capsule, which will depend on the mechanical properties of the capsule. The ultimate load of the posterior capsule for people between 16 and 98 years old decreases with age in the range of 1.590 to 119 mg force in the posterior capsule and 5.240 to 560 mg in the anterior capsule.27 For the other models, the energy and the compression force had an almost linear behavior whatever the compression diameter. For the POD FT IOL, the compression force when the haptic was compressed to 10 mm was 99.99 mg, and the lowest value was for the POD F IOL with 30.009 mg. Similar results were found in previous studies.13,15 The internal energy deformation is related to the haptic deformation capacity. The lower the internal energy deformation, the higher the haptic deformation capacity. Therefore, the haptics of IOLs with lower internal energy will adapt to all capsule sizes, with low stress in the capsule and prevention of capsular bag damage.
Another important factor is that the optical part of the lens (Figure 2) did not suffer any stress or deformation for any of the IOL models at 10-mm compression diameter. As a consequence, the optical quality of the patient would not be affected.
In terms of the axial displacement, the POD FT IOL presented the maximum value, 0.703 mm for a compression diameter of 10 mm. Bozukova et al15 found a smaller displacement value than in our study, 0.073 and 0.016 mm at in silico and in vitro conditions, respectively. However, in our study, the other two IOLs (Micro F and POD F) showed an insignificant axial displacement whatever the compression diameters. Axial displacement produces a change in the effective lens position that is associated with a refractive error; however, this problem is usually corrected by spectacle correction. This variable could not be measured in vivo.
In terms of tilt, the Micro F IOL presented the highest value, 0.785° at 9.80 mm of compression diameter. Tilt could not be measured in vivo with the software used for the other variables. Theoretical simulations performed by Holladay et al28 demonstrated that aspherical lenses can undergo a decentration of up to 0.4 mm and a tilt of up to 7° before they start to show a lower performance than their spherical counterparts. A study by Piers et al29 revealed a higher tolerance to malposition, the resulting threshold (critical) values being at 0.8 mm of decentration and 10° of tilt. The results achieved with the lenses assessed in our study are much better than the limits proposed by Holladay et al.28
In terms of rotation, all IOL models were rotationally stable. The POD F IOL had the maximum magnitude of rotation at all compression diameters, varying from 0.375° to 0.256°. These values would not affect clinically observable visual acuity. When the rotation was evaluated in vivo, the three lenses showed good results. The lowest rotation was obtained for the POD FT IOL (the toric lens), although the average mean was shown to be less than 3° for all lenses during the 1-day and 3-month postoperative follow-up visits. These values demonstrate high stability, even following potential capsular contractions.
The mean total lateral decentration for all models at different compression diameters was 0.084 ± 0.004 mm, although Bozukova et al15 found higher displacement values for the POD FT IOL. The three lenses showed minimal displacement 1 day after implantation, but also 3 months after surgery, when in many cases capsular contractions had already occurred that can increase the forces supported by the lenses. The greatest lateral decentration in x-direction occurred, at 1 day, with the POD FT IOL, with an average value of 0.10 mm. After 3 months, the POD F IOL showed the largest decentration in the same direction (0.285 mm), although this was not clinically significant and there were no statistically significant differences between the three lenses. The average values for the lateral decentration in y-direction were close to 0 for the three lenses and no statistically significant differences were found between them. Although statistically significant differences were found when comparing the displacement, tilt, and rotation between the different lenses, in some cases, these differences cannot be considered clinically relevant.
The results obtained with the in silico simulation and in vivo indicate that the POD FT IOL, which has the highest haptic–optic junction, has higher values of stress in the haptic–optic junction than the other models. This suggests that the haptic–optic junction is essential for the mechanical stability of the IOL in the capsular bag. We found that the POD FT IOL, which has a slightly wider haptic–optic junction than the POD F IOL, showed better stability in terms of rotation than the POD F and Micro F IOLs. It would be interesting to perform similar studies, under both in vivo and in silico conditions, with different platforms to evaluate how the haptic design affects the stability of the lenses.
The two methods used in this study have limitations, so some of the parameters cannot be measured with both systems. Numerical tests have not been done because it is not possible to measure in vivo, with any of the devices currently on the market, either the size of the capsular bag or the force compression inside the bag. More in vivo studies including very short-sighted and long-sighted patients should be performed to evaluate the behavior of the lenses in these extreme conditions.
In this study, we found several limitations of the ISO 11979-3.24 It establishes that the compression force shall be measured for IOLs intended for capsular bag placement, with the haptics compressed to a diameter of 10 mm. This is a limitation because the actual clinical outcomes might differ based on a patient's capsule size. In our study, the IOL mechanics were evaluated for different compression diameters simulating different capsular bag sizes. Another limitation of the ISO standard is that the axial rotation (as a parameter) is not the subject of the ISO 11979-3 standard,24 and it is a major issue of designs that are not rotationally symmetric. In our study, we performed the additional analysis of the axial rotation for each IOL model. Finally, there are no guide values of compression force in the ISO 11979-3, so we compared with the ultimate load provided by the study of Krag and Andreassen.27
Although the in silico evaluation does not reproduce the same conditions as those experienced by the lens in the capsular bag, it can nevertheless be considered an appropriate choice for analyzing the biomechanical stability of IOLs, fundamentally at the design stage. Although statistically significant differences have been found when comparing the displacement, tilt, and rotation between the different lenses, these differences cannot be considered clinically relevant, which would suggest that all three IOL models yield excellent stability in those terms.
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- Pérez-Merino P, Marcos S. Effect of intraocular lens decentration on image quality tested in a custom model eye. J Cataract Refract Surg. 2018;44(7):889–896. doi:10.1016/j.jcrs.2018.02.025 [CrossRef]
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- Gimbel HV, Neuhann T. Continuous curvilinear capsulorhexis. J Cataract Refract Surg. 1991;17(1):110–111. doi:10.1016/S0886-3350(13)81001-2 [CrossRef]
- Shah GD, Praveen MR, Vasavada AR, Vasavada VA, Rampal G, Shastry LR. Rotational stability of a toric intraocular lens: influence of axial length and alignment in the capsular bag. J Cataract Refract Surg. 2012;38(1):54–59. doi:10.1016/j.jcrs.2011.08.028 [CrossRef]
- Vounotrypidis E, Lackerbauer C, Kook D, Dirisamer M, Priglinger S, Mayer WJ. Influence of total intraocular lens diameter on efficacy and safety for in the bag cataract surgery. Oman J Ophthalmol. 2018;11(2):144–149.
- Remón L, Siedlecki D, Cabeza-Gil I, Calvo B. Influence of material and haptic design on the mechanical stability of intraocular lenses by means of finite-element modeling. J Biomed Opt. 2018;23(3):1–10.
- Lombardo M, Carbone G, Lombardo G, De Santo MP, Barberi R. Analysis of intraocular lens surface adhesiveness by atomic force microscopy. J Cataract Refract Surg. 2009;35(7):1266–1272. doi:10.1016/j.jcrs.2009.02.029 [CrossRef]
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- Bozukova D, Pagnoulle C, Jérôme C. Biomechanical and optical properties of 2 new hydrophobic platforms for intraocular lenses. J Cataract Refract Surg. 2013;39(9):1404–1414. doi:10.1016/j.jcrs.2013.01.050 [CrossRef]
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- Bozukova D, Werner L, Mamalis N, et al. Double-C loop platform in combination with hydrophobic and hydrophilic acrylic intraocular lens materials. J Cataract Refract Surg. 2015;41(7):1490–1502. doi:10.1016/j.jcrs.2014.10.042 [CrossRef]
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- Zhong X, Long E, Chen W, et al. Comparisons of the in-the-bag stabilities of single-piece and three-piece intraocular lenses for age-related cataract patients: a randomized controlled trial. BMC Ophthalmol. 2016;16(1):100. doi:10.1186/s12886-016-0283-4 [CrossRef]
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- Cabeza-Gil I, Ariza-Gracia MA, Remón L, Calvo B. Systematic study on the biomechanical stability of C-loop intraocular lenses: approach to an optimal design of the haptics. Ann Biomed Eng. 2020;48(4):1127–1136. doi:10.1007/s10439-019-02432-9 [CrossRef]
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Compression Force, Axial Displacement, Tilt, Rotation, and Lateral Decentration at Different Compression Diameters for All IOLs Measured at In Silico Condition
|Variable||Micro F||POD F||POD FT|
|Compression force (mg)|
| 9.5 mm||190.324||43.123||109.234|
| 10 mm||120.230||30.009||99.997|
| 10.5 mm||47.676||21.342||162.253|
| 11 mm||–||18.923||78.239|
|Axial displacement (mm)|
| 9.5 mm||0.212||0.032||0.954|
| 10 mm||0.158||0.025||0.703|
| 10.5 mm||0.000||0.000||0.352|
| 11 mm||–||0.000||0.000|
| 9.5 mm||0.785||0.085||0.087|
| 10 mm||0.048||0.074||0.075|
| 10.5 mm||0.032||0.055||0.234|
| 11 mm||–||0.049||0.085|
| 9.5 mm||0.204||0.375||0.295|
| 10 mm||0.157||0.348||0.255|
| 10.5 mm||0.050||0.325||0.054|
| 11 mm||0.000||0.256||0.014|
|Lateral decentration (mm)|
| 9.5 mm||0.082||0.083||0.093|
| 10 mm||0.083||0.083||0.091|
| 10.5 mm||0.080||0.083||0.087|
| 11 mm||–||0.080||0.080|
Lateral Decentration and Rotation Measured at In Vivo Condition 1 Day and 3 Months Postoperativelya,b
|Follow-up Visit||Micro F||POD F||POD FT|
| Dx (a), mm||0.065 ± 0.168 (−0.182 to 0.329)||0.084 ± 0.191 (−0.281 to 0.454)||−0.101 ± 0.163 (−0.272 to 0.213)|
| Dx (abs), mm||0.141 ± 0.107 (0.000 to 0.329)||0.167 ± 0.118 (0.010 to 0.454)||0.170 ± 0.081 (0.044 to 0.272)|
| Dy (a), mm||0.068 ± 0.077 (−0.032 to 0.211)||0.373 ± 0.435 (−0.388 to 1.641)||0.097 ± 0.137 (−0.135 to 0.287)|
| Dy (abs), mm||0.081 ± 0.062 (0.011 to 0.211)||0.3251 ± 0.394 (0.004 to 1.641)||0.145 ± 0.079 (0.010 to 0.287)|
| Rotation (a), degrees||0.680 ± 1.755 (−1.578 to 3.864)||0.160 ± 3.490 (−4.631 to 4.634)||−0.973 ± 1.087 (−3.314 to 0.354)|
| Rotation (abs), degrees||1.567 ± 1.007 (0.560 to 3.864)||3.084 ± 1.419 (0.165 to 4.634)||1.129 ± 0.911 (0.213 to 3.314)|
| Dx (a), mm||−0.058 ± 0.259 (−0.490 to 0.320)||−0.285 ± 0.139 (−0.434 to 0.129)||0.037 ± 0.140 (−0.225 to 0.191)|
| Dx (abs), mm||0.220 ± 0.137 (0.005 to 0.490)||0.098 ± 0.107 (0.001 to 0.434)||0.124 ± 0.066 (0.003 to 0.225)|
| Dy (a), mm||0.015 ± 0.159 (−0.465 to 0.216)||−0.018 ± 0.453 (−1.182 to 0.865)||0.115 ± 0.209 (−0.156 to 0.503)|
| Dy (abs), mm||0.111 ± 0.111 (0.001 to 0.465)||0.299 ± 0.331 (0.004 to 1.182)||0.175 ± 0.159 (0.011 to 0.503)|
| Rotation (a), degrees||0.744 ± 2.701 (−3.684 to 3.864)||1.186 ± 2.719 (−3.465 to 6.501)||0.261 ± 1.685 (−2.245 to 2.049)|
| Rotation (abs), degrees||2.393 ± 1.323 (0.764 to 3.864)||2.482 ± 1.516 (0.761 to 6.501)||1.590 ± 0.451 (0.845 to 2.245)|
Statistical Significance Values for All Variables Evaluated for the 3 IOLs Measured at In Vivo Conditiona
|Follow-up Visit||Micro F–POD F||Micro F–POD FT||POD F–POD FT|
| Dx (a), mm||.95||.03||.02|
| Dx (abs), mm||.80||.68||.99|
| Dy (a), mm||.20||.75||.32|
| Dy (abs), mm||.08||.06||.23|
| Rotation (a), degrees||.86||.01||.47|
| Rotation (abs), degrees||.01||.43||< .001|
| Dx (a), mm||.98||.42||.23|
| Dx (abs), mm||.03||.06||.69|
| Dy (a), mm||.96||.32||.56|
| Dy (abs), mm||.12||.42||.41|
| Rotation (a), degrees||.89||.838||.51|
| Rotation (abs), degrees||.98||.09||.10|