The 7 × 4 mm display with ×3 magnification mimics viewing the world through a flat display measuring 1.24 × 0.71 m in size, located 1 m away from a standard-sized eye, corresponding to a horizontal visual field of 63.4 degrees. By varying the size of the projected image, different optical focusing designs can have higher visual acuity at the cost of a smaller visual field. Digital magnification of the camera image can also achieve high equivalent visual acuity at the cost of a smaller visual field. Because of the short separation of the projector from the retina, the depth of focus of the optical system is 0.066 mm, which necessitates careful positioning of the implant to ensure a sharp focus on the fovea. Inclusion of an externally adjustable focusing mechanism can potentially reduce the implantation precision needed. To prevent any phototoxicity and discomfort from a projector positioned within the eye, the brightness and the spectral width of the light source can be filtered to be within a physiologic range—on the order of nanowatts per mm2.27 LED microdisplays are designed for long-term human visual exposure and do not produce ultraviolet or other toxic wavelength light.
The size and mass of a device that can be maintained in the anterior segment is limited by the geometry and strength of the supporting tissue. The telescope implanted into the anterior segment in patients with macular degeneration weighs 120 mg.28 To reduce the risk of chafing on the interior eye surface, an intraocular projector implanted into the anterior segment should not exceed the size of the crystalline lens, which is about 10 mm in diameter, 6 mm in thickness, and weighs 270 to 300 mg.29 Microdisplays measure about 7 mm in length, 1.5 mm in thickness, and weigh in the range of 150 to 200 mg (Fig. 2).30 This and the focusing lens need to be contained within a waterproof housing, which can contain inert gas such as nitrogen to allow the device to be neutrally buoyant in an aqueous medium, reducing stress during eye movements. The hardware needed to receive and process power and data cannot be fitted within the eye using current technology, and thus needs to be connected to the projector using a transscleral cable, as is the case with retinal prosthesis systems.
Experience with retinal prostheses provides data on heat limits of intraocular electronics. Opie et al demonstrated 135 mW distributed over an area of 5 × 5 mm in contact with the suprachoroidal space to be the safety limit and recommended no more than 19 mW/mm2 for implants in contact with the retina.31,32 Clinical devices need to conform to international requirements of no greater than 2°C increase at any point on its surface (ISO norm 14708-1 article 17.2). Lazzi33 conducted mathematical heat modeling showing a surface temperature increase of 2°C of a conformally coated 4 × 4 mm midvitreous implant at approximately 35 mW. Intraocular projectors have a higher power budget because they are contained within a larger housing. Our demonstration projector power consumption is adjustable from 40 to 75 mW.
Any intraocular electronic device must be sealed in a biocompatible and durable waterproof housing. The interior of the eye is an immune-privileged area known to tolerate a variety of materials in the long term. Silicones, acrylics, and titanium are being used successfully in current intraocular implants.34 Because light is emitted through the lens, only a power cable needs to exit the housing of the device, making a projector simpler to package than retinal prostheses that have many stimulatory electrodes. Given that device contents remain secure from corrosion, the lifespan will be limited by its fastest wearing component. Current lifetime of the LED backlights is ∼50,000 hours (11 years at 12 hours a day).35 LCD displays have a lifetime several times longer than LED backlights.
Several approaches could be used for stable implantation of this device in the anterior segment. Implantations into the capsule or ciliary sulcus would be the least invasive and maintain normal anatomic separation of the anterior and posterior segments, but they require an intact capsular bag and appropriately small implants. Fixation to the sclera, either by placing haptics into scleral pockets, melting haptic ends, or suturing to the sclera such as that done for intraocular lenses without capsular support are possible approaches (Fig. 5).36 A pars plana incision, scleral tunnel (such as that for extracapsular cataract surgery), or an open sky trephination can be used to insert the device. If the cornea is trephinated once the device is implanted, the patient's own cornea can be placed back, even if opaque (Fig. 5F) The size of surgical wounds could put the patient at risk of expulsive hemorrhage. However, the required incision size can be reduced through incorporation of technologies such as flexible display and housing. The patient's feedback (such as with adjustable sutures for strabismus surgery) or intraocular endoscopy can be used to fine-tune the device's position and focus before conclusion of the surgery.
Ocular movements normally shift the image on the retina. Because this will not happen with the intraocular projector, as with the Argus retinal implant, patients will have to use head movements and keep their eye steady. Eye tracking can be used to shift the image on the display in accordance with the eye movements, and thereby eliminate this problem.37 Medical and surgical complications will likely be similar to those of retinal prostheses and may include leakage of fluid from the eye near the cable crossing the sclera, which may result in hypotony (10% rate in Argus) or infection.38 There is also a risk of chronic inflammation if the device is not stably implanted. Prosthetic membranes could grow along the cable or haptics onto the device, although this has not been reported to be common in retinal prostheses. Patients with severe inflammatory eye disease will likely experience an increased rate of complications such as scleral melt as compared with those undergoing retinal prostheses.
The cost of microdisplays is low due to commercial availability (less than $200 per unit), but development, manufacturing, and regulatory expenses will likely result in a first-generation device of similar cost to retinal prostheses ($100,000). This will limit the initial number of patients who are candidates for this device, but with time and refinement, a wider patient pool could benefit.
We here review past work and technologies relevant to development of an intraocular projector for treatment of corneal blindness and demonstrate a prototype device. Although much work remains to be performed, including long-term safety and efficacy studies, we believe that an intraocular projection system is within reach of current technology. Such a device will not require donor tissue or corneal implants and has the potential to offer a new paradigm in the treatment of corneal diseases: that corneal clarity is not required for high-quality vision.
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