CN117694824A

CN117694824A - Closed microfluidic network for monitoring intraocular pressure

Info

Publication number: CN117694824A
Application number: CN202311700820.7A
Authority: CN
Inventors: 伊斯梅尔·埃姆雷·阿拉西; 塞夫达·阿高歌鲁; 穆拉特·巴达伊; 普里西拉·迪普
Original assignee: Smart Contact Lens Co
Current assignee: Smart Contact Lens Co
Priority date: 2018-09-20
Filing date: 2018-09-20
Publication date: 2024-03-15
Also published as: JP2022508023A; CN113164042A; JP2023060306A; EP3852607A4; CA3112502A1; EP3852607A1; JP7244631B2; WO2020060558A1

Abstract

A microfluidic strain sensing device for monitoring intraocular pressure. The device has a contact lens and a closed microfluidic network that fits into the contact lens. The network has a volume that is sensitive to the applied strain. The network is characterized in that: (i) a gas reservoir containing a gas, (ii) a liquid reservoir containing a liquid, the liquid reservoir changing volume upon application of strain, and (iii) a sensing channel capable of retaining a liquid within the sensing channel. The sensing channel is connected at one end to a gas reservoir and at the other end to a liquid reservoir. The sensing channel establishes a liquid-gas equilibrium pressure interface and equilibrium within the sensing channel that will change fluidly as a response to a change in the radius of curvature of the cornea, or as a response to mechanical stretching and relaxation of the cornea. The liquid-gas equilibrium pressure interface and equilibrium are used to measure intraocular pressure.

Description

Closed microfluidic network for monitoring intraocular pressure

The present application is a divisional application of the chinese patent application entitled "closed microfluidic network for strain sensing in contact lenses for monitoring intraocular pressure" with application date 2018, 09, 20, 201880099640.9, application number corresponding to PCT application, 2018, 09, 20, application number (PCT/US 2018/052062).

Technical Field

The present invention relates to devices, systems and methods for monitoring intraocular pressure. In particular, the present invention relates to a microfluidic network design of strain sensors that works to monitor intraocular pressure based on mechanical amplification of the volume of the microfluidic channel.

Background

Glaucoma is a neurodegenerative disease that causes irreversible damage to the optic nerve of the eye and thus loss of vision. Continuous and long-term monitoring of intraocular pressure (IOP) is critical for the treatment of glaucoma.

IOP lowering is the only known way to reduce and/or terminate glaucoma progression. It is estimated that per 1mmHg drop in IOP, the risk of nerve damage is reduced by 11%. Drug therapy is commonly used to lower IOP, but there are important challenges to be addressed to improve the effectiveness of glaucoma treatment. Most importantly, nearly 50% of patients discontinue use of the drug after six months for various reasons. Continuous long-term IOP monitoring with the ability to measure efficacy can help patients maintain compliance and help doctors treat glaucoma. In addition, in recent years, diurnal changes in IOP have been identified as another risk factor for glaucoma, thereby even further enhancing the importance of continuous measurement.

The techniques currently available for IOP measurement are discontinuous (Goldmann Applanation Tonometry), either continuous but transient (Sensimed Triggerfish) or continuous but invasive (implantable sensor). Self-test ocular pressure devices (e.g., icare) can provide long-term data and are non-invasive, but remain uncomfortable for the patient to the extent they may require local anesthesia. In addition, it was found that the results obtained by the self-tonometry method were user-dependent.

Methods of continuous IOP telemetry have been developed and tested in animal models. Among these methods, contact lens (contact lens) based monitoring techniques are of interest because they are non-invasive. A contact lens system (Triggerfish) measures small changes in corneal curvature through a contact lens equipped with an electronic strain sensor, antenna and microchip (which is used to wirelessly process and transmit signals). This technique requires the patient to wear the receiver on the wrist for data transfer and power transfer. Because the silicone contact lens is thick (center thickness 580 μm), it is less comfortable than a daily use contact lens; mild to moderate adverse reactions were reported in up to 80% of patients. The need for trained personnel, and the inappropriately high cost associated with such contact lens platforms, preclude their use for long-term monitoring applications, but only allow testing during a 24-hour single cycle. For this reason Triggerfish was found to be more suitable for determining daily changes in IOP. However, IOP changes as a response to drugs are on a time scale of weeks. Similarly, IOP changes in response to certain lifestyle adjustments will also be on a time scale longer than 24 hours. Accordingly, there is a need for a continuous wear contact lens sensor that can monitor IOP changes over a long period of time in order to determine efficacy in order to reduce the number of times a patient needs to seek medical attention for routine IOP measurements.

Other examples of contact lens sensors are based on measuring resistance, inductance, and capacitance changes in response to pressure-induced strain. In these examples, the sensor response is typically detected remotely by measuring the resonant frequency change with an external reader coil or through a bluetooth connection. Electrical measurements require conductive components within the lens that are generally not transparent and not breathable.

Recently, kim et al used graphene-Ag-nanowires to solve the electrode transparency problem (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 of contact lenses having long-term use properties is high gas permeability to prevent hypoxia. Disadvantageously, the electrically conductive components required for electronic sensors are gas impermeable. Compared to soft materials, metals have a gas permeability of 8-10 orders of magnitude lower, and this causes slight adverse reactions in human trials when using contact lenses based on electronic sensing even during only 24 hours. Another condition of long-term use is comfort, which is achieved by preparing contact lenses that are high in water content and thin (< 200 microns). Electronic sensing methods are sensitive to the hydration level of the contact lens. Thus, the electronic sensor of the contact lens is made of silicone (which has a very low water content) instead of standard silicone/hydrogel materials. This reduces the comfort of the contact lens. There are three main reasons for sensitivity to hydration levels. First, hydrogels swell due to hydration to induce strain, and therefore, are a source of measurement error. Second, friction between the contact lens and the cornea may be sensitive to the level of hydration, thus affecting sensitivity. Finally, the electronic components are affected by humidity and therefore should be isolated by using a sealant material such as parylene-c.

The present invention advances the art and provides techniques for measuring IOP to eliminate at least some of the current difficulties and problems.

Disclosure of Invention

The present invention relates to strain sensors integrated with contact lenses that utilize microfluidic principles for IOP measurement. The materials used in the present invention are low cost, transparent, breathable and flexible. A method of embedding a microfluidic strain sensor in a silicone contact lens is provided. Microfluidic contact lens sensors (mlens) allow patients to measure their own IOP to better treat glaucoma.

Microfluidic contact lens sensors are capable of measuring IOP fluctuations due to internal factors (i.e., metabolism, blink motion, and saccadic motion) and external factors (i.e., drugs, diet, lifestyle, etc.) during the period of a patient's lifetime. The measurements will be made at will (or automatically) by the patient, where the reading will be done by a smartphone camera (or by a wearable camera for automated measurements). This allows home monitoring and continuous data recording. The data will then be sent directly to the database of the medical service provider, thereby allowing the patient and doctor to monitor IOP changes. Aspects of our technique are listed below:

1) The mlens will be constructed with a hybrid material system in which a narrow microfluidic sensing region (a ring with a width as low as 0.1mm at the edges of the mlens) is embedded in a silicone or silicone/hydrogel contact lens material. The microfluidic sensing channel will be made of a transparent soft oleophobic material. The sensing material will be 6-10 orders of magnitude more breathable than the electronic component.

2) Microfluidic sensing technology has no actively controlled components and works only on the basis of fluid physics principles. The miLenS does not contain all electronic components (no electricity). It is a low cost device. Furthermore, this provides easier usability by eliminating the tedious peripheral components (e.g., antennas, microchips, etc.) required in wearable electronic sensors for data transmission, reception and recording.

) The sensor will be strain sensitive and respond to changes in the radius of curvature of the cornea, but have low sensitivity to forces applied directly by the eyelid or due to hydration of the contact lens material. The sensor we design has a low stiffness in the lateral direction (i.e. the microfluidic device is thin and has a low elastic modulus) and a high stiffness in the radial direction (i.e. the microfluidic network channels have a small width), thereby making it insensitive to external forces (e.g. blinking, eye rubbing).

4) The miLenS can realize reading by using a smart phone camera and an optical adapter. This will provide measurements at discrete points in time. In one variation, a wearable camera that can track sensor response can also be used for continuous and automated measurements.

5) Continuous data recorded with the prior art indicate IOP fluctuations of about 5-15mmHg per day and hour, and 15-40mmHg per second. Microfluidic network circuitry we have designed has the ability to filter out large fluctuations that occur in 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 in minutes or slower. In a similar manner, the fluidic component may be designed to only show rapid changes in IOP. Sensors that can measure events occurring in different time scales can better estimate the true IOP based on the radius of curvature measurements of the cornea.

Contact lenses embedded with microfluidic strain sensors are easy to use and have continuous measurement capabilities. It requires minimal training to make a measurement and will therefore be used as a device for home medical treatment. These will enable clinical studies in which long-term IOP data needs to be recorded for a large patient population. Continuous recording of IOP and analysis thereof will improve our understanding of neurodegenerative diseases and their correlation with stress. In addition, it will help improve the efficacy and efficacy of the drugs used in the treatment of glaucoma. Thus, the miens technology provides a promising healthcare technology for better personalized care of glaucoma patients. These advantages listed above would potentially enable the patient to use the sensor for a long period of time without the assistance of trained personnel.

In one embodiment, the present invention provides a microfluidic strain sensing device for monitoring changes in intraocular pressure. The closed microfluidic network is transparent and/or oleophobic. The microfluidic strain sensing device has a contact lens and a closed microfluidic network embedded with the contact lens. The contact lens is a silicone contact lens, a hydrogel contact lens, or a combination thereof. Contact lenses have no actively controlled components or electronic components.

The closed microfluidic network has a volume that is sensitive to axial strain. The closed microfluidic network is characterized by: (i) a gas reservoir containing a gas, (ii) a liquid reservoir containing a liquid, the liquid reservoir changing volume upon induction of the strain, and (iii) a sensing channel capable of retaining the liquid within the sensing channel. The sensing channel is connected at one end to the gas reservoir and at the other end to the liquid reservoir. The sensing channel establishes a liquid-gas equilibrium pressure interface and balance within the sensing channel that will change fluidly as a response to a change in the radius of curvature of the cornea, or as a response to mechanical stretching and relaxation of the cornea. The liquid-gas equilibrium pressure interface and equilibrium are used to measure intraocular pressure.

The liquid reservoir forms at least one ring and wherein the air reservoir is located inside or outside said at least one ring. In each case, the volume of the liquid reservoir is highly sensitive to tangential forces on the eye relative to radial forces on the eye on which the contact lens is worn. The liquid reservoir has a high stiffness and/or a small channel width in the radial direction relative to the stiffness and/or microfluidic channel wall thickness in the tangential direction, resulting in the liquid reservoir becoming insensitive to external forces.

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, wherein the sensitivity increases with an increasing number of concentric rings.

In one embodiment, the surface of the liquid reservoir may be patterned. The surface of the top plate of the liquid reservoir may have a convex shape and the convex shape may curve towards the channel floor of the reservoir.

The sensing channel has a strain sensitivity of about 4.5mm 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 5-12mm, 10 ^-11 –10 ^-8 m ² Is a cross-sectional area of (c).

Drawings

Fig. 1 shows a workflow of a mlens device based on pressure monitoring according to an exemplary embodiment of the present invention.

Fig. 2A shows an image of a sensor that is only 100 microns thick, according to an exemplary embodiment of the present invention. The droplets on each side are Norland Optical Adhesive (NOA) used to seal the sensor and can be made less than 20 microns thick.

Fig. 2B shows an image (300 micron final thickness) of a sensor after embedding the sensor in a contact lens, according to an exemplary embodiment of the present invention.

Fig. 3 shows a top view of a closed system sensor with a multi-ring liquid reservoir embedded in a contact lens, according to an exemplary embodiment of the present invention.

Fig. 4 shows a side view of a multi-chamber liquid reservoir sensor a) as compared to Shan Qiangshi liquid reservoir sensor B), and their blood behavior when the sensors are stretched under tangential forces as shown in a) and B), according to an exemplary embodiment of the invention. 410-a and 410-B show possible stretch points under the stretch of the sensor. The sensor must be made of a soft material to reduce stiffness in both directions. The sensor must be thin. Fig. 4 shows this: basically, the microfluidic channel top plate thickness t1 and bottom plate thickness t2 must be small (< 20 μm). This also reduces the stiffness in both directions. The ring width w of the reservoir must be small (< 100 μm). This does not affect the tangential stiffness but increases the radial stiffness of the microfluidic channel and is critical in improving sensor performance.

Fig. 5 shows a top view of a single ring liquid reservoir in comparison to a tricyclic reservoir, according to an exemplary embodiment of the invention. The circled area of the three rings shows the ring enlarged.

FIG. 6 shows pressure responses for three different sensor types according to an exemplary embodiment of the present invention; 1. 2 and 5 reservoir rings. The ring height, width and spacing are 100 microns. 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 3 measurements were used.

Fig. 7 shows the dependence of sensitivity on the number of reservoir rings for three different ring widths according to an exemplary embodiment of the present invention. For some ring numbers, multiple data points are obtained with sensors manufactured at different times under the same parameters; fluctuations in the sensitivity values are a result of manufacturing variations. The sensitivity is linearly dependent on the number of loops with widths of 50 and 100 microns, but is not significantly affected by this for a width of 200 microns.

Fig. 8 shows a side view of the position of the miens disposed on the cornea and the fluid reservoir, according to an exemplary embodiment of the present invention. The inset shows a close-up view of the liquid reservoirs and the forces acting on them; panel a) shows a single wide liquid reservoir compressed under radial force and panel b) shows a series of concentric circles as a liquid reservoir uncompressed under the same force.

Fig. 9 shows the dependence of sensitivity on height for three different loop widths according to an exemplary embodiment of the present invention. For some heights, multiple data points are obtained with sensors manufactured at different times under the same parameters; fluctuations in the sensitivity values are a result of manufacturing variations. The sensitivity is linearly dependent on the reservoir height. Red data points 910 indicate thicker chips (300 microns) and they exhibit a 50% reduced sensitivity compared to thinner (150 microns) counterparts 920.

Fig. 10 shows a close-up view of an auxetic contact lens sensor and a liquid reservoir cross-section, according to an example embodiment of the present invention. Unlike the sensor with a rectangular channel as shown in fig. 8, this channel has a curved top layer. This top layer flattens out when tangential force is applied, as shown from our data and Comsol simulation.

Fig. 11 shows a sensor according to an exemplary embodiment of the present invention, wherein the reservoir ceiling has a circular and linear convex pattern.

Fig. 12 shows a microscopic image of a sensor with a linear pattern of liquid reservoir top plate on the left side, according to an exemplary embodiment of the present invention. The measurement sensitivity comparison of the flat roof sensor compared to the curved roof (auxetic) device is shown on the right side, 29 and 77 microns/mmHg, respectively.

Fig. 13 shows a method of manufacturing a sensor according to an exemplary embodiment of the present invention. A represents UV treatment. B represents plasma treatment (PDMS). C represents treatment APTES.1 denotes slide, 2 denotes NOA65 (uncured), 3 denotes PDMS,4 denotes NOA65 (cured). Step 1 is sandwiching NOA65 between two PDMS coated slides and UV cured to produce a 20 micron film. This was repeated twice. Step 2 is to drop NOA65 onto a mold and plasma treat the 20 micron film from step 2. Step 3 is to sandwich the two layers from step 2 together and UV cure. Step 4 is plasma treating the 70 micron layer from step 3. The 20 micron layer from step 1 was subjected to plasma treatment and APTES treatment. Step 5 is to clamp the two layers from step 4 together.

Fig. 14 shows a method of embedding a sensor in a contact lens according to an exemplary embodiment of the present invention. B represents plasma treatment (PDMS). C represents treatment APTES. D represents a curing (heat) treatment. 5 denotes a hemispherical mold for contact lens manufacture, 6 denotes a sensor, and 7 denotes the top layer of a contact lens. Step 6 is pouring PDMS onto the contact lens mold. Curing, plasma and APTES treatments were then performed at 80 degrees celsius. The bottom surface of the sensor is subjected to plasma treatment. Step 7 is to place the sensor bottom surface on the PDMS coated contact lens mold. The sensor reservoir is filled with a working liquid and sealed. Step 8 is to pour more PDMS onto the sensor and cure it at room temperature. Step 9 is to peel the contact lens from the surface of the mold.

Fig. 15 shows a step in the fabrication of a top plate layer of an auxetic microfluidic sensor according to an exemplary embodiment of the present invention.

Fig. 16 shows a strain sensor for biomechanics of cancer cells according to an exemplary embodiment of the present invention. In the bottom channel, a strain sensor is placed while cells are seeded in the top channel.

Fig. 17 shows top and side views of a contact lens, and the location of the shape, according to an exemplary embodiment of the present invention. Other example shapes may be provided in addition to the star shape in top and side views. Combinations of these shapes may also be used.

Fig. 18 shows a Comsol result in which 50 micron high and 50 micron wide channels provide near optimal sensitivity while maintaining a thin device, according to an exemplary embodiment of the invention. The star shape in fig. 18 shows the best geometry for maximum volume change (i.e. sensitivity) while keeping the device thin.

Detailed Description

The IOP measuring devices reported so far do not consider the directional nature of the forces acting on the sensor. For example, capacitive measurement-based sensors of Chen et al (g.—z.chen, i.— S.Chan, L.K.K.Leung, and d.c. lam, "Soft wearable contact lens sensor for continuous intraocular pressure monitoring," Medical Engineering & Physics, vol.36, no.9, pp.1134-1139, sep 2014) respond to radial forces applied to the lens, for example, due to blinking. An ideal contact lens sensor should be sensitive only to strain imposed by changes in the radius of the cornea, but should not be affected by forces applied perpendicularly to the lens (i.e., radial forces). In view of this, we have employed COMSOL simulation and experimental measurements to develop strain sensors that are more sensitive to tangential forces than radial forces acting on the eye. Embodiments of the present invention are based on microfluidic sensing for IOP measurement and such desired strain sensor force response.

Fig. 1 shows an example of a workflow of IOP self-measurement technology. The mlens differs from other sensors in that the patient will be able to place and remove it by himself similar to a conventional contact lens. As IOP fluctuates, the radius of curvature of the cornea changes (every 1mmHg change in IOP results in a 4 μm change in radius of curvature). In this technique, the fluid level in the microfluidic sensing channel of the sensor will change in response to a change in the radius of curvature of the cornea. The sensor response will be detected with a smartphone camera equipped with an optical adapter and then converted to a pressure value by the smartphone application. It will eliminate the security and health problems associated with radio frequency or bluetooth data transmission methods. We have demonstrated an IOP detection limit of 1mmHg on enucleated porcine eyes that is sufficient for IOP monitoring applications.

Microfluidic circuits, like electronic circuits, can act as low-pass or high-pass filters (the resistance and capacitance being 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. A sensor with a large RC value will not respond to rapid changes but will be sensitive to slowly changing diurnal changes. A sensor with a small RC value will have the ability to detect the effects of blinks and eye beats.

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 fig. 3-4, the sensor 300 has a sensor material 302, the sensor 300 is embedded in a contact lens 310, featuring a liquid reservoir 320 (expanding the displaced liquid volume and shown as a liquid reservoir ring in this embodiment), a gas reservoir 330, and a sensing channel 340 connected at one end to the liquid reservoir 320 and at the other end to the gas reservoir 330. First, the liquid reservoir 320 is filled with a working liquid such as oil by capillary action, and then sealed. This creates a stable gas/liquid interface 350 in the sensing channel 340 and forms a closed microfluidic network. IOP fluctuations change the radius of curvature of the cornea; the radius of curvature of the cornea is at most 4 μm for every 1mmHg increase in IOP. 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 in the sensing channel 340 and displaces the gas/liquid location 350 toward the liquid reservoir 320. As the sensing channel cross-sectional area decreases, the linear liquid displacement required to accommodate the reservoir volume change increases, and thus the sensitivity improves.

Fig. 5 shows a top view of two example designs (single ring 510 versus tri-ring 520) for a microfluidic strain sensor for a liquid reservoir. Increasing the vertical wall surface area of the liquid reservoir increases the sensitivity of the sensor to IOP changes. Testing was performed in two ways: i) Increasing the number of walls; ii) increasing the height of the channel walls. First, we designed and manufactured a sensor with multiple liquid reservoir rings (as shown by e.g. 520), thereby increasing the overall wall surface area. The sensitivity results are shown in fig. 6 to 7 for different numbers of loops. We have found that increasing the number of walls by adding more rings increases the sensitivity of the device in a linear manner. Conversely, the width of the reservoir has no significant effect on sensitivity. This phenomenon is a direct result of the interaction between collapse-induced tangential strain and radial force, as shown in fig. 8. To test the effect of this reservoir wall height, we constructed three types of sensors (50, 100 and 330 μm height) and compared their sensitivity. As shown in fig. 9, when the reservoir is highly doubled, the sensitivity is doubled. When we increase the reservoir height to 330 μm, the sensitivity also increases to 3 times (only 200 μm width is shown), demonstrating the effect of vertical wall height. Fig. 9 further illustrates the effect of sensor stiffness. When comparing a 150 μm thick sensor with a 300 μm thick sensor (shown by 100T and 330T), the thicker sensor has a sensitivity of 50% lower.

In summary, we have scanned a large parameter range through experiments to understand and optimize sensor performance. We have fabricated sensors with different numbers of reservoir rings (1-5), ring widths (w=50-500 μm), reservoir heights (50, 100, 330 μm) and chip thicknesses (130 μm, 300 μm) and different young's modulus-l MPa (PDMS) vs-10 MPa (NOA 65) and-100 MPa (NOA 61). The results of these sensitivity tests indicate that: i) The 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 desired sustained wear contact lens properties); iii) Stiffness (young's modulus (E) x chip thickness (t)/width (w)) does not significantly change sensitivity, but it needs to be optimized for other factors such as comfort and lens/corneal mechanical interaction. Auxetic metamaterials for microfluidic strain sensing

In another variation of the sensor, the microfluidic channel network height increases in response to the applied tangential strain 1010. The volume increase is achieved by poisson's ratio modification of the lithographic patterning of the elastomeric sensor. Fig. 10 shows the principle of operation of auxetic metamaterials (auxetic metamaterial) for strain sensing via a cross section of a contact lens sensor. The top plate of the microfluidic channel has a convex shape, i.e. is curved towards the inside of the channel, as shown. This is achieved by patterning the top sheet film in a circular or linear pattern as shown in fig. 11. While these are just patterns we tested, other patterns may be used to achieve the same effect. When tangential force is applied (i.e., due to IOP variations), the top plate is deformed outward due to the convex top plate, as shown in fig. 10, as opposed to the collapse observed when a flat top plate is used. This deformation toward the sensor front causes the channel height to increase, and therefore, according to our COMSOL simulation, as shown in U.S. provisional patent application 62/556366 filed on 9/2017 (fig. 14 herein), which is incorporated by reference, the liquid reservoir volume expansion is exaggerated. This amplification improves the sensitivity of the sensor.

Fig. 12 shows on the left side an image of the liquid reservoir on the auxetic sensor with a linear pattern of convex structures on the top plate. Fig. 12 shows on the right an experimental sensitivity comparison between a flat and a curved (auxetic) device. The sensitivity increase was 2.5 times.

The biocompatible and electronics-free micro-fluidic mechanical metamaterial enables the fabrication of highly sensitive and reliable strain sensors. The tangential strain sensing method we developed is specific for IOP as demonstrated by our experiments. We have used this method to monitor IOP in pig eyes and present an l-mmHg detection limit (corresponding to 0.05% strain) and reliability over several cycles. Microfluidic strain sensors can measure ocular strain in a clinically relevant range due to shape changes in response to IOP.

Production of

We use photolithography and soft etching techniques to build the sensor. First, a Polydimethylsiloxane (PDMS) flexible mold was fabricated. Polyurethane-based Norland optical adhesive 65 (NOA 65) was chosen as the sensor material for transparency, flexibility, oleophobicity and biocompatibility. The NOA65 films with the desired characteristics are then fabricated and bonded together to fabricate the sensor, as shown in fig. 13. For the purposes of the present invention, we have developed specific fabrication methods to build extremely thin (-100 μm) microfluidic devices. The polyurethane used in our device has a gas permeability of 6-8 orders of magnitude lower than the metals used in wearable electronics.

We have first cut the strain sensor into the desired shape and embed a flat 100 μm strain sensor (fig. 2A) in the PDMS contact lens. While we have constructed our sensors as flat, they can also be constructed curved if curved molds are used. We have developed a manufacturing scheme in which we can construct contact lenses with a radius of curvature of 8-15mm and a radius of 10-14mm, as shown in fig. 2B. We have used a dome-shaped plastic mold where we pour PDMS onto them to obtain a 10-100 μm silicone film of the desired radius of curvature, and we bond our sensors to the silicone film by (3-aminopropyl) triethoxysilane (APTES) chemistry. Then, more silicone was poured to fully embed the sensor in the silicone. Details are shown in fig. 14. Finally, we cut out the lenses with a circular punch after allowing the silicone to cure overnight at room temperature. We have developed methods and techniques to construct sensors as low as 50 μm thick so that the overall contact lens sensor can be less than 150 μm.

For the auxetic sensor variant, the only difference in production is step 4 of fig. 13, where we have used a patterned film instead of a smooth film as the bottom layer. Patterning is performed as shown in fig. 15.

Variations and modifications

1) Microfluidic strain sensing principles can be used in a wide range of medical applications requiring strain sensing. Biomedical applications other than glaucoma treatment may be listed: physical therapy monitoring (e.g., hand joint injury), voice recognition, fetal/infant monitoring, tremor disease, robotics, and the like.

2) Microfluidic strain sensing may be used for biosensing and biochemical sensing. For example, it may be used for monitoring to measure the strain exerted by cells on a surface. Mechanical cues play an important role in cellular processes such as cell differentiation, apoptosis and motility. The cells sense and exert forces on the substrate on which they grow. Tumor cells produce more force than normal cells. Shear stress, one of the major physical cues, is leading to upregulation of genes activated by mechanical signals. Understanding the mechanical cues generated by cells will be critical to understanding the progression of cancer triggered by mutations in the mechanical force transduction pathway of the cells. Our strain sensor will provide direct monitoring of direct cancer cell signaling exposed to different physical and mechanical cues. Thus, it will bring a novel approach in cancer research. By using our sensor, new biomarkers will be discovered and new drug therapies can be administered. These devices will also contribute to several other conditions, including regulating the synaptic plasticity of neurons, as force is one of the key factors in the progression of synaptic plasticity.

To understand the response of cells to different conditions. The two-layer microfluidic channel can be constructed as shown in fig. 16. As the cells grow, we can image the strain sensor on the bottom channel. This will provide tissue hardening. The top channel may also be manipulated by applying different flows that change the shear stress. In this design, cellular mechanical responses can be observed as they are being mechanically manipulated. This design will be used for biomarker and drug development.

Cancer tissues exhibit more sclerotic characteristics as they progress. Cancer cells will harden on average 4 times as much as normal tissue. It is understood that earlier sclerosis of cancer cells will lead to earlier cancer detection. The strain sensor may be incorporated in a patch that may be applied externally to the skin. In particular, it can be used in skin cancer and breast cancer types. In the case of ovarian, liver and brain cancers, such patches with infrared beads embedded in microchannels can be optimized and implanted in internal organs. These patches may be implanted to monitor cancer recurrence, especially after severe tumor removal surgery. The combination of a microfluidic based strain sensor with flexible silicon electronics will enable multiplexed measurements of three-dimensional soft tissue in vivo. This signal may be transmitted to the cloud-based system using wi-fi embedded technology. In general, strain sensors incorporating advanced electronics will provide continuous monitoring of tissue with a high likelihood of cancer recurrence.

3) The miLenS may be produced by: i) Embedding a strain sensor having a desired shape/size in a contact lens, as described; or ii) patterning the desired topography directly onto the surface of the contact lens by soft etching, wherein features on the mold are transferred to the contact lens.

4) The distance between the microscopic geometric features on the contact lens can be measured directly instead of using microfluidics. This distance will change as a function of IOP. The geometry and pattern of these features should be carefully selected to maximize sensitivity to IOP. IOP will be measured based on imaging a contact lens sensor (geoLenS) with geometrical features similar to miLenS. FIG. 17 shows a top view of an example geoLenSA drawing and a side view. The location and shape of microfeatures for IOP determination are shown. Other example shapes are provided in addition to the star shape shown in top and side views. Combinations of these shapes may also be used. In top view, the radius of the contact lens is denoted by r and the value of r may be between 0.5 and 1 cm. θ shows the angle between features located at the edge of the contact lens and it determines the number of features that will be angularly disposed on the contact lens. θ may be between 10 ° (36 features at the edge) and 180 ° (two features at the edge). A minimum of two features are required on a contact lens. d, d ₁ 、d ₂ 、d ₃ 、...d _n Represents the distance between successive features and may be 0.01 to 1cm. Total distance d=d ₁ +d ₂ +d ₃ +...+d _n Should be less than r. The radius of curvature rc of the contact lens shown in side view may be 0.5 to 1cm. The characteristic width w of the feature may be 0.001 to 0.5cm.

As IOP changes, the distance between edge features, e.g., d ₁ Changes and can be used as a measure of IOP changes. Distance between central features, e.g. d ₂ Or d ₃ Or the width w of any feature may be used as a reference measure because they do not change in response to IOP. The distance between the opposing features at the edge (total distance is 2 d) changes maximally in response to IOP changes. The distance between any of the geoLenS features to the known feature of the eye (i.e., iris boundary) can be detected as a measure of IOP.

To test the feasibility of the concept presented above, we manufactured contact lenses made of PDMS and having a thickness of 250um. For testing, we produced a realistic eye model made of PDMS, as shown in U.S. provisional patent application 62/556366 filed on 9/2017 (where fig. 19-left), which is incorporated herein by reference.

The radius of curvature of the eye model varies by 4 μm/mmHg (3 μm/mbar) and this is very close to the behavior of the human eye.

We place the mark on the contact lens and we place it on the eye model we manufactured as shown in U.S. provisional patent application 62/556366 filed on 9/2017 (where fig. 19-right), which is incorporated herein by reference. These markers act as probes and enable us to measure the change in distance between different locations on the contact lens as a function of the applied pressure. We varied the applied pressure from 25mbar to 100mbar for 4 levels in the eye model. We sampled 4 locations (3 distance measurements) on the contact lens and plotted the distance between these locations as a function of applied pressure, as shown in U.S. provisional patent application 62/556366 filed on 9, 2017 (fig. 29 herein), which is incorporated by reference. The point located on the center of the contact lens is marked as position '1' and the number increases as the point location is farther from the center (e.g., position '2'). The distance between the differently marked points (e.g., positions '1' to '2') is measured. In fig. 20, blue, red and green lines show the distances of positions 1 to 2, positions 2 to 4 and positions 4 to 6, respectively, as a function of the applied pressure. The corresponding linear fit is also plotted. In summary, the preliminary results show that the distance between the different positions on the geoLenS corresponds to a linear function of the applied pressure, and this is in a measurable range.

5) The geoLenS features may be manufactured similarly to miLenS or they may be only marked with ink.

6) The mlens reservoir channel may have a serpentine shape instead of a circular shape.

7) The device may be used as a temperature sensor because it is sensitive to thermal expansion of the material.

8) The device is insensitive to air pressure variations. It may be used for vacuum applications such as space applications.

9) The image may be taken by a smartphone camera, a special hand-held camera, or a wearable camera. The image may be taken directly across the eye, at an angle of 45 deg. or at an angle of 90 deg. or at any angle between 0 deg. -90 deg..

10 Front and rear cameras of the smartphone can be used for imaging.

11 The images may be collected by the patient at will or automatically while the patient is reading on the cell phone.

12 Image analysis may be performed by the microprocessor of the camera or may be transmitted to a host server for further processing.

13 A patient may pay for subscription to cloud services such as data storage, analysis, etc.

14 The mlens channel may be filled with a colored liquid to improve contrast on the iris or sclera.

Additional technical features

The present invention relates to closed microfluidic networks for strain sensing applications. The device has a strain sensitivity of 2-15mm interface movement/1% strain, depending on the number of rings. The sensitivity can be even further improved by increasing the number of rings. It is robust enough to withstand the pressure changes applied during 24 hours and has a service life of months. These features make it attractive for applications where extreme strain levels of less than 0.1% need to be measured over a period of longer than 2 hours. We have embedded sensors in contact lenses to monitor intraocular pressure (IOP). The required IOP detection limit is 1mmHg. This corresponds to a strain of 0.05%. We have achieved this strain detection limit by designing a liquid reservoir network comprising a plurality of microfluidic channels as a liquid reservoir. The liquid reservoir network is connected to the sensing channel and the sensing channel is connected to the air reservoir. These three components form a closed system. A sensor (with its three components in one possible configuration) is shown in fig. 3. Fig. 3 is a top view of the sensor, showing it when it is embedded in a contact lens. The sensor is filled with working liquid from the inlet by using only capillary forces. When the working liquid reaches the outlet, both the inlet and outlet are sealed by the use of a sensor material to form a closed system with a fixed liquid volume inside. At this point, the liquid fills the sensing channel, about half its total length, creating a liquid/air interface. Both the contact lens and the sensor are made of elastomers such as silicone and polyurethane, but may also be made of other materials such as silicone/hydrogel.

The sensor operates based on the expansion of the volume of the microfluidic liquid reservoir network when it is stretched by tangential forces. The principle of operation of this sensor is depicted in fig. 4. For simplicity, another configuration of the sensor elements is used here, in which they are distributed linearly rather than radially. Comparing side views of a sensor with a liquid reservoir, the liquid reservoir may have multiple chambers a) as opposed to a single wide chamber B). When the sensor is stretched by tangential forces, the shape of the sensor and its components changes as shown in a) and B), respectively. 410-a and 410-B are illustrations of possible stress areas on the sensor near the liquid reservoir. For reference, the total initial length of the sensor is shown as l-1', the total initial width of the liquid reservoir network is shown as 2-2', and the initial position of the liquid air interface is shown as 3. There are three significant mechanical variations that can occur when such closed microfluidic networks are stretched by tangential forces;

i) Elongation: when a) and B) are compared with a) and B), respectively, it is seen that the total sensor length (l-1') will increase due to elongation. Similarly, the liquid reservoir network width (2-2') will also increase.

ii) collapse: in the case of a single reservoir, the membrane above the liquid reservoir will collapse due to induced stress and due to the low stiffness of the membrane, as shown in B). When using multiple chambers with higher stiffness films, collapse will not occur, or will be significantly reduced, as shown in a).

iii) Liquid reservoir volume increase and resulting vacuum effect:

when the liquid reservoir width is prolonged, the total volume of the membrane will increase if it can be prevented or significantly reduced. If the liquid reservoir consists of a plurality of chambers with small widths, this volume increase can be enlarged, as shown in B). This expansion will be even greater when building up an auxetic pattern on the membrane of the small reservoir chamber. When the volume of the liquid reservoir increases, this results in a vacuum effect and this vacuum pulls the liquid/air interface position (3) towards the liquid reservoir. The movement (μm)/IOP change (mmHg) of this interface is defined as sensitivity. According to literature, IOP changes per 1mmHg resulted in a strain of 0.05%. This strain causes a positional change of about 100 μm to the interface position.

The present invention provides embodiments including, but not limited to, the following:

1. a microfluidic strain sensing device for monitoring intraocular pressure changes, comprising:

(a) A contact lens;

(b) A closed microfluidic network mated with the contact lens, wherein the closed microfluidic network has a volume that is sensitive to an applied strain, and wherein the closed microfluidic network further comprises:

(i) A gas reservoir for containing a gas,

(ii) A liquid reservoir containing a liquid, said liquid reservoir changing volume upon application of said strain, and

(iii) A sensing channel capable of holding the liquid within the sensing channel,

wherein the sensing channel is connected at one end to the gas reservoir and at the other end to the liquid reservoir,

wherein the sensing channel establishes a liquid-gas equilibrium pressure interface and balance within the sensing channel that will change fluidly as a response to a change in radius of curvature of the cornea, or as a response to mechanical stretching and relaxation of the cornea, and

wherein the liquid-gas equilibrium pressure interface and equilibrium are used to measure the intraocular pressure.

2. The intraocular pressure monitoring device of embodiment 1 wherein the liquid reservoir forms at least one ring and wherein the air reservoir is located inside the at least one ring.

3. The intraocular pressure monitoring device of embodiment 1 wherein the volume of the liquid reservoir is highly sensitive to tangential forces on the eye relative to radial forces on the eye on which the contact lens is worn.

4. The intraocular pressure monitoring device of embodiment 1 wherein the liquid reservoir has a high stiffness and/or a small channel width in radial direction relative to stiffness in tangential direction, resulting in the liquid reservoir becoming insensitive to external forces.

5. The intraocular pressure monitoring device of embodiment 1 wherein the contact lens is a silicone contact lens, a hydrogel contact lens, or a combination thereof.

6. The intraocular pressure monitoring device of embodiment 1 wherein the sensing channel has a strain sensitivity of about 4.5mm interface movement per about 1% strain applied to the fluid reservoir.

7. The intraocular pressure monitoring device of embodiment 1 wherein the sensing channel has an inner diameter of about 1-10 mm.

8. The intraocular pressure monitoring device of embodiment 1 wherein the sensing channel has an inner diameter of 5-12mm, 10 ^-11 –10 ^-8 m ² Is a cross-sectional area of (c).

9. The intraocular pressure monitoring device of embodiment 1 wherein the fluid reservoir has one or more chambers.

10. The intraocular pressure monitoring device of embodiment 1 wherein the liquid reservoir has one or more chambers with concentric rings.

11. The intraocular pressure monitoring device of embodiment 1 wherein the liquid reservoir has one or more chambers with concentric rings, wherein the concentric rings are connected at one or more locations.

12. The intraocular pressure monitoring device of embodiment 1 wherein the fluid reservoir has one or more chambers with concentric rings, wherein sensitivity increases with increasing number of concentric rings.

13. The intraocular pressure monitoring device of embodiment 1 wherein a surface of the liquid reservoir is patterned.

14. The intraocular pressure monitoring device of embodiment 1 wherein a surface of the top plate of the fluid reservoir has a convex shape, wherein the convex shape curves toward the channel floor of the reservoir.

15. The intraocular pressure monitoring device of embodiment 1 wherein the contact lens is devoid of actively controlled components or electrical components.

16. The intraocular pressure monitoring device of embodiment 1 wherein the closed microfluidic network is transparent.

17. The intraocular pressure monitoring device of embodiment 1 wherein the closed microfluidic network is oleophobic.

Another factor that should be considered for maximum sensitivity is the young's modulus (E) of the sensor material. Increasing this E reduces comfort. When contact lenses with high lubricity are used for improved comfort, contact friction between the cornea and the sensor/lens will be reduced, thereby resulting in slipping and reduced sensitivity, especially for high E sensors. According to our experimental and simulation results, the optimal E is in the range of 0.2-10MPa for maximum sensitivity and comfort. When E decreases below 2MPa, the width of the reservoir channel must also decrease below 100 μm.

Claims

(a) A contact lens;

(i) A gas reservoir for containing a gas,

2. The intraocular pressure monitoring device of claim 1, wherein the liquid reservoir forms at least one ring and wherein the air reservoir is located inside the at least one ring.

3. The intraocular pressure monitoring device of claim 1, wherein a volume of the liquid reservoir is highly sensitive to tangential forces on the eye relative to radial forces on the eye on which the contact lens is worn.

4. The intraocular pressure monitoring device of claim 1, wherein the liquid reservoir has a high stiffness and/or a small channel width in radial direction relative to stiffness in tangential direction, resulting in the liquid reservoir becoming insensitive to external forces.

5. The intraocular pressure monitoring device of claim 1, wherein the contact lens is a silicone contact lens, a hydrogel contact lens, or a combination thereof.

6. The intraocular pressure monitoring device of claim 1, wherein the sensing channel has a strain sensitivity of about 4.5mm interface movement per about 1% strain applied to the liquid reservoir.

7. The intraocular pressure monitoring device of claim 1, wherein the sensing channel has an inner diameter of about 1-10 mm.

8. The intraocular pressure monitoring device of claim 1, wherein the sensing channel has an inner diameter of 5-12mm, 10 ^-11 –10 ^-8 m ² Is a cross-sectional area of (c).

9. The intraocular pressure monitoring device of claim 1, wherein the liquid reservoir has one or more chambers.

10. The intraocular pressure monitoring device of claim 1, wherein the liquid reservoir has one or more chambers with concentric rings.