An intelligent contact lens can be used as an excellent interface between the human body and an electronic device for wearable health applications. Despite extensive investigations of smart contact lenses for diagnostic applications, there has been no report on electrically controlled drug delivery in combination with real-time biometric analysis. Here, we have developed smart contact lenses for continuous glucose monitoring and treatment of diabetic retinopathy. The smart contact lens device, built in a biocompatible polymer, contains flexible and ultrathin electrical circuits and a microcontroller chip for real-time electrochemical biosensor, controlled delivery of medications on demand, wireless power management and data communication. In models of diabetic rabbits, we were able to measure the levels of tear glucose to be validated by conventional invasive blood glucose tests and trigger drugs to be released from reservoirs for the treatment of diabetic retinopathy. Together, we have successfully demonstrated the feasibility of smart contact lenses for non-invasive, continuous diabetic diagnosis and therapy for diabetic retinopathy.
Recently, light bioelectronics has been extensively investigated to take advantage of its properties inherent to polymers and organic electronics for wearable and implantable health devices (1 1, 2) Based on this innovation, many types of medical devices have been developed for diagnosis (3), therapeutic (4) and teranotic applications (5) Wearable devices have been successfully applied for continuous glucose monitoring (5), electrocardiography (6), electromyography (7), photoplethysmography and pulse oximetry (8) They can provide important medical information for monitoring healthcare and diagnosing various relevant diseases. In addition, a pioneer semiconductor implantable drug delivery device has been developed for applications in the subcutaneous fluid (9) and triggered the development of delivery systems for implantable drugs on demand (10) Combining these technologies, many types of healthcare devices have been developed for teranostic applications at the interface of biological technologies, nanoscale and electronics (5, 11–13)
Among several portable health devices, smart contact lenses have attracted a lot of commercial attention for health applications (14, 15) The surface of the cornea presents a convenient and non-invasive interface to the physiological conditions in the human body. The eyes are directly connected to the brain, liver, heart, lung and kidney and can serve as a window to the body (16) In this context, Sensimed launched a product approved by the US Food and Drug Administration (FDA), Triggerfish, to monitor the intraocular pressure of patients with glaucoma in 2016 (14, 15) In addition, Google developed Google lenses for the diagnosis of diabetic patients in collaboration with Novartis (15) These smart contact lenses are especially important because they enable non-invasive and continuous monitoring of glaucoma and diabetes, respectively. In addition, smart wearable sensor systems integrated into soft contact lenses have been developed to measure the change in resistance of graphene sensors by binding glucose for remote monitoring of diabetes (17, 18) However, the electrical current and color changes in the sensors were proportional on a log scale to glucose concentrations, which may not be suitable for measuring the actual glucose concentration for an accurate diabetic diagnosis.
Here, we have developed a remotely controllable smart contact lens for non-invasive glucose monitoring and controlled medication administration to treat diabetic retinopathy. The multifunctional smart contact lens consists of five main parts: a real-time electrochemical biosensor, a flexible on-demand drug delivery system (f-DDS), a wireless resonant inductive energy transfer system, a circuit-based microcontroller complementary (IC) complementary chip with a power management unit (PMU) and a radio frequency remote communication (RF) system (Figure 1) The real-time amperometric biosensor is designed to detect glucose in tears, replacing the need for invasive blood tests. Medicines can be released from the self-regulating pulsatile f-DDS by remote communication. The resonant inductive coupling to a copper (Cu) receiving coil allows the wireless supply of an external power source with a transmitting coil. The device communicates with an external controller via RF communication. We evaluated and discussed the feasibility of this smart contact lens for diagnosis and therapy for diabetic retinopathy.
The smart contact lens is incorporated with a biosensor, an f-DDS, a wireless power transmission system from a transmitter coil to a receiver coil, an ASIC chip and a remote communication system as a ubiquitous platform for various therapeutic and diagnostic applications.
Preparation and characterization of silicone contact lens hydrogels
The silicone contact lens hydrogels were prepared with a chemical structure as shown schematically in fig. S1A. Silicone hydrogels were manufactured in the form of contact lenses with a diameter of 14 mm, thickness of 200 μm and radius curvature of 8.0 mm. Total attenuated reflectance – Fourier transform infrared spectroscopy (ATR-FTIR) showed clear peaks corresponding to the chemical bonding of the added monomers (fig. S1B). The five peak wavelengths were well compatible with a lotrafilcon A commercial silicone hydrogel contact lens. The silicone hydrogel contact lenses exhibited transmittance almost comparable to that of the poly (hydroxyethyl methacrylate) hydrogel contact lens. ) (PHEMA) as a control in the visible wavelength range (fig. S1C). The equilibrium water content (EWC) of the silicone hydrogel contact lens was 33.6%, higher than that of the PHEMA hydrogel contact lens (21.3%) and lotrafilcon A (24%) (fig. S1D), due to the high proportion of hydrophilic monomers containing silicone. The diameter of the silicone hydrogel lens increased by only 1 to 15 mm, while that of the PHEMA hydrogel lens increased from 2 to 16 mm. The hydrophilicity of the contact lens surface with silicone hydrogel was controlled by ozone plasma treatment. The contact lens with silicone hydrogel treated on the surface showed a lower contact angle with water than the contact lens with PHEMA hydrogel at all times (fig. S1E), and the drop of water was quickly absorbed by the contact lenses with silicone hydrogel (fig. S1F).
Real-time in vitro electrical detection of glucose concentrations in tears
An eye glucose sensor has been designed with three electrodes to have a low electrical resistance for the electrochemical reaction of facilitated glucose (Fig. 2A) The working electrode (WE) and the counter electrode (CE) were prepared with platinum (Pt) for the efficient electrochemical reaction. To improve the adhesion between polyethylene terephthalate (PET) and Pt, a layer of Cr was deposited on the PET substrate as an adhesive layer before the deposition of the Pt layer. The reference electrode (RE) coated with silver / chloride silver (Ag / AgCl) increased the accuracy of the amperometric electrochemical glucose sensor in the fluidic environment, providing a constant voltage to the WE during glucose measurement. To monitor the lacrimal glucose content with high sensitivity and stability, we coated a mixed solution of glucose oxidase (GOx), bovine serum albumin (BSA), poly (vinyl alcohol) (PVA) and chitosan in WE. After drying, glutaraldehyde was added to the crosslinking of chitosan and PVA for the immobilization of GOx with BSA. To confirm the strong correlation between blood glucose levels and tears, glucose concentrations in normal and diabetic rabbits were measured before and after feeding and fasting three times. Diabetic rabbits had higher blood and blood glucose concentrations than normal rabbits (Fig. 2B) These blood glucose levels and tears appear to be in the reasonable range, because the normal blood glucose level for non-diabetics during fasting is between 70 and 130 mg dl−1 (19) Due to the large sampling time interval, it was not possible to observe the latency time in the increase of glucose concentrations between the blood and the tear, as reported elsewhere (19) However, we make clear the strong repetitive correlation between blood and glucose levels in tears. These results indicated the feasibility of measuring the level of tear glucose as an alternative to measuring blood glucose for the diagnosis of diabetic diseases.
(AN) Schematic illustration of an eye glucose sensor with three electrodes (WE, working electrode; ER, reference electrode; CE, against electrode) and the mechanism for measuring glucose in tears. (B) Correlation between blood glucose and tear levels in normal and diabetic rabbit models. (Ç) Real-time electrical detection of glucose concentrations compared to that of PBS. (D) Current glucose sensor change showing selectivity to 0.35 and 0.7 mg dl−1 ascorbic acid (AA), 22.5 and 45 mg dl−1 lactate, 18 and 36 mg dl−1 urea and 5 mg dl−1 glucose. (AND) The long-term stability of the glucose sensor after storage for 0, 21, 42 and 63 days (n = 3)
As shown in Fig. 2C, we could measure the glucose concentration in real time from the electric current change in vitro using a potentiostat. The current increased from 0.41 to 3.12 μA with an increase in glucose concentrations from 5 to 50 mg dl−1. This current change interval may be suitable for remote monitoring of physiological glucose levels. To assess glucose selectivity, we apply potentially interfering molecules of ascorbic acid (A), lactate (L) and urea (U) to the tear (Fig. 2D) ALU concentrations are reported to be about 0.70 mg dl−1 for (20), 18 to 45 mg dl−1 for L (21) and 36 mg dl−1 for you (20) in tears. When the corresponding concentrations of interfering molecules (A, L and U) were added to the glucose detection system, only a little noise was observed with an insignificant change in current. Unlike A, L and U, addition of 5 mg dl−1 of glucose rapidly increased the current to 0.42 μA. In addition, we evaluated the long-term stability of glucose sensors (Fig. 2E) After manufacture, smart contact lenses were stored in sterile phosphate buffered saline (PBS) at 20 ° to 25 ° C, similar to the actual contact lens storage environment, for 21, 42 and 63 days. The performance of the glucose sensors was maintained steadily with a deviation of less than 2% for up to 63 days (n = 3)
On-demand release of f-DDS
The f-DDS was manufactured with dimensions of 1.5 mm by 3 mm by 130 μm (Fig. 3, A and B) A layer of exfoliation and a buffered silicone oxide (SiO2) were deposited on a glass substrate and the drug reservoir was covered with a defective Au anode electrode. The laser removal process (LLO) using an excimer laser melted locally and dissociated the exfoliation layer. A SiO buffer2 The layer supported the top layer of the device during the LLO process and blocked the heat flow generated during laser-induced exfoliation. In addition to controlling the duration of the laser firing, the thickness of the SiO buffer2 The layer was an important factor in minimizing thermal damage to the device during the LLO process. We used two photoresists different from SU8-5 and SU8-50. SU8-5 has lower viscosity and strength than SU8-50. Consequently, SU8-5 was used to isolate the electrode, except that the drug delivery site for the stable operation of the f-DDS and SU8-50 was used to build the DDS. Transverse scanning electron microscopy (SEM) showed the electrodes and isolated layers of the reservoir (Fig. S2). The mechanical flexion test was performed to assess the mechanical reliability of the f-DDS on a flexible substrate (fig. S3, A and B). The operating current of the f-DDS was maintained without noticeable changes during the mechanical durability test of up to 1000 cycles (fig. S3C).
(AN) Schematic illustration of the manufacture of f-DDS. i) Growth of buffer silicon dioxide (SiO2) layer on a glass substrate; (ii) deposition of Ti, Au and Ti metals for anode and cathode electrodes; (iii) standardize the SU8 drug reservoirs; (iv) loading of drugs; (v) attach PET and laser scanning of the device; (vi) disconnect the f-DDS; and (vii) Ti attack with SU8 isolation. (B) Photograph of f-DDS. Photo credit: Beom Ho Mun, KAIST. (Ç) SEM images of f-DDS before and after the gold electrochemical test. Scale bar, 250 μm. (D) Microscopic images of confocal fluorescence of rhodamine B dye released from drug reservoirs. Scale bars, 300 μm (left) and 500 μm (right). (AND) Current change to f-DDS. (F) Pulsatile release of genistein. (G) Standardized content of genistein released in reservoirs (n = 6) compared to the initial load content.
The loaded medications were selectively released from the medication reservoir by the voltage on / off control. As shown in the SEM image of the Au anode electrode, a thin Au membrane covered the entire area of drug-loaded reservoirs without any drug leaks (Fig. 3C, left). After applying an electric voltage of 1.8 V, the Au membrane was dissolved in 40 s (Fig. 3C, right). The Au layer was melted in PBS under constant tension in the form of AuCl4–. Fluorescent confocal microscopy showed the rhodamine red dye released from a reservoir applying the electrical potential (Fig. 3D) The current between the anode and cathode electrodes increased to 6.08 ± 0.16 μA, and Au anodes were dissolved slowly under a slight decrease in current from 6.08 ± 0.16 μA to 4.35 ± 0 , 11 μA (Fig. 3E) Genistein has been released in a pulsatile manner in three different drug reservoirs (Fig. 3F) The anode was slowly dissolved by the microscale current, and the drug was released almost completely after the current was recovered to its initial state. We can detect 89.97 ± 37.10% of the genistein loaded in PBS, confirming that a therapeutic amount of medication can be released from f-DDS (Fig. 3G) In addition, a diabetic therapeutic amount of metformin can be released from smart contact lenses by synchronized feedback from therapy at the point of care and other teranostic applications (fig. S3D).