A closed microfluidic network for strain sensing, embedded within a contact lens to monitor intraocular pressure.

Abstract

To provide a closed microfluidic network for strain sensing embedded in a contact lens to monitor intraocular pressure.SOLUTION: The device has a contact lens and a closed microfluidic network embedded in the contact lens. The network has a volume that is sensitive to an applied strain. The network distinguishes (i) a gas reservoir containing a gas, (ii) a liquid reservoir containing a liquid that changes volume when the strain is applied, and (iii) a sensing channel capable of holding the liquid within the sensing channel. The sensing channel connects to the gas reservoir on one end and to the liquid reservoir on another end. The sensing channel establishes a liquid-gas equilibrium pressure interface and equilibrium within the sensing channel, which can fluidically change as a response to curvature radius variations on a cornea or as a response to mechanical stretching and release of the cornea.SELECTED DRAWING: Figure 1

Description

[Technical Field]

[0001] The present invention relates to devices, systems, and methods for monitoring intraocular pressure. In particular, the present invention relates to a microfluidic network design for a strain sensor that functions based on the mechanical amplification of the volume of microfluidic channels for monitoring intraocular pressure. [Background technology]

[0002] Glaucoma is a neurodegenerative disease that causes irreversible damage to the optic nerve of the eye, and therefore loss of vision. Continuous and long-term monitoring of intraocular pressure (IOP) is important for the management of glaucoma.

[0003] Reducing IOP is the only known method to slow and / or halt the progression of glaucoma. For every 1 mmHg reduction in IOP, the risk of nerve damage is estimated to be reduced by 11%. While drug therapy is commonly used to reduce IOP, there are significant challenges that need to be addressed to improve the effectiveness of glaucoma treatment. Most importantly, about 50% of patients discontinue medication use after 6 months for various reasons. Sustained, long-term IOP monitoring with the ability to measure drug efficacy can help patients stay in compliance with glaucoma management and can assist physicians in this regard. Furthermore, in recent years, diurnal variation in IOP has been established as another risk factor for glaucoma, which further increases the importance of continuous monitoring.

[0004] Current technologies available for IOP measurement are either non-sustaining (Goldmann Applanation Tonometry), persistent but temporary (Sensimed Triggerfish), or persistent but invasive (Implantable Sensors). Self-administered tonometry devices (e.g., Icare) can provide long-term data, and while non-invasive, they are still uncomfortable for the patient to the extent that topical anesthetics may be required. Furthermore, the results obtained by self-administered tonometry have been found to be user-dependent.

[0005] Approaches for continuous IOP measurement using telemetry have been developed and tested in animal models. Among these approaches, contact lens-based monitoring techniques are attractive because they are non-invasive. One contact lens system (Triggerfish from Sensimed AG) measures minute changes in corneal curvature by using a contact lens equipped with an electrical strain sensor, antenna, and microchip used to process and transmit signals wirelessly. This technique requires the patient to wear a receiver on their waist for data transmission and power transfer. Due to the thick silicone contact lens (center thickness 580 μm), it is not as comfortable as contact lenses used daily, and mild to moderate adverse reactions have been reported in up to 80% of patients. The requirement for trained personnel, as well as the discomfort and high cost associated with this contact lens platform, hinder its use in long-term monitoring applications, although it allows for testing over a single 24-hour cycle. For this reason, Triggerfish is found to be more suitable for determining IOP changes on a daily scale. However, IOP changes as a response to drugs are on a timescale of several weeks. Similarly, IOP changes in response to lifestyle modifications will also occur over time scales longer than 24 hours. Therefore, there is a need for continuous-wear contact lens sensors that can monitor IOP fluctuations over the long term in order to determine drug efficacy and reduce the number of hospital visits patients need to make for regular IOP measurements.

[0006] Other embodiments of contact lens sensors are based on measuring changes in electrical resistance, inductance, and capacitance in response to pressure-induced strain. In these embodiments, the sensor response is typically detected remotely by measuring changes in resonant frequency using an external reader coil, or by Bluetooth® connectivity. Electrical measurements require conductive components inside the lens, which are typically opaque and impermeable.

[0007] More recently, Kim et al. have used graphene-Ag nanowires to address the problem of electrode transparency (J. Kim et al., "Wearable smart sensor systems integrated on soft contact lenses for wireless ocular diagnostics", Nature Communications, vol. 8, Apr 2017, Art. no. 14997). The first requirement for contact lenses with long-term wearability is high breathability to prevent hypoxia. Unfortunately, the conductive components required by electrical sensors are impermeable to gases. Metals have gas permeability 8 to 10 orders of magnitude lower than soft materials, which causes mild adverse reactions in human clinical trials, even when electrically sensing-based contact lenses are used over a single 24-hour period. Another requirement for long-term wear is comfort, which is achieved by producing contact lenses with high water content and thinness (<200 micrometers). Electrically sensing methods are sensitive to the hydration level of the contact lens. Therefore, contact lens electrical sensors are made from silicone with a very low water content, instead of standard silicone / hydrogel material. This reduces the comfort of the contact lens. There are three main reasons for the sensitivity to hydration levels. Firstly, the expansion of the hydrogel due to hydration induces strain, and therefore this is a source of error in measurement. Secondly, friction between the contact lens and the cornea can be sensitive to hydration levels and therefore affects sensitivity. Finally, electrical components are affected by humidity and should therefore be isolated by using a sealant material such as parylene-C.

[0008] This invention provides a technique for measuring IOP that advances the field of the art and eliminates at least some of the current problems or concerns. [Prior art documents] [Non-patent literature]

[0009] [Non-Patent Document 1] J. Kim et al., “Wearable smart sensor systems integrated on soft contact lenses for wireless ocular diagnostics”, Nature Communications, vol. 8, Apr 2017, Art. no. 14997 [Overview of the project] [Means for solving the problem]

[0010] This invention relates to a strain sensor that uses microfluidic principles and is integrated with a contact lens for IOP measurement. The material used in this invention is low-cost, transparent, breathable, and flexible. A method for embedding a microfluidic strain sensor within a silicone contact lens is provided. The microfluidic contact lens sensor (miLenS) enables patients to measure their own IOP and better manage glaucoma.

[0011] The microfluidic contact lens sensor can measure increases and decreases in IOP (intraocular pressure) throughout a patient's lifetime due to internal factors (i.e., metabolism, blinking, and saccadic eye movements) as well as external factors (i.e., medications, diet, lifestyle, etc.). Measurements are performed at the patient's discretion (or automatically), and readings are performed by a smartphone camera (or a wearable camera for automated measurement). This enables home monitoring and continuous data recording. The data is then transmitted directly to the healthcare provider's database, which allows patients and physicians to monitor IOP fluctuations. Aspects of our technology are listed below.

[0012] 1) miLenS will be constructed using a hybrid material system in which a narrow microfluidic sensing region (a ring as small as 0.1 mm in width around the periphery of the miLenS) is embedded within a silicone or silicone / hydrogel contact lens material. The microfluidic sensing channel will be made from a transparent, flexible, and oleophobic material. The sensing material will have a permeability 6 to 10 orders of magnitude higher than that of the electronic components.

[0013] 2) Microfluidic sensing techniques have no actively controlled components and function solely based on the principles of fluid physics. miLenS has no electrical components whatsoever (no power). This makes it a low-cost device. In addition, it offers easier usability by eliminating the cumbersome peripheral components (e.g., antennas, microchips, etc.) required for data transmission, reception, and recording in wearable electronic sensors.

[0014] 3) The sensor is sensitive to strain and responds to changes in the radius of curvature of the cornea, but has low sensitivity to forces directly applied by the eyelid or forces resulting from the hydration of the contact lens material. The sensor designed by the inventors has low rigidity in the lateral direction (i.e., the microfluidic device is thin and has a low modulus of elasticity) and high rigidity in the radial direction (i.e., the microfluidic network channel has a small width), which will reduce sensitivity to external forces (e.g., blinking, eye rubbing).

[0015] 4) miLenS enables reading using a smartphone camera and optical adapter. This will provide measurements at discrete points in time. In one variation, a wearable camera capable of tracking sensor responses can also be used for continuous and automated measurements.

[0016] 5) Continuous data recorded using existing techniques show that IOP increases or decreases by approximately 5–15 mmHg per day and per hour, and by approximately 15–40 mmHg per second. The microfluidic network circuit designed by the inventors has the ability to filter out large increases or decreases that occur on short time scales due to blood pressure or muscle contraction. In this case, the sensor actually acts as a fluid low-pass filter, which responds only to changes that occur over several minutes or slower. In a similar manner, fluid components can be designed to register only rapid IOP changes. Sensors capable of measuring events occurring on different time scales can make a better estimate of the true IOP based on corneal radius of curvature measurements.

[0017] Microfluidic strain sensor-embedded contact lenses are convenient to use and possess continuous measurement capabilities. They require little training to perform measurements and will therefore be used as devices for home medicine. These will enable clinical research requiring the recording of long-term IOP data on large patient populations. Continuous recording and analysis of IOP will improve our understanding of neurodegenerative diseases and their relationship to pressure. In addition, this will be useful in improving the efficiency and efficacy of drugs used for glaucoma treatment. Therefore, miLenS technology offers a promising healthcare technology for better personalized treatment of glaucoma patients. These advantages listed above will potentially allow patients to use the sensors permanently and without the assistance of trained personnel.

[0018] In one embodiment, the present invention provides a microfluidic strain sensing device for monitoring intraocular pressure changes. The closed microfluidic network is transparent and / or oleophobic. The microfluidic strain sensing device has a contact lens and a closed microfluidic network embedded in the contact lens. The contact lens is a silicone contact lens, a hydrogel contact lens, or a combination thereof. The contact lens does not have any actively controlled components or electrical components.

[0019] The closed microfluidic network has a volume that is sensitive to axial strain. The closed microfluidic network is distinguished into (i) a gas reservoir containing gas, (ii) a liquid reservoir containing a liquid that changes volume when strain is induced, and (iii) a sensing channel capable of holding the liquid within the sensing channel. The sensing channel connects the gas reservoir on one end and the liquid reservoir on the other end. The sensing channel establishes a liquid-gas equilibrium pressure interface and equilibrium within the sensing channel that will vary fluidically in response to changes in the radius of curvature on the cornea or in response to mechanical stretching and release of the cornea. The liquid-gas equilibrium pressure interface and equilibrium are used to measure intraocular pressure.

[0020] The liquid reservoir forms at least one ring, and the air reservoir is positioned inside or outside at least one ring. In each case, the liquid reservoir volume is very sensitive to radial forces on the eye and tangential forces on the eye with respect to wearing the contact lens. The liquid reservoir has a high rigidity in the radial direction and / or a channel width smaller than the rigidity and / or microfluidic channel wall thickness in the tangential direction, resulting in the liquid reservoir being less sensitive to external forces.

[0021] In one embodiment, the liquid reservoir has one or more chambers. These chambers may have concentric rings. These chambers may also have concentric rings that are interconnected at one or more locations. These chambers may also have concentric rings, and the sensitivity increases as the number of concentric rings increases.

[0022] In one embodiment, the surface of the liquid reservoir may be patterned. The surface of the liquid reservoir ceiling may have a convex shape, and the convex shape may curve towards the reservoir channel floor.

[0023] The sensing channel has a strain sensitivity of about 4.5 mm of interface movement per about 1% strain applied to the liquid reservoir. In one embodiment, the sensing channel has an inner diameter of about 1 - 10 mm. In another embodiment, the sensing channel has an inner diameter of 5 - 12 mm with a cross-sectional area of 10 -11 ~10 -8 m 2 and an inner diameter of 5 - 12 mm with a cross-sectional area of 10 The present invention provides, for example, the following. (Item 1) A microfluidic strain sensing device for monitoring intraocular pressure changes, (a) A contact lens, and (b) A closed microfluidic network embedded in the contact lens comprising, the closed microfluidic network has a volume that is sensitive to an applied strain, and the closed microfluidic network further comprises (i) A gas reservoir containing gas, and (ii) A liquid reservoir, wherein the liquid reservoir contains a liquid that changes volume when the strain is applied, (iii) The sensing channel capable of holding the liquid within the sensing channel comprising, the sensing channel connects the gas reservoir on one end and the liquid reservoir on the other end, The sensing channel establishes a liquid-gas equilibrium pressure interface and equilibrium within the sensing channel, which will change fluidly in response to changes in the radius of curvature on the cornea, or in response to mechanical stretching and release of the cornea. The liquid-gas equilibrium pressure interface and equilibrium are microfluidic strain sensing devices used to measure the intraocular pressure. (Item 2) The intraocular pressure monitoring device according to item 1, wherein the liquid reservoir forms at least one ring, and the air reservoir is located inside the at least one ring. (Item 3) The intraocular pressure monitoring device according to item 1, wherein the liquid reservoir volume is much more sensitive to tangential forces on the eye than to radial forces on the eye on which the contact lens is worn. (Item 4) The intraocular pressure monitoring device according to item 1, wherein the liquid reservoir has high rigidity in the radial direction and / or a smaller channel width compared to the rigidity in the tangential direction, resulting in the liquid reservoir being less sensitive to external forces. (Item 5) The intraocular pressure monitoring device described in item 1, wherein the contact lens is a silicone contact lens, a hydrogel contact lens, or a combination thereof. (Item 6) The intraocular pressure monitoring device according to item 1, wherein the sensing channel has a strain sensitivity of approximately 4.5 mm interface movement per approximately 1% strain applied to the liquid reservoir. (Item 7) The sensing channel has an inner diameter of approximately 1 to 10 mm, as described in item 1 of the intraocular pressure monitoring device. (Item 8) The sensing channel is 10 -11 ~10 -8 m 2 An intraocular pressure monitoring device as described in item 1, having an inner diameter of 5 to 12 mm with a cross-sectional area. (Item 9) The intraocular pressure monitoring device according to item 1, wherein the liquid reservoir has one or more chambers. (Item 10) The intraocular pressure monitoring device according to item 1, wherein the liquid reservoir has one or more chambers with concentric rings. (Item 11) The intraocular pressure monitoring device according to item 1, wherein the liquid reservoir has one or more chambers with a concentric ring, and the concentric ring is connected at one or more locations. (Item 12) The intraocular pressure monitoring device according to item 1, wherein the liquid reservoir has one or more chambers with concentric rings, and the sensitivity increases as the number of concentric rings increases. (Item 13) The surface of the liquid reservoir is patterned, as described in item 1 of the intraocular pressure monitoring device. (Item 14) The intraocular pressure monitoring device according to item 1, wherein the surface of the liquid reservoir ceiling has a convex shape, and the convex shape curves toward the reservoir channel floor. (Item 15) The intraocular pressure monitoring device according to item 1, wherein the contact lens does not have any actively controlled or electrical components. (Item 16) The closed microfluidic network is transparent, as described in item 1, for the intraocular pressure monitoring device. (Item 17) The closed-type microfluidic network is oleophobic, as described in item 1, for the intraocular pressure monitoring device. [Brief explanation of the drawing]

[0024] [Figure 1] Figure 1 shows a workflow of a pressure-monitoring-based miLens device according to an exemplary embodiment of the present invention. [Figure 2A] Figure 2A shows an image of a sensor with a thickness of only 100 micrometers according to an exemplary embodiment of the present invention. The small drops on each side are Norland Optical Adhesive (NOA), which is used to seal the sensor and can be fabricated with a thickness of less than 20 micrometers. [Figure 2B] Figure 2B shows an image of the sensor (final thickness of 300 micrometers) after it has been embedded in a contact lens, according to an exemplary embodiment of the present invention. [Figure 3] Figure 3 shows a top view of a closed-type system sensor with multiple ring-shaped liquid reservoirs embedded in a contact lens, according to an exemplary embodiment of the present invention. [Figure 4] Figure 4 shows side views of the individual behavior of sensors when stretched under tangential forces, such as the multi-chamber liquid reservoir sensor A) versus the single-chamber liquid reservoir sensor B) and A*) and B*) according to exemplary embodiments of the present invention. 410-A and 410-B indicate possible stretching points under sensor stretching. The sensor needs to be made from a soft material that reduces stiffness in both directions. The sensor needs to be thin. Figure 4 illustrates this, and essentially, the microfluidic channel ceiling thickness t1 and floor thickness t2 need to be small (<20 μm). This also reduces stiffness in both directions. The reservoir width w needs to be small (<100 μm). This does not affect tangential stiffness but is important in increasing the radial stiffness of the microfluidic channel and improving sensor performance. [Figure 5] Figure 5 shows a top view of a single-ring liquid reservoir versus a three-ring reservoir according to an exemplary embodiment of the present invention. The circled area relating to the three rings shows an enlarged view of the rings. [Figure 6] Figure 6 shows the pressure responses of three different sensor types, namely, one, two, and five reservoirs, according to exemplary embodiments of the present invention. The ring height, width, and spacing are 100 micrometers. The slope value is the sensitivity and is shown below the corresponding curve in mm / mmHg units. For each curve, the mean and standard deviation of at least three measurements are used. [Figure 7]Figure 7 shows the sensitivity dependence on the number of reservoir rings for three different ring widths according to an exemplary embodiment of the present invention. Multiple data points for some of the ring numbers were acquired using sensors machined at different time points with identical parameters, and the increase or decrease in sensitivity values ​​is a result of the differences in machining. Sensitivity is linearly dependent on the number of rings with widths of 50 and 100 micrometers, but is not significantly affected by the number of rings with a width of 200 micrometers. [Figure 8] Figure 8 shows a side view of the placement of the miLenS on the cornea and the position of the liquid reservoir according to an exemplary embodiment of the present invention. The insets show enlarged views of the liquid reservoirs and the forces acting thereon, inset a) showing a single broad liquid reservoir compressed under a radial force, and inset b) showing a series of concentric circles as liquid reservoirs not compressed under the same force. [Figure 9] Figure 9 shows the sensitivity dependence on height for three different ring widths according to an exemplary embodiment of the present invention. Multiple data points for some of the heights were acquired using sensors machined at different time points with the same parameters, and the increase or decrease in sensitivity values ​​is a result of the differences in machining. Sensitivity depends linearly with reservoir height. The red data points 910 represent thicker chips (300 micrometers), which show a 50% reduction in sensitivity compared to their thinner (150 micrometer) counterparts 920. [Figure 10] Figure 10 shows an enlarged cross-section of an auxetic contact lens sensor and liquid reservoir according to an exemplary embodiment of the present invention. Unlike sensors with rectangular channels as shown in Figure 8, this channel has a curved upper layer. This upper layer is flattened when a tangential force is applied, as shown by our data and Consol simulations. [Figure 11] Figure 11 shows a sensor with a patterned reservoir ceiling having circular and linear convex shapes, according to an exemplary embodiment of the present invention. [Figure 12]Figure 12 shows a microscopic image on the left of a sensor with a linearly patterned liquid reservoir ceiling according to an exemplary embodiment of the present invention. On the right, a comparison of measured sensitivities of 29 micrometers / mmHg and 77 micrometers / mmHg for a flat ceiling sensor versus a curved ceiling (auxetic) device is shown, respectively. [Figure 13] Figure 13 shows a method for processing a sensor according to an exemplary embodiment of the present invention. A refers to UV treatment. B refers to plasma treatment (PDMS). C refers to APTES treatment. 1 refers to a glass slide, 2 refers to NOA65 (uncured), 3 refers to PDMS, and 4 refers to NOA65 (cured). In step 1, NOA65 is sandwiched between two PDMS-coated glass slides and UV-cured to create a 20-micrometer film. This is repeated twice. In step 2, NOA65 is dropped onto a mold and the 20-micrometer film from step 2 is plasma-treated. In step 3, the two layers from step 2 are sandwiched together and UV-cured. In step 4, the 70-micrometer layer from step 3 is plasma-treated. The 20-micrometer layer from step 1 is plasma-treated and APTES-treated. In step 5, the two layers from step 4 are sandwiched together. [Figure 14]Figure 14 shows a method for embedding a sensor in a contact lens according to an exemplary embodiment of the present invention. B refers to plasma treatment (PDMS). C refers to APTES treatment. D refers to curing (heat) treatment. 5 refers to a hemispherical mold for contact lens processing, 6 refers to the sensor, and 7 refers to the upper layer of the contact lens. In step 6, PDMS is poured onto the contact lens mold. It is then cured at 80 degrees Celsius and subjected to plasma and APTES treatment. The bottom surface of the sensor is plasma treated. In step 7, the bottom surface of the sensor is placed on the PDMS-coated contact lens mold. The sensor reservoir is filled with working fluid and sealed. In step 8, more PDMS is poured onto the sensor and cured at room temperature. In step 9, the contact lens is peeled off the surface of the mold. [Figure 15] Figure 15 shows the fabrication step of the ceiling layer of an auxetic microfluidic sensor according to an exemplary embodiment of the present invention. [Figure 16] Figure 16 shows a strain sensor for the biomechanics of cancer cells according to an exemplary embodiment of the present invention. A strain sensor is installed in the bottom channel, while cells are seeded in the upper channel. [Figure 17] Figure 17 shows a top and side view of the contact lens and shape location according to an exemplary embodiment of the present invention. In addition to the star shape in the top and side views, other exemplary shapes can also be provided. Combinations of these shapes can also be used. [Figure 18] Figure 18 shows the Consol results for an exemplary embodiment of the present invention, in which a 50-micrometer height and 50-micrometer width channel provides near-optimal sensitivity while maintaining a thin device. The star shape in Figure 18 indicates the optimal geometric parameters for maximum volume change (i.e., sensitivity) while maintaining the thinness of the device. [Modes for carrying out the invention]

[0025] Previously reported IOP measurement devices do not consider the directionality of the forces acting on the sensor. For example, the capacitance measurement-based sensor by Chen et al. (G.-Z. Chen, I.-S. Chan, LKK Leung, and DCC Lam, "Soft wearable contact lens sensor for continuous intraocular pressure monitoring", Medical Engineering & Physics, vol. 36, no. 9, pp. 1134-1139, Sep 2014) responds to radial forces applied to the lens, such as those caused by blinking. An ideal contact lens sensor should be sensitive only to strains applied as a result of changes in corneal radius and should not be affected by forces applied perpendicular to the lens (i.e., radial forces). With this in mind, the inventors used COMSOL simulations and experimental measurements to develop a strain sensor that is more sensitive to tangential forces than to radial forces on the eye. Embodiments of the present invention are based on microfluidic sensing for IOP measurement and such a desired strain sensor force response.

[0026] Figure 1 shows an example of the IOP self-monitoring technique workflow. miLenS is distinctly different from other sensors because the patient will be able to insert and remove it themselves, just like a regular contact lens. As IOP increases or decreases, the radius of curvature of the cornea changes (each 1 mmHg change in IOP causes a 4 μm change in the radius of curvature). In this technique, the fluid level in the sensor's microfluidic sensing channel will change in response to the change in the radius of curvature on the cornea. The sensor response will be detected using a smartphone camera equipped with an optical adapter and then converted into a pressure value by a smartphone app. This will eliminate security and health concerns associated with radio frequency or Bluetooth® data transfer methods. The inventors have demonstrated an IOP detection limit of 1 mmHg on the eye of an enucleated pig, which is sufficient for IOP monitoring applications.

[0027] Microfluidic circuits, similar to electronic circuits, can function as low or high-pass filters (electrical resistance and capacitance are replaced by the fluid resistance (R) and compliance (C) of the compressible material, respectively). The RC value will determine the time constant of the sensor response. Sensors with large RC values ​​will not respond to rapid changes but will be sensitive to slowly fluctuating diurnal variations. Sensors with small RC values ​​will have the ability to detect the effects of blinking and eye pulsation.

[0028] In one exemplary embodiment, a microfluidic strain sensor (Figure 2A) is integrated into a PDMS contact lens (Figure 2B) for wearable sensing applications. Referring to the sensor 300 with sensor material 302 in Figure 3-4, the sensor 300 is embedded within the contact lens 310 and is distinguished into a liquid reservoir 320 (which amplifies the displaced liquid volume and is shown in this embodiment as a liquid reservoir ring), a gas reservoir 330, and a sensing channel 340 connected to the liquid reservoir 320 at one end and to the gas reservoir 330 at the other end. First, the liquid reservoir 320 is filled with a working fluid such as oil using capillary action and then sealed. This creates a stable gas / liquid interface 350 within the sensing channel 340, forming a closed microfluidic network. Increasing or decreasing IOP changes the radius of curvature of the cornea, with each 1 mmHg increase in IOP increasing the radius of curvature of the cornea by 4 μm. This increases the liquid reservoir volume due to the strain applied to the elastic wall of the liquid reservoir. The increased reservoir volume creates a vacuum, shifting the gas / liquid position 350 in the sensing channel 340 toward the liquid reservoir 320. As the cross-sectional area of ​​the sensing channel is reduced, the linear liquid displacement required to adapt to the change in reservoir volume increases, and therefore, sensitivity improves.

[0029] Figure 5 shows top views of two exemplary designs of a microfluidic strain sensor, namely a single ring 510 versus three rings 520 with respect to a liquid reservoir. Increasing the vertical wall surface area of ​​the liquid reservoir increases the sensor's sensitivity to changes in IOP. This was tested in two ways: i) increasing the number of walls, and ii) increasing the height of the channel walls. First, the inventors designed and fabricated a sensor with multiple liquid reservoir rings, for example, indicated by 520, thus increasing the total wall surface area. Sensitivity results for different numbers of rings are presented in Figures 6-7. The inventors found that increasing the number of walls by adding more rings increased the sensitivity of the device in a linear manner. In contrast, the width of the reservoir had no significant effect on sensitivity. This phenomenon is a direct result of the interaction between tangential strain and radially force-induced collapse, as shown in Figure 8. To test the effect of reservoir wall height, the inventors constructed three types of sensors (50, 100, and 330 μm heights) and compared their sensitivities. As shown in Figure 9, as the reservoir height doubled, the sensitivity also doubled. When the inventors increased the reservoir height to 330 μm, the sensitivity also tripled (shown only for a 200 μm width), demonstrating the effect of vertical wall height. Figure 9 further shows the effect of sensor stiffness. When a 150 μm thick sensor is compared to a 300 μm thick one (shown by 100T and 330T), the thicker sensor has approximately 50% lower sensitivity.

[0030] In summary, the inventors experimentally examined a wide range of parameters to understand and optimize sensor performance. They fabricated sensors with various numbers of reservoirs (1 to 5), ring widths (w = 50 to 500 μm), reservoir heights (50, 100, 330 μm), tip thicknesses (130 μm, 300 μm), and different Young's moduli of approximately 1 MPa (PDMS) versus approximately 10 MPa (NOA 65) and approximately 100 MPa (NOA 61). The results of these sensitivity tests demonstrated the following: i) Increased liquid reservoir height increases sensitivity. ii) The inventors can improve sensitivity by adding more reservoirs to the design as needed (e.g., depending on the required continuous wear contact lens properties). iii) Stiffness (Young's modulus (E) × tip thickness (t) / width (w)) does not significantly alter sensitivity; however, it needs to be optimized to take into account other factors such as comfort and lens / corneal mechanical interaction.

[0031] Ausetic metamaterials for microfluidic strain sensing In another version of the sensor, the height of the microfluidic channel network increases in response to the applied tangential strain 1010. The volume increase is achieved by correcting Poisson's ratio through lithographic patterning of the elastomer sensor. Figure 10 shows the operating principle of the auxetic metamaterial for strain sensing through a cross-section of the contact lens sensor. The ceiling of the microfluidic channel has a convex shape, i.e., curves inward toward the channel, as shown. This is achieved by patterning a ceiling film with either a circular or linear pattern, as shown in Figure 11. These are the only patterns tested by the inventors, but other patterns can be used to obtain the same effect. As shown in Figure 10, when a tangential force is applied (i.e., due to a change in IOP), the ceiling deforms outward due to the convex ceiling, in contrast to the collapse observed when a flat ceiling is used. This deformation toward the front of the sensor, according to our COMSOL simulations as shown in U.S. Provisional Patent Application No. 62 / 556366 filed September 9, 2017 (Figure 14 therein, incorporated herein by reference), causes an increase in channel height, and therefore an amplification of liquid reservoir volume expansion. This amplification increases the sensitivity of the sensor.

[0032] Figure 12 shows an image of a liquid reservoir on an auxetic sensor with a linear pattern of convex structures on the ceiling on the left. Figure 12 shows an experimental sensitivity comparison between flat and curved (auxetic) devices on the right. The sensitivity increase is 2.5 times.

[0033] The biocompatible, electronically-free microfluidic mechanical metamaterials are highly sensitive, enabling the fabrication of reliable strain sensors. The tangential strain sensing method developed by the inventors is specific to intraocular pressure (IOP), as demonstrated by their experiments. The inventors used this approach to monitor IOP in pig eyes, demonstrating a detection limit of 1 mmHg (corresponding to 0.05% strain) and reliability over several weeks. The microfluidic strain sensor can measure eye strain resulting from shape changes in response to IOP within a clinically relevant range.

[0034] manufacturing The inventors constructed sensors using photolithography and soft lithography techniques. First, a polydimethylsiloxane (PDMS) soft mold was fabricated. As the sensor material, polyurethane-based Norland optical adhesive 65 (NOA65) was selected due to its transparency, flexibility, oleophobicity, and biocompatibility. Next, thin NOA65 films with the required characteristics were fabricated and bonded together to fabricate the sensor, as shown in Figure 13. For the purposes of the present invention, the inventors developed specific fabrication methods to construct extremely thin (approximately 100 μm) microfluidic devices. The gas permeability of the polyurethane used in the inventors' devices is 6 to 8 orders of magnitude lower than that of metals used in wearable electronic devices.

[0035] The inventors first cut a strain sensor into a desired shape and embedded a flat 100 μm strain sensor (Figure 2A) in a PDMS contact lens. Although the inventors constructed their sensors flat, they can also be constructed curved if a curved mold is used. The inventors developed a processing protocol that allows them to construct contact lenses with curvature radii of 8–15 mm and radii of 10–14 mm, as shown in Figure 2B. The inventors used dome-shaped plastic molds, into which they poured PDMS to obtain a 10–100 μm silicone film with a desired curvature radius, and then bonded their sensors to the silicone film by the (3-aminopropyl)triethoxysilane (APTES) chemical structure. More silicone was then poured to completely embed the sensors in the silicone. Further details are shown in Figure 14. Finally, the inventors cured the silicone overnight at room temperature and then cut out the lens using a circular puncher. The inventors developed a process and technique for constructing a sensor as thin as 50 μm in thickness, so that the entire contact lens sensor may be less than 150 μm thick.

[0036] Regarding the auxetic sensor version, the only difference in manufacturing is in step 4 of Figure 13, where the inventors used a patterned film instead of a flat film as the bottom layer. The patterning was done as shown in Figure 15.

[0037] Variations and modifications 1) The microfluidic strain sensing principle can be used for a wide range of medical applications where strain sensing is required. Biomedical applications other than glaucoma management can be listed as physical therapy monitoring (e.g., in hand joint injuries), speech recognition, fetal / infant monitoring, tremor disorders, robotics, etc.

[0038] 2) Microfluidic strain sensing can be used for biosensing and biochemical sensing. For example, it can be used to monitor and measure strain applied by cells on a surface. Mechanical cues play a crucial role in cellular processes such as cell differentiation, apoptosis, and motility. Cells sense and impart forces to the substrate on which they grow. Tumor cells generate more forces than normal cells. Shear stress, one of the representative physical cues, causes increased expression of genes activated by mechanical signals. Understanding the mechanical cues generated by cells will be important for understanding cancer progression induced by mutations in the cellular mechanical signaling pathways. Our strain sensors will provide direct monitoring of direct cancer cell signaling under exposure to different physical and mechanical cues. Thus, this will bring about a novel approach in cancer research. By using our sensors, new biomarkers may be discovered, as well as new drug therapies may be implemented. These devices will also be useful in several other conditions, including the modulation of synaptic plasticity in neurons, since force is one of the important factors in the progression of synaptic plasticity.

[0039] To understand the cellular response to different conditions, two layers of microfluidic channels can be constructed as shown in Figure 16. As cells grow, strain sensors on the bottom channel can be imaged. This will provide tissue rigidity. The upper channel can also be manipulated by applying different flow rates that vary the shear stress. In this design, the mechanical response of cells can be observed while they are being mechanically manipulated. This design will be used in biomarker and drug development.

[0040] Cancerous tissue exhibits increased rigidity as it progresses. On average, cancer cells will have four times the rigidity of normal tissue. Understanding the earlier rigidity of cancer cells will lead to earlier cancer detection. Strain sensors can be incorporated into patches that can be used externally on the skin. Specifically, this can be used in skin and breast cancer types. Such patches, with infrared beads embedded in microchannels, can be optimized and implanted in viscera in the case of ovarian, liver, and brain cancers. In particular, these patches can be implanted after severe tumor removal surgery to monitor cancer recurrence. Combining microfluidic-based strain sensors with flexible silicon electronics would enable multiplexing of three-dimensional soft tissue in vivo. These signals can be transmitted to a cloud-based system using Wi-Fi implantation technology. Overall, strain sensors incorporating advanced electronics would provide continuous monitoring of tissues with a high probability of cancer recurrence.

[0041] 3) miLenS can be manufactured by either i) embedding a strain sensor with the desired shape / size within the contact lens as described, or ii) directly patterning the desired topography onto the surface of the contact lens through soft lithography, in which the features on the mold are transmitted to the contact lens.

[0042] 4) Instead of using microfluidics, the distance between the fine geometric features on the contact lens can be directly measured. This distance will vary as a function of IOP. The geometric shape and pattern of these features should be carefully selected to maximize sensitivity to IOP. IOP will be measured based on imaging of a contact lens sensor (geoLenS) with geometric features similar to miLenS. FIG. 17 shows a top view and a side view of an exemplary geoLenS. The location and shape of the fine features to be used for IOP determination are shown. In addition to the star shape shown in the top view and side view, other exemplary shapes are also provided. Combinations of these shapes can also be used. In the top view, the radius of the contact lens is represented by r, and the value of r can be 0.5 to 1 cm. θ represents the angle between the features positioned at the periphery of the contact lens, which determines the number of features that will be angularly placed on the contact lens. θ can be 10° (36 features at the periphery) to 180° (2 features at the periphery). At least two features are required on the contact lens. d1, d2, d3,... d n represents the distance between consecutive features and can be 0.01 to 1 cm. The total distance d = d1 + d2 + d3 +... + d n should be smaller than r. The radius of curvature r of the contact lens shown in the side view c can be 0.5 to 1 cm. The characteristic width w of the features can be 0.001 to 0.5 cm.

[0043] As the IOP changes, the distance between the peripheral features, e.g., d1, changes and can be used as a measure of the IOP change. The distance between the central features, e.g., d2 or d3, or the width w of any feature can be used as a reference measurement because they do not change in response to the IOP. The distance between the opposing features at the periphery (total distance is 2d) changes the most in response to the IOP change. The distance of any one of the geoLenS features to a known feature of the eye (i.e., the boundary of the iris) can be detected as a measure of the IOP.

[0044] To test the feasibility of the concept proposed above, the inventors fabricated a contact lens made from PDMS with a thickness of approximately 250 μm. For testing, the inventors fabricated a realistic eye model made from PDMS as shown in U.S. Provisional Patent Application No. 62 / 556366, filed September 9, 2017 (left side of Figure 19 therein) (incorporated herein by reference).

[0045] The radius of curvature of the eye model changes by approximately 4 μm / mmHg (3 μm / mbar), which is very close to the behavior of the human eye.

[0046] As shown in U.S. Provisional Patent Application No. 62 / 556366, filed September 9, 2017 (right side of Figure 19 therein) (incorporated herein by reference), the inventors marked a contact lens and placed it on an eye model they had created. These marks served as probes, enabling the inventors to measure changes in distance between different locations on the contact lens as a function of applied pressure. The inventors applied four levels of pressure to the eye model, varying from 25 mbar to 100 mbar. The inventors sampled four locations on the contact lens (three distance measurements), and the distances between these locations were plotted as a function of applied pressure, as shown in U.S. Provisional Patent Application No. 62 / 556366, filed September 9, 2017 (Figure 29 therein) (incorporated herein by reference). Points located on the center of the contact lens were labeled as location "1," and the number of locations increased as the points moved further from the center (e.g., location "2"). The distance between different marked points (e.g., location "1" - location "2") was measured. In Figure 20, the blue, red, and green lines show the distance as a function of the applied pressure for locations 1-2, 2-4, and 4-6, respectively. The corresponding linear fits are also plotted. Overall, the preliminary findings indicate that the distance between different locations on geoLenS follows a linear function of the applied pressure, and this is within the measurable range.

[0047] 5) geoLenS features can be processed in the same way as miLenS, or they can simply be marked using ink.

[0048] 6) The miLenS reservoir channel may have a meandering shape instead of being circular.

[0049] 7) This device can be used as a temperature sensor because it is sensitive to the thermal expansion of materials.

[0050] 8) This device is less sensitive to changes in atmospheric pressure. This means it can be used in a vacuum, for example, in space applications.

[0051] 9) Images can be captured by a smartphone camera, a special handheld camera, or a wearable camera. Images can be captured directly in front of the eyes, at a 45° angle, at a 90° angle, or at any angle between 0° and 90°.

[0052] 10) The front or rear camera of the smartphone can be used for imaging.

[0053] 11) Images can be collected voluntarily by the patient or automatically when the patient is reading something on the phone.

[0054] 12) Image analysis can be performed by the camera's microprocessor, or it can be transferred to a main server for further processing.

[0055] 13) Patients can pay for subscriptions to cloud services such as data storage and analysis.

[0056] 14) The miLenS channel can be filled with a colored liquid to improve contrast with the iris or sclera.

[0057] Additional technical notes This invention relates to a closed microfluidic network for strain sensing applications. The device has strain sensitivity of 2 to 15 mm interface displacement per 1% strain, depending on the number of rings. Sensitivity can be further increased by increasing the number of rings. It is robust enough to withstand pressure changes applied over 24 hours and has a lifespan of several months. These features make it attractive for applications where extreme strain levels less than 0.1% need to be measured over time periods longer than 2 hours. The inventors embedded a sensor in a contact lens to monitor intraocular pressure (IOP). The required detection limit for IOP is 1 mmHg, which corresponds to a strain of 0.05%. The inventors achieved this strain detection limit by designing a liquid reservoir network containing multiple microfluidic channels as a liquid reservoir. The liquid reservoir network is connected to a sensing channel, which is connected to an air reservoir. These three components form a closed system. A sensor with these three components in one possible configuration is shown in Figure 3. Figure 3 is a top view of the sensor, showing it when embedded in a contact lens. The sensor is filled with a working fluid from the inlet using only capillary force. When the working fluid reaches the outlet, both the inlet and outlet are sealed with the sensor material, forming a closed system with a fixed fluid volume inside. At this point, the fluid fills the sensing channel to about half its total length, creating a liquid / air interface. Both the contact lens and the sensor are made from elastomers such as silicone and polyurethane, but can also be made from other materials such as silicone / hydrogel.

[0058] The sensor operates based on the volume amplification of a microfluidic liquid reservoir network when it is stretched under a tangential force. The operating principle of the sensor is illustrated in Figure 4. Here, for simplification, another configuration of the sensor components is used, in which they are linearly distributed instead of radially distributed. Side views of the sensor with a liquid reservoir that may have multiple chambers A) versus a single wide chamber B) are compared. When the sensor is stretched under a tangential force, the shapes of the sensor and its components change, respectively, A) * ) and B * ) changes as described. 410-A and 410-B are representations of possible stress regions on the sensor in the vicinity of the liquid reservoir. For reference, the total initial length of the sensor is shown as 1-1', the total initial liquid reservoir network width is shown as 2-2', and the initial position of the liquid-air interface is shown as 3. There are three significant mechanical changes that can occur when such a closed microfluidic network is stretched under tangential forces.

[0059] i) Elongation: A * ) and B * When these are compared to A) and B), respectively, it can be seen that the total sensor length (1-1') will increase due to the elongation. Similarly, the liquid reservoir network width (2-2') will also increase.

[0060] ii) Collapse: In the case of a single reservoir, the thin membrane above the liquid reservoir is B * As shown in ), it will collapse due to the induced stress and the low rigidity of the film. When multiple chambers with films of higher rigidity are used, collapse will occur. * As shown in ), it will not occur or will be significantly reduced.

[0061] iii) Increase in liquid reservoir volume and resulting vacuum effect: If the liquid reservoir width is stretched, its total volume will increase if film collapse can be prevented or significantly reduced. This volume increase means the liquid reservoir is B * If it consists of multiple chambers with small widths, as shown in ), it can be amplified. The amplification will be even greater when the auxetic pattern is created on a film of small reservoir chambers. As the volume of the liquid reservoir increases, this causes a vacuum effect, which pulls the liquid / air interface position (3) toward the liquid reservoir. The displacement of this interface in μm units per IOP change in mmHg units is defined as sensitivity. According to the literature, each 1 mmHg IOP change causes a strain of 0.05%. This strain causes a positional change of approximately 100 μm relative to the interface position.

[0062] Another factor to consider for maximum sensitivity is the Young's modulus (E) of the sensor material. Increasing E reduces comfort. When contact lenses with high lubricity are used for improved comfort, the contact friction between the cornea and the sensor / lens decreases, which will cause sliding and reduced sensitivity, especially with respect to high-E sensors. According to our experimental and simulation results, the optimal E is in the range of 0.2 to 10 MPa for maximum sensitivity and comfort. As E is reduced below 2 MPa, the width of the reservoir channel also needs to be reduced below 100 μm.

Claims

1. A microfluidic strain sensing system for monitoring changes in intraocular pressure, (a) Contact lenses and, (b) A microfluidic network embedded in the contact lens, wherein the microfluidic network has a volume that is sensitive to applied strain, The aforementioned microfluidic network is (i) A gas reservoir containing gas, (ii) A liquid reservoir containing a liquid whose volume changes when the strain is applied, (iii) A sensing channel, wherein the sensing channel is capable of holding the liquid within the sensing channel, and the sensing channel has the gas reservoir connected to one end and the liquid reservoir connected to the other end. Furthermore, (iv) The sensing channel establishes a fluidly changing liquid-gas equilibrium pressure interface and equilibrium within the sensing channel in response to changes in the radius of curvature on the cornea, or in response to mechanical stretching and release of the cornea. (v) The liquid-gas equilibrium pressure interface and equilibrium are used to measure the intraocular pressure, and the liquid reservoir has a microfluidic network having one or more chambers with concentric rings, (c) an imaging device configured to perform one or more of the following: to activate intraocular pressure measurement, or to interpret the measured intraocular pressure. A microfluidic strain sensing system equipped with the following features.

2. The system according to claim 1, wherein the concentric rings are connected at one or more locations.

3. The system according to claim 1, wherein the sensitivity of the volume of the liquid in the liquid reservoir increases as the number of concentric rings increases.

4. The system according to claim 1, wherein the imaging device comprises a smartphone camera.

5. The system according to claim 1, wherein the imaging device comprises a wearable camera.

6. The system according to claim 1, wherein the imaging device is configured to activate the automated measurement of intraocular pressure.

7. The system according to claim 1, wherein the imaging device is configured to activate the measurement of intraocular pressure at the discretion of one or more of the wearer or healthcare provider.

8. The system according to claim 1, wherein the imaging device is configured to transmit the measured intraocular pressure to a healthcare provider's database.

9. The system according to claim 1, wherein the imaging device is configured to continuously read the measured intraocular pressure.

10. The system according to claim 1, wherein the imaging device is configured to read the measured intraocular pressure at discrete time points.

11. The system according to claim 1, further comprising an optical adapter coupled to the imaging device, wherein the optical adapter is configured to be positioned between the imaging device and the eye, and is configured to assist the imaging device in capturing an image of the microfluidic network.

12. The system according to claim 1, wherein the imaging device is configured to convert the liquid-gas equilibrium pressure interface and equilibrium into intraocular pressure values.

13. The system according to claim 1, wherein the liquid reservoir forms at least one ring, and the gas reservoir is located inside the at least one ring.

14. The system according to claim 1, wherein the volume of the liquid reservoir is much more sensitive to tangential forces on the eye than to radial forces on the eye on which the contact lens is worn.

15. The system according to claim 1, wherein the contact lens comprises silicone, hydrogel, or a combination thereof.

16. The system according to claim 1, wherein the liquid reservoir has high rigidity in the radial direction and / or a smaller channel width compared to the rigidity in the tangential direction, resulting in the liquid reservoir being less sensitive to external forces.

17. The system according to claim 1, wherein the contact lens does not have any actively controlled or electrical components.

18. The system according to claim 1, wherein the surface of the ceiling of the liquid reservoir has a convex shape, and the convex shape is curved toward the reservoir channel floor.

19. The system according to claim 1, wherein the microfluidic network is closed.

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