Wireless smart contact lens for diabetic diagnosis and therapy – ABOUT MAG 2020


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, 1113)

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.

Figure 1 Schematic illustration of smart contact lenses for diabetic diagnosis and therapy.

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.

Figure 2 In vitro electrical detection of ocular glucose sensors.

(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).

Fig. 3 On-demand delivery of medicines using an f-DDS.

(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).

Wireless power transmission and remote communication

A wireless energy transmission system was developed via resonant inductive coupling. The receiver coil incorporated in the smart contact lens received different electrical energy from the transmitter coil, depending on the distance (fig. S4A). The efficiency of wireless energy transmission between two coils was measured with a network analyzer, which was inversely proportional to the distance (fig. S4A). The required energy consumption of the PMU, the sensor reading block and the remote communication unit (RCU) in the smart contact lens was 43, 34.4 and 2.3 mW, respectively (fig. S4B). The RCU transmitted data at a rate of 445 kbit · s−1 in the 433 MHz industry-science-medicine (ISM) frequency range using on-off coding modulation and can be controlled to disable energy savings when data is not transmitted. Using resonant inductive coupling, the application-specific integrated circuit chip (ASIC) connected to an additional capacitor for energy storage successfully received electromagnetic energy at a distance of 1 cm from the transmitter coil, with an efficiency of 2%. The efficiency was sufficient to maintain the basic operation and remote communication of the smart contact lenses. The average output code of the ASIC chip’s analog-to-digital converter (ADC) was proportional to the input current (fig. S5, A and B). Total input conversion was available up to 4.1 μA with a resolvable 150 pA input, which was suitable for electrical glucose detection using the eye glucose sensor. The eye glucose sensor and f-DDS were operated under the control of the ASIC chip, applying the corresponding polarization voltages (fig. S5, B and C). The converted data from the biosensor were serialized by the ASIC chip and successfully transmitted to an external device of the personal computer (PC) using the power systems and wireless remote communication (fig. S5D).

Manufacturing and evaluation of the integrated smart contact lens

Based on preliminary experimental results, an intelligent contact lens was manufactured by chemical cross-linking of the silicone hydrogel precursor solution containing a PET film, which was incorporated with a glucose biosensor, an f-DDS, an ASIC chip, a power of Receiver cu and RF communication coils and passivated with Parylene C (fig. S6A). The reader coil, connected to a commercial power amplifier, wirelessly transferred enough electrical energy to the smart contact lenses for the real-time detection of glucose in tears and the remote control of the f-DDS (fig. S6B). A constant potential was applied to the ER of the electrochemical glucose sensor, allowing high sensitivity and stability. The biosensor output data was transmitted wirelessly via remote communication using a custom ASK receiver (amplitude shift shift keying), an Alf Vergard Risc (AVR) and a PC. The data transferred remotely showed that the current change in the glucose sensor was proportional to the level of glucose applied in vitro, confirming the feasibility of detecting wireless electrical glucose in real time using the smart contact lens (fig. S6C). The output current variation values ​​from 0.40 to 3.13 μA were similar to those of glucose measurement using an in vitro potentiostat in Fig. 2C. In addition, the remote control of the ASIC chip demonstrated the administration of medications on demand to apply a constant voltage of 1.8 V to the f-DDS (fig. S6C). High water content silicone hydrogel contact lenses did not cause substantial damage to the biosensor, f-DDS and other medium-sized components.

In vivo diagnostic and therapeutic applications of the smart contact lens

Before in vivo applications, the safety of integrated smart contact lenses was assessed in the eyes of New Zealand white rabbits for a period of 5 days (fig. S7). Histological analysis of the eyes of rabbits extracted with hematoxylin and eosin (H&E) staining showed no noticeable damage to the epithelia, stroma and endotheliums of the cornea of ​​rabbits after using smart contact lenses for 3 and 5 days compared to the cornea normal rabbits. Although our smart contact lens induced some degree of corneal swelling, it did not trigger an inflammatory reaction after 5 days. The corneal swelling was probably caused by low oxygen transfer through the closed eyelid during sleep while wearing the contact lens, which leads to the accumulation of lactic acid and water inside the cornea as a result of the osmotic change. There were no serious infections or adverse reactions on the ocular surface or changes with the lens in place. Overall, our results demonstrated the preliminary safety of smart contact lenses while being placed in the eye.

Then, we performed the evaluation of the smart contact lens integrated in the eyes of a diabetic rabbit for biosensor applications and drug administration, as shown schematically in Fig. 4A. The integrated wireless smart contact lens for glucose detection only (fig. S8A) or for glucose detection and drug delivery (fig. S8B) was used in the rabbit’s eye and operated by wireless energy transfer between an external coil of the transmitter and a receiving coil on the smart contact lens (fig. S8C). The portable power transmission system can be installed on smartphones or smart phones, as shown schematically in Fig. 4A. Diabetic rabbits were injected with insulin, anesthetized with ketamine and equipped with our smart contact lens (film S1). After using the smart contact lens, the eye glucose sensor indicated an increase in glucose concentration by up to 30.53 mg dl−1 contacting the tear glucose and then decreasing to 16.72 mg dl−1 by the effect of insulin on glucose metabolism, which corresponded well to the blood glucose concentration profile determined by a glucometer (Fig. 4B) The actual level of tear glucose measured by the glucose test was well compatible with the glucose level converted from the current output values. Parviz’s group previously developed a contact lens sensor system and performed wireless glucose monitoring using a polydimethylsiloxane (PDMS) eye model (20, 22) While the output current of the online sensor was in the range of 0 to 400 nA for the glucose concentration of 0 to 10.81 mg dl−1 (20), the output current of the wireless sensor was in the range of 0 to 80 nA for the glucose concentration of 0 to 36.03 mg dl−1 (22) On the other hand, we wirelessly measure the actual level of tear glucose over a wide physiologically significant range from 0 to 49.9 mg dl−1 in vitro and in vivo with improved sensitivity (Figs. 2C and 4B and fig. S6C).

Fig. 4 Aplicações in vivo de sistemas de lentes de contato inteligentes.

(AN) Ilustração esquemática do diagnóstico diabético in vivo e terapia das lentes de contato inteligentes. (B) Medição sem fio in vivo em tempo real dos níveis de glicose lacrimal com a lente de contato inteligente. Os níveis de glicose no sangue e nas lágrimas foram medidos (i) após a injeção de insulina e anestesia para o uso de lentes de contato inteligentes em PBS. (ii) O nível de glicose nas lágrimas aumentou devido à glicose nas lágrimas e diminuiu, refletindo a diminuição no nível de glicose no sangue devido à insulina injetada. O nível de glicose no sangue foi medido a cada 5 minutos com um glicosímetro comercial. (C) Imagens microscópicas de fluorescência de medicamentos absorvidos na córnea, esclera e retina de coelhos usando lentes de contato inteligentes carregadas com (linha superior) e sem (linha inferior) genisteína. Barra de escala, 0,1 mm. (D) Análise de câmera térmica infravermelha para a temperatura do olho, lente de contato inteligente e bobina de transmissão após operar por 0, 15 e 30 min.

Além disso, poderíamos disparar remotamente a liberação da genisteína antiangiogênica do f-DDS nas lentes de contato inteligentes, aplicando o potencial elétrico sob demanda. Figura 4C mostra as imagens microscópicas de fluorescência da córnea, esclera e retina com seção criogênica. A genisteína liberada pelas lentes de contato inteligentes parecia ser efetivamente liberada através da córnea para a retina. A fraca fluorescência na esclera revelou que a genisteína havia passado pela esclera com pouca absorção. No caso do controle, não foi observada fluorescência nos tecidos seccionados criogênicos de coelhos que usavam lentes de contato inteligentes sem genisteína ou lentes de contato inteligentes com genisteína sem gatilho elétrico para sua liberação (Fig. 4C, abaixo). A partir dos resultados, pudemos confirmar a viabilidade das lentes de contato inteligentes para a entrega de medicamentos terapêuticos oculares sob demanda, controlados eletricamente (tabela 1)

tabela 1 Comparação de várias lentes de contato inteligentes.

Uma câmera térmica infravermelha não mostrou alteração notável de temperatura no corpo da lente de contato inteligente nos olhos de coelho (Fig. 4D) No início, a temperatura da lente de contato inteligente era de 32,4 ° C, a da superfície ocular era de 34,4 ° C e a da bobina externa era de 32,0 ° C. Após 30 minutos de operação, a temperatura da lente de contato inteligente era de 33,8 ° C com aumento de temperatura de 1,4 ° C, a da superfície ocular era de 34,8 ° C com aumento de temperatura de 0,4 ° C e a da bobina externa foi de 29,7 ° C com uma diminuição de temperatura de 2,3 ° C. O ligeiro aumento de temperatura revelou a segurança térmica de nossas lentes de contato inteligentes.

Efeito terapêutico da genisteína liberada pelas lentes de contato inteligentes na retinopatia diabética

Os coelhos brancos da Nova Zelândia foram divididos em cinco grupos para avaliar o efeito terapêutico da genisteína liberada pelas lentes de contato inteligentes na retinopatia diabética em comparação com uma série de grupos controle e comparador. The left eyes of rabbits were treated with a topical eye drop of PBS as a negative control in group 1, a topical eye drop of genistein in group 2, intravitreal injection of genistein in group 3, and intravitreal injection of Avastin as a positive control in group 4. The right eyes of all groups were treated with smart contact lenses containing genistein (which collectively comprised group 5). Transmission electron microscopy (TEM) visualized the inhibitory effect of genistein released from the smart contact lens on the deformation of retinal vascular structure (Fig. 5A) The diabetic retinal vessels in Fig. 5A(iv) (left eye of group 4) and Fig. 5A(v) had a round shape surrounded by the thick vascular endothelial cell (EC) layers, which were comparable to that of the healthy rabbit (23) However, the vascular basement membrane appeared to be irregular and folded without the clear vascular EC layer in Fig. 5A(i) (left eye of group 1), reflecting increased vascular permeability and the blood-retinal barrier breakdown. At the Fig. 5A(ii) (left eye of group 2) and Fig. 5A(iii) (left eye of group 3), the vessels had a round shape, but the surrounding vascular EC layers were not as thick as those in Fig. 5A(iv and v).

Fig. 5 In vivo therapeutic effect of genistein released from the smart contact lens.

The eyes of diabetic rabbits were treated with (i) an eye drop of PBS (control), (ii) an eye drop of genistein, (iii) intravitreal injection of genistein, (iv) intravitreal injection of Avastin, and (v) genistein released from the smart contact lens. (AN) Electron micrographs of the retinal vessels. L, lumen of vessel; EC, endothelial cell; RBC, red blood cell. Scale bar, 1 μm. (B) Fluorescence angiograms of the retina (arrowheads, retinal vessels). Scale bar, 0.2 mm. (C) Histological analysis for the damage to the retinal pigment epithelium (RPE) and choroidal vessels (CVs) (arrowheads, damage in CV). Scale bar, 0.1 mm. (D) Apoptosis detection in retina by TUNEL assay. Scale bar, 0.1 mm. (AND) Merged images of immunohistochemistry staining for collagen type 4 (red) and PECAM-1 (green) with nuclear staining by 4′,6-diamidino-2-phenylindole (blue). Scale bar, 0.1 mm. (F) Fluorescence intensity of retinal choroidal neovascularization lesion quantified from the images of (B). (G) Fluorescence intensity of TUNEL assay quantified from the images of (D). (H) Immunochemical fluorescence intensity (E) of collagen type 4 (filled box) quantified from the images in fig. S9A (red) and PECAM-1 (dashed box) quantified from the images in fig. S9B (green)[[[[n = 3, *P < 0.05 and **P < 0.01 versus the control sample of (i)].

Figure 5B shows fluorescence angiograms for the morphology of retinal vessels. While no clear morphology of vessels was observed in Fig. 5B(i and ii), retinal vessels (arrowheads) with clear morphology were observed with the notably decreased retinal vascular permeability in Fig. 5B(iv and v). Fluorescence was observed throughout the retinal parenchyma owing to the increased vascular leakage after blood-retinal barrier breakdown, as quantified in Fig. 5F. At the Fig. 5B(iii), little fluorescence was observed with only a scant vasculature. The results of histological H&E analysis were consistent with those of TEM images and fluorescence angiograms (Fig. 5C) In addition, retinal cell death was validated by terminal deoxynucleotidyl transferase–mediated deoxyuridine triphosphate nick end labeling (TUNEL) assay in retinal cross-sectioned images (Fig. 5D) Fluorescence of TUNEL assay was quantified by ImageJ program. When the mean fluorescence intensity in Fig. 5D(i) was set to be 100%, the mean percentage of fluorescence intensity was 76.0% in Fig. 5D(ii), 69.0% in Fig. 5D(iii), 37.0% in Fig. 5D(iv), and 45.1% in Fig. 5D(v) (Fig. 5G) Furthermore, the immunohistochemical staining for collagen type 4 and platelet EC adhesion molecule–1 (PECAM-1) revealed the therapeutic effect of genistein released from the smart contact lens (Fig. 5E) The expression degree of collagen type 4 and PECAM-1 was lower in fig. S9(iv and v) than in fig. S9(i to iii) (Fig. 5H)


Smart electronic contact lens devices have been widely investigated for diagnostic applications, especially for continuous glucose monitoring and intraocular pressure monitoring. In addition, there have been many reports on the electrical and optical glucose sensing with improved sensitivity using various nanomaterials (2426) To improve the sensitivity, stability, and reproducibility, we immobilized GOx in the chitosan and PVA hydrogels together with BSA. PVA appeared to mitigate the problem of uneven coating and cracking by increasing the viscosity of the GOx mixture solution with the increased loss modulus (27) PVA was also reported to have a substantial effect on the sensitivity of glucose sensors (28., 29) As shown in Fig. 2, the glucose concentrations could be accurately measured from the electrical current change using our glucose sensor, showing the stability for the repeated glucose sensing even after storage for more than 63 days (Fig. 2E) and enabling the real-time continuous tear glucose monitoring in live rabbit eyes in comparison with the blood glucose sensing by a glucometer (Fig. 4B) In contrast, Parviz’s group used a model eye and Park’s group dropped glucose samples directly onto the rabbit eyes after wearing the smart contact lens for the assessment of their electrochemical glucose sensors, and there is no scientific journal report on in vivo glucose sensing of the Google lens (tabela 1)

Despite the intensive effort for the commercial development of Google lens, they recently reported that there was insufficient consistency in their measurements of the correlation between tear glucose and blood glucose concentrations to support the requirements of a medical device. The disappointing clinical results might be associated with the challenges of obtaining reliable tear glucose readings in the complex on-eye environment. Although the correlation between tear and blood glucose concentrations remains controversial, there are many reports supporting the strong correlation between them (15, 1719) As shown in Fig. 4B, we could perform real-time continuous tear glucose monitoring in live rabbit eyes, which was strongly correlated with the blood glucose concentrations. We believe that with proper calibration and baseline monitoring, the changes in glucose concentrations can be measured reliably for each patient using the smart contact lens. This is similar to that of the FDA-approved Triggerfish lens that measures changes in intraocular pressure rather than an absolute intraocular pressure.

On top of that, our smart contact lens has a unique function of ocular drug delivery. To date, a variety of drug-eluting contact lenses have been developed using biodegradable polymer nanoparticles and micelles to improve the efficiency of ocular drug delivery. However, there has been no report on smart contact lenses with an electrically controlled on-demand DDS, possibly due to the difficulty in the miniaturization of all these electronic components onto the small contact lens. Antiangiogenic genistein and the glucose level–controlling metformin could be delivered from the f-DDS on the smart contact lens (Figs. 3 and 4 and fig. S3). The released genistein could be delivered through the cornea to the retina as shown in Fig. 4, exhibiting the therapeutic effect on diabetic retinopathy. This smart contact lens for wireless biosensing and therapeutic drug delivery might pave a new avenue to ubiquitous health care for further theranostic applications. Although metformin has been commercialized as an oral drug, its therapeutic effects through various other delivery routes have been well documented, such as transdermal delivery (25) and ocular delivery (30, 31) Berstein (31) reported that metformin is not simply an oral drug and that it influences many reactions and processes such as proliferation, apoptosis, angiogenesis, and oxidative stress in cell lines and, given these findings, stated that it is very reasonable to target metformin for topical and ocular delivery applications.

Concerning the safety issue of the smart contact lens, the wireless energy transfer system should be carefully investigated because of the possible ocular damage by the generated heat of the smart contact lens. In this context, we measured the heat from operating the contact lens using an infrared thermal camera, which showed no notable temperature change in the smart contact lens on rabbit eyes (Fig. 4D) The only slight temperature increase revealed the thermal safety of our smart contact lens. Optical images and histological analyses of corneas in the eyes of New Zealand white rabbits also confirmed the safety of our smart contact lens (fig. S7). From all these results, we could confirm the preliminary safety of our smart contact lens for further applications. Moreover, the FDA approval for the clinical use of Triggerfish is an important supporting information on the safety of smart contact lenses.

In summary, a smart electrochemical contact lens has been successfully developed with a glucose biosensor and an f-DDS controlled by wireless power and remote communication systems for both diabetic diagnosis and therapy. We demonstrated the real-time biosensing of glucose concentrations in the tear and on-demand therapeutic drug delivery of genistein for the treatment of diabetic retinopathy in diabetic rabbit eyes. The ocular glucose biosensor uniformly coated with GOx immobilized in the cross-linked hydrogels of chitosan and PVA with BSA showed high sensitivity, linearity, and stability for the repeated applications after long-time storage for 63 days. The genistein delivered from the smart contact lens through the cornea to the retina showed a comparable therapeutic effect to that by the intravitreal injection of Avastin on diabetic retinopathy. This smart theranostic contact lens will be investigated further as a next-generation wearable device to achieve the real-time biosensing of ocular biomarkers and on-demand medication for ubiquitous health care applications to various ocular and other diseases.


Preparation of contact lens materials

Silicone contact lens hydrogels were prepared under nitrogen by the photocrosslinking of 2-hydroethylmethacrylate (HEMA), silicone-containing monomers of 3-(trimethoxysilyl)propyl methacrylate, 3-[tris(trimethylsiloxy)silyl]propyl methacrylate, and a cross-linker of ethyleneglycol dimethacrylate (EGDMA) for 15 min using a photoinitiator of Darocur TPO, diphenyl(2,4,6-trimethylbenzoyl)phosphine oxide. As a control, PHEMA contact lens hydrogels were prepared by mixing HEMA and EGDMA with the photoinitiator. To form a contact lens shape, the precursor solution was loaded on a polypyrrole mold under ultraviolet (UV) light at a wavelength of 254 nm for 8 min. Silicone and PHEMA hydrogel contact lenses were detached from the mold and surface-treated under oxygen plasma (OptiGlow ACE, Glow Research). The prepared contact lens was completely submerged in PBS at 37°C for a day before use.

Characterization of contact lens materials

ATR-FTIR (Tensor 27, Bruker) of dehydrated silicone hydrogel contact lens and lotrafilcon A was recorded over the 400 to 4000 cm−1 range. The transmittance of silicone and PHEMA hydrogel contact lenses was measured using a UV-visible spectrometer (SD-1000, Scinco) after soaking in PBS for 24 hours. Both samples were placed in quartz plates, and the transmittance was measured at the wavelength range of 250 to 1000 nm. The EWC was determined by weighing the dried contact lens (Wdry) and the hydrated contact lens with soaking in PBS for 24 hours (Wwet) The value of EWC was calculated as the percentage of the weight gain during hydration and dehydration using the following equation: EWC = (WwetWdry)/Wdry × 100 (32) The water contact angles on dried silicone and PHEMA contact lenses were measured in static mode by dropping 5 μl of water every 2 min (SmartDrop, FemtoFAB).

Fabrication of ocular glucose sensor

Three WE, CE, and RE in the glucose sensor were patterned with 20-nm-thick chromium (Cr) and 80-nm-thick Pt on a 0.23-μm-thick PET substrate using an electron beam evaporator. RE was additionally treated to form a 200-nm-thick silver (Ag) layer. For the long-term stability, all parts of the glucose sensor except WE, CE, and RE were passivated with Parylene C. For chlorination, the Ag layer was dipped in FeCl3 (1 M, Sigma-Aldrich) solution for 1 min. Then, PVA[2weight%(wt%)100000gmol[2weight%(wt%)100000gmol[2weight%(wt%)100000gmol[2weight%(wt%)100000gmol−1, Sigma-Aldrich]was dissolved in deionized water and chitosan (0.5 wt %, mid molecular weight, Sigma-Aldrich) was dissolved in acetic acid (1 M, Sigma-Aldrich) with vigorous stirring at 80°C for 12 hours. BSA (10 mg ml−1, Sigma-Aldrich) and GOx (50 mg ml−1, Sigma-Aldrich) were dissolved in 2 wt % of PVA solution, which was mixed with the chitosan solution. The mixed solution was stored in a desiccator to remove bubbles. To uniformly fabricate a GOx layer only on the WE, all areas of the sensor except WE were passivated with PDMS. Then, glucose sensors were treated with UV in the presence of ozone for 10 min. After removing PDMS, 1.8 μl of the prepared GOx mixture solution was dropped onto WE and dried in a desiccator. Last, 1.8 μl of glutaraldehyde (2 wt %, Sigma-Aldrich) was dropped on the GOx layer and dried slowly at 4°C.

In vitro electrical detection of glucose

In vitro electrical glucose measurements were conducted using a potentiostat (Ivium Tech. Co., AJ Eindhoven, The Netherlands) and a computer-controlled ADC (6030E, National Instruments). A 50-ml beaker was filled with 10 ml of PBS (1 M, pH 7.4). The glucose sensor was put into the beaker to dip the sensing area sufficiently in PBS. The glucose sensor detected the change of electrical current under a constant potential of 0.7 V versus Ag/AgCl for steady-state amperometric current responses. After stabilizing the glucose sensor, a high concentration of glucose solution (10,000 mg dl−1, Wako) was added in PBS to slowly change the glucose concentration in the beaker from 5 to 50 mg dl−1, and the change of current was monitored for the glucose quantification. To investigate the selectivity and specificity of the glucose sensor, the change of current was measured after adding the potentially interfering molecules such as A (0.1 M, Sigma-Aldrich), L (10 M, Sigma-Aldrich), and U (10 M, Sigma-Aldrich) in PBS. The long-term storage stability and the repeated usability of the glucose sensor were assessed at days 0, 21, 42, and 63 after fabricating the glucose sensors. The glucose sensors were stored at 20° to 25°C in 5 ml of sterilized PBS (1 M, pH 7.4), similar to the conventional contact lens storage condition.

Fabrication and characterization of f-DDS

On-demand f-DDS was prepared by the LLO process. First, hydrogenated amorphous silicon (a-Si:H) exfoliation and SiO2 buffer layers were grown by plasma-enhanced chemical vapor deposition. Anode and cathode electrodes of the f-DDS were covered with 10-nm-thick Ti, 80-nm-thick Au, and 10-nm-thick Ti by e-beam evaporation and lithography. The reservoirs were patterned with 100-μm-thick negative photoresists (SU8-5 and SU8-50) with dimensions of 500 μm by 500 μm. As a model drug, 25 nl of genistein (3 M, Sigma-Aldrich) or metformin (2 M) with rhodamine B (Sigma-Aldrich) dye was loaded in the reservoirs. Subsequently, drug-loaded reservoirs were sealed with a flexible PET film. The XeCl excimer laser was exposed on the back side of the glass substrate to separate the SU-8 drug reservoir on the PET film from the glass substrate. For the mechanical bending test, the entire f-DDS was bent with a bending radius in the range of 5 to 30 mm and the electrical current was measured with a probe station. The durability of the f-DDS was assessed by applying 1000 bending cycles at a fixed bending radius of 5 mm.

Characterization of f-DDS

The drug release in response to applied voltage was investigated by connecting anode and cathode electrodes with the probe station. The constant electrical potential of 1.8 V was applied between anode and cathode electrodes for 1 min. Rhodamine dye released from the reservoir was visualized by confocal microscopy (Leica) using the corresponding imaging software (FluoView). The excitation wavelength was 543 nm and the emission wavelength was in the range of 560 to 610 nm. The concentration of released genistein and metformin in PBS was quantified with a spectrofluorometer (Thermo Fisher Scientific) at excitation/emission wavelengths of 355/460 nm and 485/538 nm, respectively.

Fabrication of power transmission coils

To fit into a contact lens, a wireless power receiver composed of a copper (Cu) coil was prepared with a thickness of 0.1 mm and an outer diameter of 1.2 mm. PDMS was spin-coated on a glass substrate, attaching 0.1 mm of Cu foil (Sigma-Aldrich). After polymerization of PDMS in an oven at 70°C for 1 hour, the Cu foil was patterned by photolithography. The foil was wet-etched in 5 ml of ammonium persulfate solution (12 mg ml−1) for 6 hours and detached from the PDMS. Then, the Cu coil was rinsed with acetone, ethanol, and distilled water for 10 min with sonication, respectively. The power transmitting coil was fabricated using four-turned Cu wire (Sigma-Aldrich) with a thickness of 1 mm and an outer diameter of 5 cm.

Power transmission efficiency measurement

The wireless power transmission system consisted of a Cu power transmitter coil, a Cu power receiver coil in a contact lens, a function generator (AFG 3101, Tektronix), a commercial power amplifier module (MAX 7060), and an ASIC chip. The power amplifier module was used to supply sufficient power to the ASIC chip. The transmitter coil transferred the power to the receiver coil by resonant inductive coupling. The receiver coil embedded in the contact lens was aligned in parallel to the transmitter coil with a distance from 0 to 4 cm to measure its efficiency. The efficiency of wireless power transmission between two coils was measured by using a network analyzer (N5230A, Agilent).

Design and fabrication of the ASIC chip

The ASIC chip is custom-built by multiwafer process fabrication. The ASIC chip was fabricated by Taiwan Semiconductor Manufacturing Company using a 180-nm complementary metal-oxide semiconductor (CMOS) process. The PMU rectified incoming alternating current (ac) energy from the coil to direct current (dc) supply voltage and generated various regulated voltages for other subunits. An RCU transmitted data through 433-MHz on-off keying modulation. A reference clock generator (CLKREF) was implemented with a relaxation oscillator for the system timing. A potentiostat with three nodes (WE, RE, and CE) was integrated into the ASIC chip by Au flip-chip bonding. The potentiostat applied a voltage bias of 1.2 V on the RE and 1.85 V on the WE using an operational amplifier with negative feedback. The change of electrical current was monitored in real time by dropping the glucose sample solution. An integrated ADC received the current input from the potentiostat and converted it to a 15-bit digital output code (33) The output codes were then externally transmitted through the ISM frequency band of 433 MHz using the RCU. The current sensing performance of ΔΣ ADC was measured by applying current input from a current supplier (B2961A, Agilent). To suppress the effect of large noise from the equipment, software-based filtering was applied to the measured digital codes. The RF receiver module passed the received data to the AVR, and the AVR communicated with a PC using an RS-232 protocol. The software decoded the data packets and displayed the raw data to the PC.

Power management of the ASIC chip

The PMU wirelessly received AC power and converted it into DC with a MOS-based rectifier, generating the external rectified voltage (VEXT) A bandgap reference circuit generated a reference voltage of 1.2 V, which was up-converted to 1.85 V and buffered with a regulator to provide an internal supply voltage (VINT), driving overall control logic blocks of the ASIC chip. For controlled drug delivery, anode and cathode electrodes in the f-DDS were connected to the PMU that selectively operated the f-DDS according to control commands received from the external reader.

Remote communication system

The RCU consisted of a 433-MHz tuned inductor-capacitor (LC) transmitter and its control logics. Control logics serialized the ADC output and patched a predefined header to define the packet boundary. The carrier frequency was determined by internal capacitors with an external loop antenna (L). Data modulation was performed by controlling the impedance change of the LC transmitter that could be observed by the external reader. An ASK receiver in the reader demodulated the impedance change, recovering transmitted data from the ASIC chip. The remote telemetry was formed with the ASIC chip, a receiver module, an AVR (Atmega-128), and data processing software written in Java.

Overall fabrication of the integrated smart contact lens

Because of the restriction to the ocular field of vision, a power receiver coil, a biosensor, and an f-DDS were fabricated on the peripheral area of a contact lens. The Cu power receiver coil was attached onto the ultrathin PET film (25 μm) with f-DDS using adhesive PDMS. The ASIC chip was implemented through the standard 0.18-μm CMOS process and diced into dimensions of 1.5 mm by 1.5 mm by 0.2 mm by chemical polishing and mechanical sawing. Afterward, the diced ASIC chip was attached, and WE, CE, and RE of the biosensor were deposited on the PET substrate. The power receiver coil, electrodes of the biosensor, and f-DDS were electrically connected with the ASIC chip using Au flip-chip bonding. For insulation and waterproofing, all devices on the PET substrate were coated with Parylene C and PDMS except for the sensing channel of the biosensor and the exposed electrodes of the f-DDS. Last, the integrated devices were molded into silicone hydrogels to fabricate a smart contact lens.

Preparation of diabetic retinopathy model rabbits

For in vivo glucose monitoring and diabetic retinopathy treatment, streptozotocin (STZ)–induced diabetic rabbit models were prepared by single injection of STZ (65 mg·kg−1) (1% STZ solution, diluted with 0.1 M citrate buffer, pH 4.4) to New Zealand white rabbits (2.0 kg) via the ear vein after fasting for 12 hours. After STZ injection, the rabbits with a plasma glucose concentration higher than 140 mg dl−1 were considered diabetic.

In vivo electrical detection of tear glucose levels

For in vivo real-time glucose monitoring, smart contact lenses were worn on each diabetic rabbit’s eye, and the power transmitter coil was placed outside the eyes to transfer the wireless power to the receiver coil on the smart contact lens. The voltage was applied onto the glucose sensor in a pulsed manner, and the electrical measurement of glucose concentration was performed in real time with remote data transmission. Before 15 min of wireless tear glucose sensing, 2 U of insulin was injected to decrease the blood glucose level. After 5 min, ketamine was injected into diabetic rabbits for anesthetization. PBS was dropped onto the diabetic rabbit’s eyes, and the smart contact lens was worn on the eye to start the wireless tear glucose monitoring.

Analysis of genistein penetration in vivo

The penetration of genistein released from smart contact lenses into eyes was investigated after positioning of the genistein-loaded smart contact lens onto rabbit eyes with wireless powering to operate the f-DDS. After 1 hour, the penetration of genistein was confirmed by fluorescence microscopic analysis in cryo-sectioned tissue of cornea, sclera, and retina using a fluorescence microscope (Fluoroskan Ascent, Thermo Fisher Scientific) at an excitation wavelength of 355 nm and an emission wavelength of 460 nm.

Electron microscopy and histological analysis

For the electron microscopic analysis of retinal blood vessels, the retinas were enucleated and fixed in 4 wt % glutaraldehyde and 1 wt % osmium tetroxide solution. The samples were dehydrated with ethanol and sectioned to observe the cross section of retinal blood vessels by TEM (JEM-1010, JEOL). Histological analysis was performed with H&E staining of retinas fixed in 4% (w/v) paraformaldehyde for 24 hours.

In vivo treatment of diabetic retinopathy

The treatment of diabetic retinopathy using the smart contact lens was performed for 5 days on the right eyes of rabbits in five groups. The electrical power was wirelessly transmitted at a frequency of about 433 MHz using a power transmission coil to operate the f-DDS. As a control, an eye drop of PBS (0.05 ml, group 1), an eye drop of genistein (0.4 mM, 0.05 ml, group 2), and intravitreal injection of genistein (0.4 mM, 0.05 ml, group 3) were performed on the left eyes of each rabbit at the same time with the smart contact lens treatment. In addition, intravitreal injection of Avastin (0.05 ml, group 4) was performed on the left eye of rabbits. The right eyes of all groups were treated with smart contact lenses containing genistein (group 5).

Whole-mount retina immunofluorescence staining

The rabbit eyes were placed in 4% paraformaldehyde for 45 min. After fixation, retinas were dissected and flattened by applying curve-relieving cuts. The retinas were then fixed for an additional 1 hour. The retinas were washed twice with PBS and incubated with a 0.2% solution of Triton X-100 in PBS at room temperature for 1 hour. Last, vessels were stained with fluorescein isothiocyanate–labeled lectin from Bandeiraea simplicifolia (1:100, Sigma-Aldrich).

Immunostaining analysis

TUNEL assay was performed following the standard protocol. The immunostaining of collagen type IV and PECAM-1 was performed according to the manufacturer’s protocols. The following antibodies were used: PECAM-1 antibody (sc-18916, Santa Cruz Biotechnology) and collagen type IV antibody (ab6586, Abcam). Nuclei were counterstained with 4′,6-diamidino-2-phenylindole. The images of vasculature were obtained at ×10 magnification. All fluorescence intensity was quantified by ImageJ program.

Study approval

All experiments were performed in accordance with the Association for Research in Vision and Ophthalmology Statement for the Use of Animals in Ophthalmic and Vision Research. The animal protocol was approved by the Institutional Animal Care and Use Committee at the College of Medicine, the Catholic University of Korea.

Statistical analysis

We performed one-sided statistical analyses using Student’s t tests or one-way analysis of variance (ANOVA) with Bonferroni posttest. P < 0.05 was considered statistically notable. The quantification of fluorescence images was performed using ImageJ program. All data points were derived from three or more biological or technical replicates, as indicated for each experiment.

This is an open-access article distributed under the terms of the Creative Commons Attribution-NonCommercial license, which permits use, distribution, and reproduction in any medium, so long as the resultant use is not for commercial advantage and provided the original work is properly cited.

Acknowledgments: Funding: This work was financially supported by Samsung Science & Technology Foundation (SRFC-IT1401-03) in Korea. This research was supported by the Center for Advanced Soft-Electronics (Global Frontier Project, CASE-2015M3A6A5072945) and the Basic Science Research Program (2017R1E1A1A03070458 and 2020R1A2C3014070) of the National Research Foundation (NRF) funded by the Ministry of Science and ICT, Korea. This work was also supported by the World Class 300 Project (S2482887) of the Small and Medium Business Administration (SMBA), Korea. D.M. was supported by the National Eye Institute (K08EY028176 and P30-EY026877) and the Research to Prevent Blindness Foundation. Author contributions: S.K.H. conceived and supervised the project, designed experiments, interpreted data, and wrote the manuscript. D.H.K. and S.-K.K. performed experiments, collected samples, analyzed and interpreted data, and wrote the manuscript. J.K., C.J., B.H.M., K.J.L., E.K., and S.H.Y. contributed to preparing and designing the smart contact lens. G.-H.L., S.S., J.-Y.S., and Z.B. contributed to designing and performing the electrical experiments. J.W.M. and C.J. contributed to designing and performing the animal experiments. D.M. contributed to analyzing and interpreting the data and revising the manuscript. All authors contributed to critical reading and revision of this manuscript. Competing interests: S.H.Y., E.K., K.J.L., D.H.K., C.-K.J., and S.K.H. are inventors on a patent related to this work filed by Harvard Medical School and PHI Biomed Co. (no. US 2016/0223842A1, filed 4 August 2016). K.J.L., B.H.M., D.H.K., and S.K.H. are inventors of a patent related to this work filed by POSTECH and PHI Biomed Co. [no. US 10,399,291B2, filed 3 September 2019, registered in the United States and Korea (10-2016-0050139), and applied in Japan (2018-507476) and Europe (16783461.3)]. The authors declare that they have no other competing interests. Data and materials availability: All data needed to evaluate the conclusions in the paper are present in the paper and/or the Supplementary Materials. Additional data related to this paper may be requested from the authors.

Paula Fonseca