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Surgeons treat cataracts by replacing the clouded lens with an intraocular lens (IOL), but patients are required to wear reading glasses for tasks requiring near vision. We suggest a new voltage-controlled accommodating IOL made of an ionic polymer metal composite (IPMC) actuator to change focus. An in vitro experiment was conducted where an actuator was placed inside the eye and moved with applied voltage. The lens attached to the actuator was deformed by its movement to change the patient’s focus. The results showed that this system can accommodate a change of approximately 0.8 diopters under an applied voltage of ± 1.3 V.
© 2016 Optical Society of America
1. Background
Cataracts are one of the most common visual disorders, and surgical treatment has been established for it. Cataracts are caused by clouding of the lens. Currently, about 285 million people have been blinded by this condition. Cataracts are responsible for 33% of all cases of blindness [1].
In most cases, cataracts are treated by replacing the clouded lens with an intraocular lens (IOL). The first IOL introduced was monofocal with one focal length; therefore, the vision of a patient with this type of IOL is limited to the range of the depth of focus [2]. The subjective depth of focus has been reported to be 0.59 diopters (D) for a pupil diameter of 4 mm. Patients have to wear reading glasses to perform daily activities or work involving near-vision. To solve this problem, Hoffer et al. [3] introduced the multifocal IOL, which has two focal lengths so that the patient can simultaneously see far and near objects. However, the mechanism of the multifocal lens has three major problems. First, the multifocal IOL has a large aberration because its shape is far from the ideal lens shape. This reduces the contrast difference in patient sight. Second, it cannot attain focus when the patient reads tiny letters. This problem occurs because the multifocal IOL has both far and near foci; if the patient wishes to see objects that are too far or too near, the lens cannot attain focus, and the patient has to wear glasses. The third problem is the occurrence of glare and halos. This phenomenon occurs when the patient sees light under dark conditions; this reduces the contrast difference, and the information of the light is scattered. A multifocal IOL has edges between the near-focus lens and far-focus lens. These edges do not have the ideal lens shape and cause glare and halos.
Some companies and researchers have suggested new multifocal IOLs with reduced glare and halos. Lentis M plus is a multifocal IOL with two lenses [4]. One lens allows far vision and forms the upper part of the IOL. The other lens allows near vision and forms the lower part of the IOL. This IOL has only one edge for reduced glare and halos. However, this IOL is not a fundamental solution to the above problems. A comprehensive solution requires an IOL that can change focus, i.e., an accommodating IOL.
Many researchers around the world have suggested accommodating IOLs. There are three general types of accommodating IOLs, as shown in Fig. 1: moving, sliding, and deforming lenses [5].
Well-known moving-lens-type IOLs include Crystalens and Trulign (Bausch & Lomb Inc.). These IOLs are set in the capsula lentis and have springs at its sides. When the ciliary muscle contracts, the capsula lentis is deformed. As the IOL spring is bent by the capsula lentis, the IOL position changes. Consequently, the focus of the patient’s eyes is changed.
Akkolens International BV has proposed a sliding-type IOL called Lumina IOL [6]. This IOL is also set in the capsula lentis. This IOL contains two lenses that slide by deformation of the capsula lentis by the ciliary muscle. Figure 1 shows the shapes of these lenses along with their focus change mechanisms.
NuLans Ltd. proposed another method to change the focus [7]. Similar to the other methods, it uses the deformation of the capsula lentis to change focus except that a soft lens material is used; the surface shape is changed by the deformation of the capsula lentis. The Flex Optic IOL (Advanced Medical Optics, Inc., Irvine) is a single-optic IOL [5] with flexible lens and an arch. When the ciliary muscle moves, the arch is deformed, which changes the shape of the lens surface. McCafferty et al. suggested a similar deforming-type IOL that uses an arch and silicone gel [8]. The FluidVision Lens (PowerVision Inc., Belmont, CA) drives the fluid of a polymer-matched refractive index [5]. This lens also controls the focus by changing the lens surface with the ciliary muscle.
The above methods can be used to control the focus of the lens of a patient; however, a significant problem exists. The focus control mechanism is based on the deformation of the capsula lentis, and such an IOL cannot be used for old patients with presbyopia whose capsula lentis has hardened and cannot be deformed. Presbyopia usually starts at an age of 40–45 years [9], and cataracts also increase after an age of 45 years. This means that most patients are already or will be presbyopia patients. Thus, a new accommodating IOL system for presbyopia patients is needed.
Some researchers have attempted to address this problem. Peng et al. suggested an accommodating IOL that uses a conductive liquid moved by an electric field [10]. The surface of the liquid is changed by the electric field, which changes the focus. Hasan et al. suggested an accommodating IOL that uses a shape memory alloy spring controlled by electric heat [11]. Vdovin et al. suggested an accommodating IOL that uses liquid crystal lens deformed by conductive metal moved by an electrical field [12]. Wei et al. used dilectric elastomers to control the focus [13]. These methods can be used for presbyopia patients but ignore electrical leakage. Electrolysis of water occurs above 1.5 V; to ensure safety, IOL systems should not use over 1.5 V, but this was not the case for the above approaches. A low-voltage system is needed.
Our aim was to develop a new control system for changing the focus of an accommodating IOL. To control the changing focus, we used an ion polymer metal composite (IPMC) actuator that can induce accommodative changes in a targeted lens. In this study, we used swine crystalline lenses because they were easy to obtain. We were aware that the swine crystalline lens is not optimum material for accommodation studies but we found it was better than other artificial materials.
2. Materials and methods
Figure 2 shows our proposed system, which is set under the cornea. A soft, clear material is placed in the capsula lentis that functions as the IOL. It is deformed by using the IPMC actuator. Figure 3(a) shows the shape of the IPMC actuator, which has two major components: a moving part and fixed part. The moving part has six wings to push soft materials. These wings have a planar figure because such a shape has the lowest second moment of area and thus is most suitable for movement. If the number of wings is increased, the deformation of the soft material becomes more spherical. In contrast, the fixed part has a spherical surface to mimic the eye, and this part is hard to move because of the high second moment of area. The fixed part should be hard because, if this part deforms, the moving part also deforms, its second moment of area increases, and its range of movement decreases. A coil is used to apply voltage to the moving part. In the future, we will apply a variable magnetic field and control this actuator with a wireless system [12]. To prevent electric leakage, this actuator should be coated with a nonconductive membrane.
2.1 Ion polymer metal composite actuator
The IPMC actuator is a soft actuator that falls in the category of ionic electromechanically active polymers (IEAPs) [14, 15]. This actuator has two major advantages for use in eyes. First, it has a simple shape in the form of a membrane with a thickness of less than 0.2 mm. Thus, it can easily be placed inside the eye. The biggest difference between an IPMC actuator and other mechanical actuators such as a motor is the proportion of the actuator part. A motor has a coil, line, bush, and other parts. Only one rod can rotate, and the other parts cannot move. However, the IPMC actuator can move its whole body. Thus, it has a huge advantage in narrow spaces, such as inside the eye. Second, this actuator can be moved by applying a voltage of under 1.5 V, which is sufficiently low for use in the body. For use inside the body, we have to consider the leakage of electricity and its influence. In this case, the actuator will be used inside the eye, which contains a great deal of water. If a high voltage is used, electrolysis of the water occurs, and the eye will burst. Thus, we have to choose a voltage low enough that electrolysis will not occur.
Figure 3 shows the mechanism of the IPMC actuator. It consists of an ionic conductive polymer coated with gold or platinum. Water molecules and cations exist inside the polymer. A cation and water molecule are connected by hydrogen bonding, and they collectively act as a large molecule. When a voltage is applied to the metal membrane, an electric field is generated in the membrane. Under the influence of the electric field, the cations move and gather at one side with water molecules; consequently, there is a difference in volume between the anode and cathode. As a result, the membrane is bent to one side. The performance of an IPMC actuator is determined by the thickness, water content ratio, and equivalent weight. If two IPMC actuators are of the same material, the thinner one has a larger movement area and weaker power. The thinner one can move a larger area because of its low second moment of area. However, it has weaker power because the amounts of cations and water are less than those for the thicker one. If we change the physical properties of the IPMC actuator, the performance increases with the water content ratio and equivalent weight.
In this study, we used Nafion (Sigma-Aldrich Co.) as the ionic conductive polymer and gold for the metal membrane. The thickness of the actuator was 0.2 mm, and the diameter of its moving part was 13 mm. This diameter is slightly greater than that of the lens of swine. We adopted this size for the in vitro experiment.
2.2 In vitro experiment
We conducted an in vitro experiment to evaluate the performance of our system. Figure 4 shows the experimental system, which was intended to be a preliminary step.
Fig. 4 In vitro experiment: (A) whole system, (B) zig for actuator and lens, (C) pathway of the light, and (D) close-up view of zig and lens.
To measure the applied voltage precisely, we used a wire connected to a potentiostat. To determine the refracting power at each point of the lens, we performed measurements using a compact Shack-Hartman wavefront aberrometer (CSHWA) [16], as shown in Fig. 4. The CSHWA had dimensions of 0.21 m × 0.17 m × 0.06 m with a weight of 0.55 kg. The light source for wavefront sensing was a super-luminescent diode (SLD) with a wavelength of 840 nm. The illumination system was designed to form a small spot on the retina through the eye optics by using narrow Gaussian beam characteristics. The refraction ranged from −25 to + 21 D. We set up the CSHWA with other instruments, and the CSHWA readily measured the refraction and aberrations of unknown targets. We analyzed the measured images with the SCWA by using Zernike polynomials up to the sixth order. The diameter of the pupils for the analysis was 2 mm. We converted the second-order defocus term to spherical equivalents to determine the power of the accommodation [17]. We performed wavefront sensing at 3 Hz for 14 s. We used the lens of a swine as the soft, clear material and the retina of a swine as a diffuser panel. Because swine lenses are much harder than human lenses, the results of this experiment should also be useful for human beings. The average age of swine in this experiment was 6 months. The average diameter of the swine lenses was 11 mm, and the height was 7 mm. We conducted this experiment in a saline solution to keep the swine lens wet, and the IPMC actuator showed good performance in water.
The light beam from the SLD of CSHWA was lead to retina thought the mirror, then reflected at the retina, went through the mirror and lens system, and finally returned to the CSHWA, as shown in Fig. 4. We measured each deviation of the light spot on the image sensor of the CSHWA, which is corresponding the lens array, and calculated the refracting power at each point of the accommodated lens. When we applied a voltage of 1.3 V to the IPMC actuator, the swine lens was deformed and the reflecting power was changed. We applied square wave, 1.3 V amplitude and 0.25 Hz frequency. We performed the wavefront sensing continuously before and after the change of the voltage.
3. Results
We conducted the experiments six times, and the results are shown in Figs. 5 and 6. The accommodating range was 0.53–1.11 D, and the average was 0.86 D (Table 1).
Fig. 6 Results of experiment No. 1 (time: 1000 ms). Upper left: actuator and lens. Lower left: position of light emitted from the sensor and its path through the lens and back to the sensor. Upper right: total aberration. Lower right: higher-order aberrations.
Table 1. Results of spherical equivalent.
4. Discussion
4.1 Efficiency of refraction
For efficiency of refraction, two points need to be discussed. The first point is the amplitude of the spherical equivalent. For an accommodating IOL, a range of at least 2 D (equal to that of a 50-year-old person) is required. However, this system controls the refracting power in a range of 1 D under an applied voltage of ± 1.3 V. There are two methods to increase the refraction range. One is to maintain a spherical shape. As shown in Fig. 6, the lens did not maintain a spherical shape; rather, it took the shape shown in Fig. 7 because a difference in deformation occurred between the part attached to the IPMC actuator and the part that was not attached. To prevent this difference in deformation, we should put a thin ring between the IPMC actuator and lens. The second method is to use a soft, clear material for the lens. In the present experiments, we used the lens of a swine, which is much harder than human lenses. We believe that we can increase the range of refraction by at least 1 D by using these two methods. The second point is the aperture diameter. In this study, we measured a diameter of 2 mm, but an accommodating IOL needs a 4 mm diameter. This is because the centerline of the lens deviates from that of the IPMC actuator. As a result, the range of the accommodating IOL is less than the 4 mm diameter. In the future, we will optimize the shape of the IPMC actuator design so that it will fit the lens.
4.2 Noise of results
The results of the experiments had a significant noise component. We believe the phenomenon explained in Fig. 7 also increased the noise. To calculate the refractive index of the lens, we used a mathematical expression that assumes a spherical lens. However, in the present system, the lenses could not maintain a spherical shape. We believe that, if a ring is placed between the lens and IPMC actuator, this noise will also disappear.
4.3 Alternative method for accommodating IOL
In this study, we adopted the deforming lens method presented in Fig. 1. However, the other methods are also useful, and the IPMC actuator can be used for both. For the moving-lens- and sliding-lens-type IOLs, we have to connect a small lens to the IPMC actuator. Because no such connection is required for the deforming-lens-type IOL, we adopted it. In the future, we will develop a method to connect the clear, soft material to the IPMC actuator and evaluate the performance.
4.4 Electrical supply
In this experiment, we used a wire to supply voltage to the IPMC actuator. This is possible in an in vitro experiment but not for patients. In our prototype system, we induced an electromotive force to supply the voltage. However, this method has a disadvantage in that the patient requires a magnetic power generator. This is not the ideal method for realizing an accommodating IOL. However, because of advances in technology, the size of batteries is decreasing. We believe that a battery small enough to install in the eye will be developed in the near future.
4.5 Artificial lens
In this experiment, we used swine lenses as soft material, as shown in Fig. 2. For clinical use, however, the soft material should be made of an artificial biocompatible material, such as silicone. An artificial material also has the advantage that it can be designed to have the desired shape and hardness. DeBoer et al. suggested a valved deformable liquid balloon lens that is softer than a human lens [18]. By using such a material, our system can achieve better performance. We also have to consider other methods. Figure 1 shows three types of accommodating IOLs, and we think the optimal shape of the lens for each type is different. In the future, we will design lenses and measure the performances of each type.
4.6 Biological compatibility
Because this study was the primary step, we ignored biological compatibility in this experiment. We have to consider two problems: the biological compatibility of the IPMC actuator and that of the whole system. Regarding the IPMC actuator, Nafion has a sulfonic acid group and is too strong an acid to insert into body directly. There are two major ways to treat this problem. One is covering Nafion with a biocompatible membrane. This is the most basic method, but the problem is that this membrane reduces the range of movement of the IPMC actuator because it also has a second moment of area. The other way is changing the material. Nafion is a strong acid because of the sulfonic acid group, but there are other ion-compatible polymers that use not the sulfonic acid group but the carboxyl group, which is a weak enough acid to insert into the body. The second problem is the whole system. Our present system is bared to the body, but some researchers have argued against bared systems and that the system should be place in a capsule [19]. A capsular interface also helps increase the biocompatibility of the IPMC actuator. In the next step, we will try capsulation.
5. Conclusion
We developed a voltage-controlled accommodating IOL system using an IPMC actuator that can be controlled by applying a voltage. Through an in vitro experiment, we determined that this system can accommodate approximately 0.8 D under an applied voltage of ± 1.3 V.
Funding
Japan Society for the Promotion of Science (JSPS) (15K15634, 15K16342).
Acknowledgments
We thank Topcon Corp. for providing the CHSWA to us and a research fund to T. M.
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