Closed microfluidic network for strain sensing embedded in 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, and in particular to a microfluidic network design for a strain sensor that functions based on mechanical amplification of the volume of a microfluidic channel 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] IOP reduction is the only known method to slow and / or stop the progression of glaucoma. It is estimated that for every 1 mmHg of IOP reduction, the risk of nerve damage is reduced by 11%. Although 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 stop using medication after 6 months for various reasons. Continuous long-term IOP monitoring with the ability to measure drug efficacy can help patients remain compliant with glaucoma management and assist physicians in this regard. Furthermore, in recent years, diurnal variation of IOP has been established as another risk factor for glaucoma, which further increases the importance of continuous measurement.

[0004] Current technologies available for IOP measurement are either non-continuous (Goldmann Applanation Tonometry), or continuous but temporary (Sensimed Triggerfish), or continuous but invasive (implantable sensors). Self-tonometry devices (e.g., Icare) can provide long-term data, which is non-invasive, but still uncomfortable for the patient to the point that it may require topical anesthetics. Furthermore, the results obtained by self-tonometry have been found to be user-dependent.

[0005] Approaches for telemetric continuous IOP measurement 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 a contact lens equipped with an electrical strain sensor, an antenna, and a microchip used to process and wirelessly transmit signals. The technique requires the patient to wear a receiver on the waist for data transmission and power transfer. Due to the thick silicone contact lens (central 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 the contact lens platform prevent its use in long-term monitoring applications, but only allows testing over a single 24-hour period. 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 time scale of weeks. Similarly, IOP changes in response to certain lifestyle modifications will also be on longer than 24 hour time scales. Thus, there is a need for a continuous-wear contact lens sensor that can monitor IOP fluctuations over the long term to determine drug efficacy and reduce the number of office visits patients need to make for regular IOP measurements.

[0006] Other examples of contact lens sensors are based on measuring electrical resistance, inductance, and capacitance changes in response to pressure-induced strain. In these examples, the sensor response is typically detected remotely, by measuring resonant frequency changes using an external reader coil, or by Bluetooth® connectivity. Electrical measurements require a conductive component on the inside of the lens, which is typically not transparent or gas-permeable.

[0007] Recently, Kim et al. have used graphene-Ag-nanowires to address the electrode transparency issue (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 condition for contact lenses with long-term use capability is high gas permeability to prevent hypoxia. Unfavorably, the conductive components required by electrical sensors are impermeable to gases. Metals have 8-10 orders of magnitude lower gas permeability compared to soft materials, which causes mild adverse reactions in human clinical trials even when electrical sensing-based contact lenses are used over a single 24-hour period. The other condition for long-term use is comfort, which is achieved by creating contact lenses with high water content and thin (<200 micrometers). Electrical sensing methods are sensitive to the hydration level of the contact lenses. Therefore, the contact lens electrical sensor is 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 level. First, the expansion of the hydrogel due to hydration induces distortion and thus this is a source of error in the measurement. Second, the friction between the contact lens and the cornea may be sensitive to the hydration level and thus affects the sensitivity. Finally, the electrical components are affected by humidity and therefore should be isolated by using a sealant material such as parylene-c.

[0008] The present invention advances the art and provides techniques for measuring IOP that eliminate 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 Summary of the Invention [Means for solving the problem]

[0010] The present invention relates to a strain sensor using microfluidic principles that is integrated with contact lenses for IOP measurement. The materials used in the present invention are low-cost, transparent, gas-permeable, and flexible. A method is provided for embedding a microfluidic strain sensor in a silicone contact lens. The microfluidic contact lens sensor (miLenS) allows patients to measure their own IOP and better manage glaucoma.

[0011] The microfluidic contact lens sensor is capable of measuring IOP increase and decrease due to internal factors (i.e. metabolism, eye blinking, and saccadic eye movements) and external factors (i.e. drugs, diet, lifestyle, etc.) during the patient's lifetime. Measurements will be taken at the patient's discretion (or automatically) and readouts will be achieved by a smartphone camera (or by a wearable camera for automated measurements). This allows for at-home monitoring and continuous data recording. Data is then sent directly to a healthcare provider's database, which allows patients and physicians to monitor IOP fluctuations. Aspects of our technology are listed as follows:

[0012] 1) miLenS will be constructed using a hybrid material system where a narrow microfluidic sensing area (a ring as small as 0.1 mm wide around the periphery of miLenS) is embedded within a silicone or silicone / hydrogel contact lens material. The microfluidic sensing channel will be made from a transparent, soft, and oleophobic material. The sensing material will be 6-10 orders of magnitude more permeable compared to the electronic components.

[0013] 2) Microfluidic sensing techniques do not have any actively controlled components and work only on the principles of fluid physics. miLenS is devoid of any electrical components (powerless). It is a low-cost device. In addition, it offers easier usability by eliminating the cumbersome peripheral components (e.g., antennas, microchips, etc.) for data transmission, reception, and recording that are required 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 applied directly by the eyelid or forces due to hydration of the contact lens material. The sensor we designed has low stiffness in the lateral direction (i.e., the microfluidic device is thin and has a low modulus of elasticity) and high stiffness in the radial direction (i.e., the microfluidic network channels have a small width), which will make it less sensitive to external forces (e.g., blinking, rubbing the eye).

[0015] 4) miLenS allows readout using a smartphone camera and optical adapter. This will provide measurements at discrete points in time. In one variation, a wearable camera that can track the sensor response can also be utilized for continuous and automated measurements.

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

[0017] Microfluidic strain sensor embedded contact lenses are convenient to use and have continuous measurement capabilities. It requires little training to perform measurements and therefore will be used as a device for family medicine. They will enable clinical studies where recording of long-term IOP data on large patient populations is required. Continuous recording of IOP and its analysis will improve the understanding of neurodegenerative diseases and their relationship with pressure. In addition, this will be useful for improving the efficiency and efficacy of drugs used for glaucoma treatment. Thus, miLenS technology provides a promising health care technology for better personalized treatment of glaucoma patients. These advantages listed above will potentially allow patients to use the sensor 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 comprises 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 or electrical components.

[0019] The closed microfluidic network has a volume that is sensitive to axial strain. The closed microfluidic network is differentiated into (i) a gas reservoir, containing a gas, (ii) a liquid reservoir, containing a liquid that changes volume when strain is induced, and (iii) a sensing channel, capable of holding a liquid in the sensing channel. The sensing channel connects the gas reservoir on one end and the liquid reservoir on another end. The sensing channel establishes a liquid-gas equilibrium pressure interface and equilibrium within the sensing channel that will fluidly change in response to radius of curvature variations 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 located inside or outside the at least one ring. In each case, the liquid reservoir volume is highly sensitive to radial forces on the eye wearing the contact lens and to tangential forces on the eye. The liquid reservoir has a high radial stiffness and / or a tangential stiffness and / or a smaller channel width compared to the microfluidic channel wall thickness, resulting in the liquid reservoir being less sensitive to external forces.

[0021] In one embodiment, the liquid reservoir has one or more chambers. The chambers may have concentric rings. The chambers may also have concentric rings that are interconnected at one or more locations. The 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, which 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 about 10 -11 ~10 -8 m 2 It has an inner diameter of 5 to 12 mm with a cross-sectional area of ​​. The present invention provides, for example, the following: (Item 1) 1. A microfluidic strain sensing device for monitoring intraocular pressure changes, comprising: (a) a contact lens; (b) a closed microfluidic network embedded in the contact lens; Equipped with The closed microfluidic network has a volume that is sensitive to an applied strain, the closed microfluidic network further comprising: (i) a gas reservoir containing a gas; (ii) a liquid reservoir, the liquid reservoir containing a liquid that changes volume when the strain is applied; and (iii) a sensing channel capable of retaining the liquid within the sensing channel; Equipped with the sensing channel connects the gas reservoir on one end and the liquid reservoir on another end; the sensing channel establishing a liquid-gas equilibrium pressure interface and equilibrium within the sensing channel that will fluidly change in response to radius of curvature variations 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 the intraocular pressure, a microfluidic strain sensing device. (Item 2) Item 14. The intraocular pressure monitoring device of item 1, wherein the liquid reservoir forms at least one ring and the air reservoir is positioned inside the at least one ring. (Item 3) 2. The intraocular pressure monitoring device of claim 1, wherein the liquid reservoir volume is highly sensitive to tangential forces on the eye compared to radial forces on the eye wearing the contact lens. (Item 4) 2. The intraocular pressure monitoring device of claim 1, wherein the liquid reservoir has a high radial stiffness and / or a smaller channel width compared to the tangential stiffness, resulting in the liquid reservoir being less sensitive to external forces. (Item 5) Item 10. The intraocular pressure monitoring device of item 1, wherein the contact lens is a silicone contact lens, a hydrogel contact lens, or a combination thereof. (Item 6) Item 1 . The intraocular pressure monitoring device of item 1 , wherein the sensing channel has a strain sensitivity of about 4.5 mm of interface movement per about 1% strain applied to the liquid reservoir. (Item 7) Item 2. The intraocular pressure monitoring device of item 1, wherein the sensing channel has an inner diameter of about 1-10 mm. (Item 8) The sensing channel is 10 -11 ~10 -8 m 2 2. The intraocular pressure monitoring device according to item 1, having an inner diameter of 5 to 12 mm with a cross-sectional area of ​​0.1 mm to 0.5 mm. (Item 9) 2. The intraocular pressure monitoring device of claim 1, wherein the liquid reservoir has one or more chambers. (Item 10) 2. The intraocular pressure monitoring device of claim 1, wherein the liquid reservoir has one or more chambers with concentric rings. (Item 11) Item 14. The intraocular pressure monitoring device of item 1, wherein the liquid reservoir has one or more chambers with concentric rings, the concentric rings being connected at one or more locations. (Item 12) 2. The intraocular pressure monitoring device of claim 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) 2. The intraocular pressure monitoring device of claim 1, wherein the surface of the liquid reservoir is patterned. (Item 14) 2. The intraocular pressure monitoring device of claim 1, wherein the surface of the liquid reservoir ceiling has a convex shape, the convex shape curving toward the reservoir channel floor. (Item 15) 2. The intraocular pressure monitoring device of claim 1, wherein the contact lens does not have any actively controlled or electrical components. (Item 16) 2. The intraocular pressure monitoring device of item 1, wherein the closed microfluidic network is transparent. (Item 17) 2. The intraocular pressure monitoring device of claim 1, wherein the closed microfluidic network is oleophobic. [Brief description of the drawings]

[0024] [Figure 1] FIG. 1 illustrates the workflow of a miLens device based on pressure monitoring, according to an exemplary embodiment of the present invention. [Figure 2A] 2A shows an image of a sensor that is only 100 micrometers thick, according to an exemplary embodiment of the invention. The small drops on each side are Norland Optical Adhesive (NOA), which is used to seal the sensor and can be made less than 20 micrometers thick. [Figure 2B] FIG. 2B shows an image of the sensor (300 micrometer final thickness) after embedding the sensor in a contact lens, according to an exemplary embodiment of the present invention. [Diagram 3] FIG. 3 shows a top view of a closed system sensor with multiple ring liquid reservoirs embedded within a contact lens, according to an exemplary embodiment of the present invention. [Figure 4] FIG. 4 shows a side view of a multi-chamber liquid reservoir sensor A) versus a single-chamber liquid reservoir sensor B) and their respective behavior when the sensors as shown in A*) and B*) are stretched under tangential force according to an exemplary embodiment of the present invention. 410-A and 410-B show possible stretching points under sensor stretching. The sensor needs to be made of a soft material which reduces the stiffness in both directions. The sensor needs to be thin. FIG. 4 illustrates this, essentially the microfluidic channel ceiling thickness t1 and floor thickness t2 need to be small (<20 μm). This also reduces the stiffness in both directions. The reservoir ring width w needs to be small (<100 μm). This does not affect the tangential stiffness, but is important in increasing the radial stiffness of the microfluidic channel and improving the sensor performance. [Diagram 5] 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 areas for the three rings show the enlarged rings. [Figure 6] FIG. 6 shows the pressure response of three different sensor types, one, two and five reservoir rings, according to an exemplary embodiment of the 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. For each curve, the average and standard deviation of at least three measurements are used. [Figure 7]7 shows 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 obtained using sensors fabricated at different times with the same parameters, and increases and decreases in sensitivity values ​​are the result of processing differences. Sensitivity is linearly dependent on the number of rings with the 50 and 100 micrometer widths, but is not significantly affected from that for the 200 micrometer width. [Figure 8] 8 shows a side view of the placement of the miLenS on the cornea and the location of the liquid reservoirs according to an exemplary embodiment of the present invention. The insets show a close-up of the liquid reservoirs and the forces acting on them, with inset a) showing a single wide liquid reservoir compressed under a radial force and inset b) showing a series of concentric circles as the liquid reservoirs not compressed under the same force. [Figure 9] 9 shows sensitivity dependence on height for three different ring widths according to an exemplary embodiment of the invention. Multiple data points for some of the heights were obtained using sensors processed at different times with identical parameters, and increases and decreases in sensitivity values ​​are the result of processing differences. Sensitivity is linearly dependent on reservoir height. Red data points 910 show thicker tips (300 micrometers) that show a 50% reduced sensitivity compared to their thinner (150 micrometer) counterparts 920. [Figure 10] 10 shows a close-up view of an auxetic contact lens sensor and liquid reservoir cross-section according to an exemplary embodiment of the invention. Unlike a sensor with a rectangular channel as shown in FIG. 8, this channel has a curved top layer that flattens when a tangential force is applied, as shown by our data and Comsol simulations. [Figure 11] FIG. 11 shows a sensor with a reservoir ceiling patterned with circular and linear convex shapes, according to an exemplary embodiment of the present invention. [Figure 12]12 shows a microscope image of a sensor with a linearly patterned liquid reservoir ceiling according to an exemplary embodiment of the invention on the left, and a comparison of measured sensitivities of 29 micrometers / mmHg and 77 micrometers / mmHg for a flat ceiling sensor versus a curved ceiling (auxetic) device, respectively, on the right. [Figure 13] FIG. 13 shows a method of fabricating a sensor according to an exemplary embodiment of the invention. A refers to UV treatment. B refers to plasma treatment (PDMS). C refers to treatment APTES. 1 refers to glass slide, 2 refers to NOA65 (uncured), 3 refers to PDMS, 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]FIG. 14 shows a method of 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 treatment APTES. 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 top 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 treated with plasma and APTES. 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 liquid 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 from the surface of the mold. [Figure 15] FIG. 15 shows processing steps for the ceiling layer of an auxetic microfluidic sensor according to an exemplary embodiment of the invention. [Figure 16] 16 shows a strain sensor for cancer cell biomechanics according to an exemplary embodiment of the present invention. In the bottom channel, the strain sensor is placed, while cells are seeded in the top channel. [Figure 17] 17 shows a top view and a side view of a contact lens and a 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 Comsol results showing that a 50 micrometer high and 50 micrometer wide channel provides near optimal sensitivity while maintaining a thin device, according to an exemplary embodiment of the invention. The star shape in Figure 18 shows the optimal geometric parameters for maximum volume change (i.e., sensitivity) while keeping the device thin. DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

[0025] Previously reported IOP measurement devices do not consider the directionality of the force 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 on the lens, such as those caused by blinking. An ideal contact lens sensor should be sensitive only to strains applied as a result of corneal radius changes and should not be affected by forces applied perpendicular to the lens (i.e., radial forces). With this in mind, we used COMSOL simulations and experimental measurements to develop a strain sensor that is more sensitive to tangential forces than radial forces on the eye. An embodiment of the present invention is based on microfluidic sensing for IOP measurement and such a desired strain sensor force response.

[0026] FIG. 1 shows an example of the workflow of the IOP self-measurement technique. miLenS is distinct from other sensors because patients would be able to place and remove it themselves, similar to a normal 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 radius of curvature). In this technique, the fluid level in the microfluidic sensing channel of the sensor would change in response to the radius of curvature variation on the cornea. The sensor response would be detected using a smartphone camera equipped with an optical adapter and then converted to a pressure value by a smartphone app. This would eliminate security and health concerns associated with radio frequency or Bluetooth data transfer methods. We have demonstrated an IOP detection limit of 1 mmHg on enucleated porcine eyes, which is sufficient for IOP monitoring applications.

[0027] Similar to electronic circuits, microfluidic circuits can function as low or high pass filters (electrical resistance and capacitance are replaced by the fluid resistance (R) and compliance (C) of compressible materials, 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 circadian variations. Sensors with small RC values ​​will have the ability to detect the effects of eye blinks and ocular pulsation.

[0028] In one exemplary embodiment, a microfluidic strain sensor (FIG. 2A) is integrated into a PDMS contact lens (FIG. 2B) for wearable sensing applications. Referring to the sensor 300 with sensor material 302 in FIG. 3-4, the sensor 300 is embedded in a contact lens 310 and differentiated into a liquid reservoir 320 (which amplifies the displaced liquid volume and is shown in this example as a liquid reservoir ring), a gas reservoir 330, and a sensing channel 340 that is connected to the liquid reservoir 320 on one end and to the gas reservoir 330 on the other end. First, the liquid reservoir 320 is filled with a working liquid, such as oil, using capillary action and then sealed. This creates a stable gas / liquid interface 350 in the sensing channel 340, forming a closed microfluidic network. IOP increases and decreases change the radius of curvature of the cornea, and for every 1 mmHg increase in IOP, the radius of curvature of the cornea increases by 4 μm. This increases the liquid reservoir volume due to the strain applied to the liquid reservoir elastic walls. The increased reservoir volume creates a vacuum and shifts the gas / liquid position 350 in the sensing channel 340 towards the liquid reservoir 320. As the sensing channel cross-sectional area is reduced, the linear liquid displacement required to accommodate the reservoir volume change increases, and therefore sensitivity is improved.

[0029] FIG. 5 shows a top view of two exemplary designs of a microfluidic strain sensor, a single ring 510 versus three rings 520 around the liquid reservoir. Increasing the vertical wall surface area of ​​the liquid reservoir increases the sensitivity of the sensor to changes in IOP. This was tested in two ways: i) increasing the number of walls, ii) increasing the channel wall height. First, we designed and fabricated a sensor with multiple liquid reservoir rings, for example as shown by 520, thus increasing the total wall surface area. Sensitivity results for different numbers of rings are presented in FIGS. 6-7. We found that increasing the number of walls by adding more rings increases the sensitivity of the device in a linear fashion. In contrast, the width of the reservoir had no significant effect on the sensitivity. This phenomenon is a direct result of the interaction between tangential strain and radial force induced collapse as shown in FIG. 8. To test the effect of reservoir wall height, we constructed three types of sensors (50, 100, and 330 μm height) and compared their sensitivity. As shown in FIG. 9, as the reservoir height doubled, the sensitivity also doubled. When we increased the reservoir height to 330 μm, the sensitivity also increased three-fold (shown only for the 200 μm width), proving the effect of vertical wall height. FIG. 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 ∼50% lower sensitivity.

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

[0031] Auxetic metamaterials for microfluidic strain sensing In another version of the sensor, the microfluidic channel network height increases in response to an applied tangential strain 1010. The volume increase is achieved by Poisson's ratio modification through lithographic patterning of the elastomeric sensor. Figure 10 illustrates the working principle of an auxetic metamaterial for strain sensing through a cross section of a contact lens sensor. The ceiling of the microfluidic channel has a convex shape, as shown, i.e., it curves towards the channel interior. This is achieved by patterning the ceiling film with either a circular or linear pattern as shown in Figure 11. These are the only patterns we have tested, but other patterns can be used to achieve the same effect. As shown in Figure 10, when a tangential force is applied (i.e., due to an IOP change), the ceiling is deformed outward due to the convex ceiling, as opposed to the collapse observed when a flat ceiling is used. This deformation towards the front of the sensor causes an increase in the channel height and therefore an amplification of the liquid reservoir volume expansion according to our COMSOL simulations as shown in U.S. Provisional Patent Application No. 62 / 556,366, filed September 9, 2017 (FIG. 14 therein), which is incorporated herein by reference. 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 a factor of 2.5.

[0033] The biocompatible, electronics-free microfluidic mechanical metamaterial enabled the fabrication of highly sensitive and reliable strain sensors. The tangential strain sensing method we developed is specific to IOP, as demonstrated by our experiments. We used the approach to monitor IOP in porcine eyes, demonstrating a detection limit of 1 mmHg (corresponding to 0.05% strain) and reliability over several weeks. The microfluidic strain sensor is capable of measuring ocular strain resulting from shape changes in response to IOP within a clinically relevant range.

[0034] manufacturing We constructed the sensor 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. Then, thin NOA65 films with the required characteristics were fabricated and bonded together to make the sensor, as shown in FIG. 13. For the purpose of this invention, we developed a specific fabrication method to build an extremely thin (about 100 μm) microfluidic device. The gas permeability of the polyurethane used in our device is 6 to 8 orders of magnitude lower than the metals used in wearable electronics.

[0035] We first cut the strain sensor to the desired shape and embedded a flat 100 μm strain sensor (FIG. 2A) into a PDMS contact lens. We built our sensors flat, but they can also be built curved if a curved mold is used. We developed a fabrication protocol that allows us to build contact lenses with 8-15 mm radius of curvature and 10-14 radius as shown in FIG. 2B. We used dome-shaped plastic molds, on which we poured PDMS to obtain 10-100 μm silicone films at the desired radius of curvature, and we bonded our sensor onto the silicone film by (3-aminopropyl)triethoxysilane (APTES) chemical structure. Then, more silicone was poured to completely embed the sensor in the silicone. Details are shown in FIG. 14. Finally, we allowed the silicone to cure overnight at room temperature and then cut out the lenses using a circular puncher. We developed processes and techniques to build sensors as thin as 50 μm thick, so that the entire contact lens sensor can be less than 150 μm.

[0036] For the auxetic sensor version, the only difference in the fabrication is in step 4 of Figure 13, where we 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 physiotherapy monitoring (e.g. in hand joint injuries), voice recognition, fetal / infant monitoring, tremor disorders, robotics, etc.

[0038] 2) Microfluidic strain sensing can be used for biological and biochemical sensing. For example, it can be used to monitor and measure strain applied by cells on a surface. Mechanical cues play an important 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 force than normal cells. Shear stress, one of the representative physical cues, causes increased expression of genes that are activated by mechanical signals. Understanding the mechanical cues generated by cells will be important for understanding the progression of cancer induced by mutations in the cellular mechanical signaling pathways. Our strain sensor will provide direct monitoring of direct cancer cell signaling under exposure to different physical and mechanical cues. This will therefore lead to a novel approach in cancer research. By using our sensor, new biomarkers can be discovered as well as new drug therapies can be implemented. These devices will also be useful in several other conditions, including the regulation of synaptic plasticity in neurons, since force is one of the key factors for the progression of synaptic plasticity.

[0039] To understand the response of cells 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 stiffness. The top channel can also be manipulated by applying different flow rates that change the shear stress. With 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] Cancer tissues exhibit stiffer characteristics as they progress. On average, cancer cells will have four times the stiffness of normal tissues. Understanding the earlier stiffness 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 internal organs in cases of ovarian, liver, and brain cancer. In particular, these patches can be implanted after severe tumor removal surgery to monitor cancer recurrence. Combining microfluidics-based strain sensors with flexible silicon electronics will enable multiple measurements on three-dimensional soft tissues in vivo. The signals can be transferred to a cloud-based system using wi-fi embedding technology. Overall, strain sensors embedded with advanced electronics will provide continuous monitoring of tissues with high chances of cancer recurrence.

[0041] 3) miLenS can be fabricated either by i) embedding strain sensors with the desired shape / size into a contact lens as described, or ii) directly patterning the desired topography on the surface of the contact lens through soft lithography, where features on a mold are transferred to the contact lens.

[0042] 4) The distance between the microscopic geometric features on the contact lens can be measured directly instead of using microfluidics. This distance will vary as a function of IOP. The geometric shape and pattern of these features should be carefully selected to maximize the 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 top and side views of an exemplary geoLenS. The location and shape of the microscopic features to be used for IOP determination are shown. Besides the star shape shown in the top and side views, 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 indicates the angle between features located at the periphery of the contact lens, which determines the number of features that will be angularly located on the contact lens. θ can be 10° (36 features at the periphery) to 180° (2 features at the periphery). A minimum of 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 side view c may be between 0.5 and 1 cm. The characteristic width w of a feature may be between 0.001 and 0.5 cm.

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

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

[0045] The radius of curvature of the eye model varies by approximately 4 μm / mmHg (3 μm / mbar), which closely resembles the behavior of the human eye.

[0046] As shown in US Provisional Patent Application No. 62 / 556366, filed on September 9, 2017 (right side of FIG. 19 therein) (incorporated herein by reference), we made marks on the contact lens, which we placed on the eye model we created. These marks acted as probes, allowing us to measure the change in distance between different locations on the contact lens as a function of the applied pressure. We applied four levels of pressure on the eye model, ranging from 25 mbar to 100 mbar. We sampled four locations on the contact lens (three distance measurements), and the distance between these locations is plotted as a function of the applied pressure, as shown in US Provisional Patent Application No. 62 / 556366, filed on September 9, 2017 (FIG. 29 therein) (incorporated herein by reference). The points located on the center of the contact lens are labeled as location "1" and the number is increased as the points are located further from the center (e.g., location "2"). The distance between the different marked points (e.g., location "1" - location "2") was measured. In FIG. 20, the blue, red, and green lines show the distance as a function of applied pressure for locations 1-2, 2-4, and 4-6, respectively. The corresponding linear fits are plotted as well. Overall, the preliminary findings indicate that the distance between different locations on the geoLenS follows a linear function of applied pressure, and this is within the measurable range.

[0047] 5) geoLenS features can be fabricated similarly to miLenS, or they can simply be marked with ink.

[0048] 6) The miLenS reservoir channel can have a serpentine shape instead of a circular shape.

[0049] 7) The device can be used as a temperature sensor since it is sensitive to the thermal expansion of the material.

[0050] 8) The device is insensitive to changes in air pressure, which allows it to be used in vacuum, e.g. in space applications.

[0051] 9) The image can be taken by a smartphone camera, a special handheld camera, or a wearable camera. The image can be taken directly in front of the eye, at a 45° angle, or 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 can be transferred to a central server for further processing.

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

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

[0057] Additional Technical Notes The present invention relates to a closed microfluidic network for strain sensing applications. The device has a strain sensitivity of 2-15 mm interface movement per 1% strain, depending on the number of rings. The sensitivity can be increased even further by increasing the number of rings. It is robust enough to withstand pressure changes applied over 24 hours and has a life span of several months. These features make it attractive for applications where extreme strain levels smaller than 0.1% need to be measured over time periods longer than 2 hours. The inventors embedded the sensor in a contact lens to monitor intraocular pressure (IOP). The required detection limit for IOP is 1 mmHg. This corresponds to a strain of 0.05%. The inventors achieved this strain detection limit by designing a liquid reservoir network that includes multiple microfluidic channels as liquid reservoirs. The liquid reservoir network is connected to a sensing channel, which is connected to an air reservoir. These three components form a closed system. The sensor with its three components in one possible configuration is shown in Figure 3. FIG. 3 is a top view of the sensor when it is embedded in a contact lens. The sensor is filled with working liquid from the inlet using only capillary forces. When the working liquid reaches the outlet, both the inlet and outlet are sealed using the sensor material to form a closed system with a fixed liquid volume inside. At this point, the liquid fills the sensing channel to about half of 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 works based on the volume amplification of a microfluidic liquid reservoir network when it is stretched under a tangential force. The working principle of the sensor is illustrated in FIG. 4. Here, a different configuration of the sensor components is used for simplicity, where they are linearly distributed instead of radially distributed. Side views of a 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 shape of the sensor and its components expands to A, respectively. * ) and B * ), where 410-A and 410-B are representations of possible stress areas 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 notable mechanical changes that can occur when such a closed microfluidic network is stretched under a tangential force.

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

[0060] ii) Squashing: In the case of a single reservoir, a thin membrane above the liquid reservoir is crushed by B * ) will collapse due to the induced stress and due to the low rigidity of this membrane. When multiple chambers with higher rigidity membranes are used, the collapse will occur at a rate of A * ), would not occur or would be significantly reduced.

[0061] iii) Liquid reservoir volume increase and resulting vacuum effect: When the liquid reservoir width is stretched, its total volume will increase if membrane collapse can be prevented or significantly reduced. This volume increase occurs when the liquid reservoir is B * ), the amplification can be even greater when the auxetic pattern is created on the membrane of a small reservoir chamber. When the volume of the liquid reservoir increases, this causes a vacuum effect, which pulls the liquid / air interface location (3) towards the liquid reservoir. The movement of this interface in μm per IOP change in mmHg is defined as the sensitivity. According to the literature, each 1 mmHg IOP change causes a strain of 0.05%. This strain causes a position change of about 100 μm to the interface location.

[0062] Another factor to be considered 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 is reduced, which will cause slippage and reduced sensitivity, especially for high E sensors. According to our experimental and simulation results, the optimal E is in the range of 0.2-10 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 intraocular pressure changes, comprising: (a) a contact lens; (b) a microfluidic network embedded in the contact lens, the microfluidic network having a volume that is sensitive to an applied strain; The microfluidic network comprises: (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, the sensing channel capable of holding the liquid therein, the sensing channel connecting the gas reservoir on one end and the liquid reservoir on another end; Further equipped with (iv) the sensing channel establishes a liquid-gas equilibrium pressure interface and equilibrium within the sensing channel that fluidly changes in response to radius of curvature variations on the cornea or in response to mechanical stretching and release of the cornea; (v) a microfluidic network, wherein the liquid-gas equilibrium pressure interface and equilibrium are used to measure the intraocular pressure; and (c) an imaging device configured to one or more of: actuate a measurement of intraocular pressure; or interpret the measured intraocular pressure; and A microfluidic strain sensing system comprising:

2. The system of claim 1, wherein the liquid reservoir has one or more chambers with concentric rings.

3. The system described in claim 2, wherein the concentric rings are connected at one or more locations.

4. The system described in claim 2, wherein the sensitivity increases as the number of concentric rings increases.

5. The system described in claim 1, wherein the imaging device comprises a smartphone camera.

6. The system described in claim 1, wherein the imaging device comprises a wearable camera.

7. The system described in claim 1, wherein the imaging device is configured to activate automated measurement of the intraocular pressure.

8. The system described in 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 a healthcare provider.

9. The system of claim 1, wherein the imaging device is configured to transmit the measured intraocular pressure reading to a healthcare provider database.

10. The system described in claim 1, wherein the imaging device is configured to continuously read the measured intraocular pressure.

11. The system of claim 1, wherein the imaging device is configured to read the measured intraocular pressure at discrete times.

12. The system of claim 1, further comprising an optical adapter.

13. The system of claim 1, wherein the imaging device is configured to convert the liquid-gas equilibrium pressure interface and equilibrium into an intraocular pressure value.

14. The system of claim 1, wherein the imaging device is configured to capture images at one or more of: (i) directly in front of the eye, (ii) at a 45 degree angle, (iii) at a 90 degree angle, or (iv) any angle between 0 degrees and 90 degrees.

15. The system of claim 1, wherein the liquid reservoir forms at least one ring and the air reservoir is positioned inside the at least one ring.

16. The system of claim 1, wherein the liquid reservoir volume is more sensitive to tangential forces on the eye compared to radial forces on the eye wearing the contact lens.

17. The system described in claim 1, wherein the contact lens is a silicone contact lens, a hydrogel contact lens, or a combination thereof.

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

19. The system described in claim 1, wherein the contact lens does not have any actively controlled or electrical components.

20. The system described in claim 1, wherein a surface of the liquid reservoir ceiling has a convex shape, the convex shape curved toward the reservoir channel floor.

21. The system described in claim 1, wherein the microfluidic network is closed.

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