JP2022508023A - Closed microfluidic network for strain sensing embedded in contact lenses to monitor intraocular pressure - Google Patents

Abstract

Microfluidic strain sensing device for monitoring intraocular pressure. 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 the strain applied. The network is capable of retaining liquid in (i) a gas reservoir, (ii) a liquid reservoir containing a liquid that changes volume when strain is applied, and (iii) a sensing channel. Distinguished into channels. The sensing channel connects the gas reservoir on one end and the liquid reservoir on the other end. The sensing channel establishes a liquid-gas equilibrium pressure interface and equilibrium within the sensing channel that will change fluidly in response to changes in radius of curvature on the cornea or in response to mechanical extension and release of the cornea. .. The liquid-gas equilibrium pressure interface and equilibrium are used to measure intraocular pressure.

Description

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

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 important for the management of glaucoma.

IOP reduction is the only known method of slowing and / or stopping the progression of glaucoma. For every 1 mmHg IOP reduction, the risk of nerve injury is estimated to be reduced by 11%. Although drug therapy is commonly used to reduce IOP, there are important challenges that need to be addressed to improve the effectiveness of glaucoma treatment. Most importantly, about 50% of patients stop using the drug after 6 months for a variety of reasons. Sustained long-term IOP monitoring with the ability to measure drug efficacy may help patients remain compliant with glaucoma management and assist physicians in that regard. Moreover, in recent years, chronotype of IOP has been established as another risk factor for glaucoma, which further increases the importance of continuous measurement.

Current techniques available for IOP measurement are either non-persistent (Goldmann Application Tonometry), persistent but temporary (Sensimed Triggerfish), or persistent but invasive. Either (embeddable sensor). Autotonocular pressure measuring devices (eg, Icare) can provide long-term data, which are non-invasive but still uncomfortable for patients to the extent that they may require surface anesthetics. Furthermore, it has been found that the results obtained by self-tension measurement are user-dependent.

Approaches for continuous IOP measurement of telemetry have been developed and tested in animal models. Among these approaches, contact lens-based surveillance techniques are attractive because they are non-invasive. One contact lens system (Sensimed AG's Triggerfish) is a contact lens equipped with an electrical strain sensor, antenna, and microchip used to process and transmit signals wirelessly, resulting in subtle changes in corneal curvature. To measure. The technique requires the patient to wear a receiver on his waist for data and power transmission. Due to the thick silicone contact lenses (center thickness 580 μm), this is not as comfortable as routinely used contact lenses, with mild to moderate adverse reactions reported in up to 80% of patients. There is. The requirements of trained personnel and the discomfort and high cost associated with this contact lens platform hinder its use in long-term surveillance applications, but allow testing over a single 24-hour cycle. For this reason, Triggerfish have been found to be more suitable for determining changes in IOP on a daily scale. However, changes in IOP as a response to the drug are on a time scale of several weeks. Similarly, IOP changes in response to certain lifestyle modifications will also be longer than the 24-hour time scale. Therefore, there is a need for continuously worn contact lens sensors that can monitor IOP fluctuations over the long term in order to determine drug efficacy and reduce the number of hospital visits a patient needs to make for regular IOP measurements. do.

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

Recently, Kim et al. However, graphene-Ag-nanowires are used to address the problem of electrode transparency (J. Kim et al., "Wearable smart sensors integrated on soft contact lensesnences for wireless Co., Ltd.". vol. 8, Apr 2017, Art. No. 14997). The first condition for contact lenses with long-term use capacity is high air permeability to prevent hypoxia. Preferably, the conductive component required by the electrical sensor is impermeable to gas. Metals have 8-10 orders of magnitude lower gas permeability compared to soft materials, which is mild in human clinical trials, even when electrically sensitive based contact lenses are used over a single 24-hour period. Causes adverse reactions. Another condition for long-term use is comfort, which is achieved by making high moisture content and thin (<200 micrometer) contact lenses. Electrical sensing methods are sensitive to contact lens hydration levels. Therefore, contact lens electrical sensors are made from silicone with a very low water content instead of standard silicone / hydrogel materials. This reduces the comfort of contact lenses. There are three main reasons for sensitivity to hydration levels. First, the expansion of the hydrogel due to hydration induces strain, which is therefore a source of error in the measurement. Second, the friction between the contact lens and the cornea can be sensitive to the level of hydration and thus affects the sensitivity. Finally, the electrical components are affected by humidity and should therefore be isolated by using a sealant material such as parylene-c.

The present invention provides techniques for measuring IOP that advance the art and eliminate at least some of the current problems or concerns.

J. Kim et al. , "Wearable smart sensors integrated on soft contacts lenses for womens ocular digitals", Nature Communications, vol. 8, April 2017, Art. no. 14997

The present invention relates to a strain sensor that uses the microfluidic principle and is integrated with a contact lens for IOP measurement. The materials used in the present invention are low cost, transparent, breathable and flexible. A method for embedding a microfluidic strain sensor in a silicone contact lens is provided. The microfluidic contact lens sensor (miLenS) allows the patient to measure his or her own IOP and better manage glaucoma.

Microfluidic contact lens sensors measure IOP increases and decreases due to internal factors (ie, metabolism, blinking, and impulsive eye movements) and external factors (ie, drugs, diet, lifestyle, etc.) throughout the patient's life. It is possible to do. The measurements will be made at the patient's discretion (or automatically) and the readings will be achieved by a smartphone camera (or by a wearable camera for automated measurements). This enables home monitoring and continuous data recording. The data is then sent directly to the healthcare provider's database, which allows patients and physicians to monitor IOP fluctuations. Aspects of the present inventors' technology are listed as follows.

1) miLenS uses a hybrid material system in which a narrow microfluidic sensing area (a ring as small as 0.1 mm around the miLenS) is embedded in a silicone or silicone / hydrogel contact lens material. Will be built. Microfluidic sensing channels will be made from clear, soft and oleophobic materials. The sensing material will be 6-10 orders of magnitude more breathable compared to the electronic components.

2) The microfluidic sensing technique has no actively controlled components and works only on the principle of fluid physics. miLenS has no electrical components (no power). This is a low cost device. In addition, it is easier to use by eliminating the cumbersome peripheral components (eg antennas, microchips, etc.) required for wearable electronic sensors for data transmission, reception, and recording. I will provide a.

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 due to hydration of the contact lens material. The sensors designed by the present inventors have low lateral rigidity (ie, the microfluidic device is thin and has a low modulus) and high rigidity in the radial direction (ie, the microfluidic network). The channel will have a small width), which will make it less sensitive to external forces (eg, blinks, eye scratches).

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

5) Sustained data recorded using existing techniques show that IOP increases or decreases by about 5-15 mmHg daily and hourly and by about 15-40 mmHg per second. The microfluidic network circuit designed by the present inventors has the ability to filter out large increases and decreases that occur on a short time scale due to blood pressure or muscle contraction. In this case, the sensor actually acts as a fluid lowpass filter, which responds only to changes that occur in minutes or later. In a similar fashion, fluid components can be designed to register only rapid IOP changes. Sensors capable of measuring events occurring at different time scales can make better estimates of true IOP based on measurements of the radius of curvature of the cornea.

Microfluidic strain sensor embedded contact lenses are convenient to use and have continuous measurement capability. It requires little training to make measurements and will therefore be used as a device for family medicine. These will enable clinical studies that require the recording of long-term IOP data for large patient populations. Sustained recording of IOP and its analysis will improve understanding of neurodegenerative diseases and their relationship to pressure. In addition, it will be useful for improving the efficiency and efficacy of drugs used for the treatment of glaucoma. Therefore, miLenS technology provides 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.

In one embodiment, the invention provides a microfluidic strain sensing device for monitoring changes in intraocular pressure. Closed microfluidic networks are transparent and / or oleophobic. The microfluidic strain sensing device has a contact lens and a closed microfluidic network embedded in the contact lens. Contact lenses are silicone contact lenses, hydrogel contact lenses, or a combination thereof. Contact lenses do not have any actively controlled or electrical components.

Closed microfluidic networks have a volume that is sensitive to axial strain. Closed microfluidic networks include (i) a gas reservoir, a liquid reservoir containing a liquid that changes volume when strain is induced, and (iii) a liquid in a sensing channel. Distinguished into sensing channels that can be retained. The sensing channel connects the gas reservoir on one end and the liquid reservoir on the other end. The sensing channel establishes a liquid-gas equilibrium pressure interface and equilibrium within the sensing channel that will change fluidly in response to changes in radius of curvature on the cornea or in response to mechanical extension and release of the cornea. do. The liquid-gas equilibrium pressure interface and equilibrium are used to measure intraocular pressure.

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 very sensitive to the radial force to the eye wearing the contact lens and to the tangential force to the eye. The liquid reservoir has high stiffness in the radial direction and / or a channel width that is smaller than the stiffness in the tangential direction and / or the microfluidic channel wall thickness, making the liquid reservoir less sensitive to external forces. Bring.

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

In one embodiment, the surface of the liquid reservoir can be patterned. The surface of the liquid reservoir ceiling can have a convex shape, which can be curved towards the reservoir channel floor.

The sensing channel has a strain sensitivity of about 4.5 mm interfacial movement per about 1% strain applied to the liquid reservoir. In one embodiment, the sensing channel has an inner diameter of about 1-10 mm. In another embodiment, the sensing channel has an inner diameter of 5-12 mm with a cross-sectional area of ^10-11 to ^10-8 m ² .

FIG. 1 shows a workflow of a miLens device based on pressure monitoring according to an exemplary embodiment of the invention. FIG. 2A shows an image of a sensor that is only 100 micrometers thick, according to an exemplary embodiment of the invention. A small drop on each side is a Norland optical adhesive (NOA) that is used to seal the sensor and can be made to a thickness of less than 20 micrometers. FIG. 2B shows an image of a sensor (final thickness of 300 micrometers) after the sensor is embedded in a contact lens according to an exemplary embodiment of the invention. FIG. 3 shows a top view of a closed system sensor with a plurality of ring liquid reservoirs embedded in a contact lens according to an exemplary embodiment of the invention. FIG. 4 shows the sensors as shown in the multi-chamber liquid reservoir sensor A) vs. single-chamber liquid reservoir sensor B) and A ^* ) and B ^* ) according to an exemplary embodiment of the invention extended under tangential force. A side view of their individual behavior when done is shown. 410-A and 410-B indicate the possible extension points of the sensor under extension. The sensor needs to be made from a soft material that reduces stiffness in both directions. The sensor needs to be thin. FIG. 4 illustrates this, and basically the microfluidic channel ceiling thickness t1 and floor thickness t2 need to be small (<20 μm). This also reduces stiffness in both directions. The reservoir 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 sensor performance. FIG. 5 shows a top view of a single ring liquid reservoir vs. 3 ring reservoir according to an exemplary embodiment of the invention. The circled area for the three rings indicates the enlarged ring. 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 mean and standard deviation of at least three measurements are used. FIG. 7 shows the sensitivity dependence on the number of reservoir rings for three different ring widths according to an exemplary embodiment of the invention. Multiple data points for some of the number of rings were obtained using sensors machined at different time points with the same parameters, and the increase or decrease in sensitivity values is the result of machined differences. Sensitivity depends linearly on the number of rings with 50 and 100 micrometer widths, but is not significantly affected by those for 200 micrometer widths. FIG. 8 shows a side view of the installation of miLenS on the cornea and the location of the liquid reservoir according to an exemplary embodiment of the invention. The inset shows an enlarged view of the liquid reservoirs and the forces acting on them, inset a) shows a single wide liquid reservoir that is compressed under radial force, and inset b) shows the same force. Shown below is a series of concentric circles as an uncompressed liquid reservoir. FIG. 9 shows the 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 machined at different time points with the same parameters, and the increase or decrease in sensitivity values is the result of machined differences. Sensitivity depends linearly on reservoir height. Red data points 910 indicate thicker chips (300 micrometers), which show a 50% reduction in sensitivity compared to the thinner (150 micrometer) counterpart 920. FIG. 10 shows an enlarged view of a cross section of an auxetic contact lens sensor and a liquid reservoir according to an exemplary embodiment of the present invention. Unlike a sensor with a rectangular channel as shown in FIG. 8, this channel has a curved top layer. The upper layer is flattened when a tangential force is applied, as shown by our data and the Comsol simulation. FIG. 11 shows a sensor with a patterned reservoir ceiling with circular and linear convex shapes according to an exemplary embodiment of the invention. FIG. 12 shows a microscopic image of a sensor with a linearly patterned liquid reservoir ceiling according to an exemplary embodiment of the invention on the left. On the right side, a comparison of measured sensitivities of 29 micrometer / mmHg and 77 micrometer / mmHg for flat ceiling sensors vs. curved ceiling (auxetic) devices is shown, respectively. FIG. 13 shows a method of processing a sensor according to an exemplary embodiment of the present invention. A refers to UV treatment. B refers to plasma treatment (PDMS). C refers to the processing APTES. 1 refers to a glass slide, 2 refers to NOA65 (uncured), 3 refers to PDMS, and 4 refers to NOA65 (cured). In step 1, NOA65 is sandwiched between two PDMS coated glass slides and UV cured to create a 20 micrometer film. This is repeated twice. In step 2, NOA65 is dropped onto the 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. 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 the processing APTES. D refers to a curing (heat) treatment. 5 refers to a hemispherical mold for processing contact lenses, 6 refers to a sensor, and 7 refers to the upper layer of a contact lens. In step 6, PDMS is poured onto the contact lens mold. It is then cured at 80 degrees Celsius and plasma and APTES treated. 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 and sealed with working liquid. In step 8, more PDMS is poured onto the sensor and cured at room temperature. In step 9, the contact lens is stripped from the surface of the mold. FIG. 15 shows the machining steps of the ceiling 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 invention. In the bottom channel, strain sensors are installed while cells are seeded in the top channel. FIG. 17 shows a top view and a side view of the location of the contact lens and shape according to an exemplary embodiment of the invention. In addition to the star shapes in the top and side views, other exemplary shapes can also be provided. A combination of these shapes can also be used. FIG. 18 shows the COMSOL results according to an exemplary embodiment of the invention, in which channels 50 micrometer high and 50 micrometer wide provide near optimum sensitivity while maintaining thin devices. The star shape in FIG. 18 shows the optimum geometric parameters for maximum volume change (ie, sensitivity) while keeping the device thin.

The IOP measuring devices reported so far do not consider the direction of the force acting on the sensor. For example, Chen et al. (G.-Z. Chen, I.-S. Chan, L. K. K. Leng, and DC Lam, "Soft wearable contact lens sensor for continuus engineer essence" The capacitance measurement-based sensor according to vol. 36, no. 9, pp. 1134-1139, Sep 2014) responds to the radial force applied to the lens, such as due to the blink of an eye. An ideal contact lens sensor should only be sensitive to the strain applied as a result of a change in the radius of the cornea and is affected by the force applied perpendicular to the lens (ie, the radial force). You shouldn't take it. With this in mind, we used COMSOL simulations and experimental measurements to develop strain sensors that are more sensitive to tangential forces than radial forces to the eye. Embodiments of the invention are based on microfluidic sensing for IOP measurements and such desired strain sensor force responses.

FIG. 1 shows an example of a workflow of the IOP self-measurement technique. The miLenS is distinctly different from other sensors because the patient will be able to install and remove it on their own, similar to regular contact lenses. As the 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 sensor's microfluidic sensing channel will change in response to changes in the radius of curvature on the cornea. The sensor response will be detected using a smartphone camera equipped with an optical adapter and then converted to a pressure value by the smartphone app. This will eliminate security and health concerns associated with radio frequencies or Bluetooth® data transfer methods. We have demonstrated an IOP detection limit of 1 mmHg on the eye of enucleated pigs, which is sufficient for IOP monitoring applications.

Similar to electronic circuits, microfluidic circuits can act as low or high pass filters (electrical resistance and capacitance are replaced by the fluid resistance (R) and compliance (C) of the compressible material, respectively. ). The RC value will determine the time constant of the sensor response. Sensors with large RC values will not respond to rapid changes, but will be sensitive to slowly fluctuating chronotypes. Sensors with small RC values will have the ability to detect the effects of blinks and eye pulsations.

In one exemplary embodiment, the microfluidic strain sensor (FIG. 2A) is integrated into a PDMS contact lens (FIG. 2B) for wearable sensing applications. Referring to the sensor 300 with the sensor material 302 of FIG. 3-4, the sensor 300 is embedded in the contact lens 310 and as a liquid reservoir 320 (amplifying the displaced liquid volume and in this embodiment as a liquid reservoir ring). Shown), a gas reservoir 330, and a sensing channel 340 connected to the liquid reservoir 320 on one end and connected 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 within the sensing channel 340 and forms a closed microfluidic network. Increasing or decreasing the IOP changes 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 wall. The increased reservoir volume creates a vacuum and shifts the gas / liquid position 350 within the sensing channel 340 towards the liquid reservoir 320. As the sensing channel cross-sectional area decreases, the linear liquid displacement required to adapt to changes in reservoir volume increases, and thus sensitivity increases.

FIG. 5 shows two exemplary designs of a microfluidic strain sensor, namely a top view of a single ring 510 vs. three rings 520 for a 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 height of the channel walls. First, we designed and machined 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 Figure 6-7. We have found that increasing the number of walls by adding more rings increases the sensitivity of the device in a linear fashion. In contrast, reservoir width had no significant effect on sensitivity. This phenomenon is a direct result of the interaction between tangential strain and radial force-induced crushing as shown in FIG. To test the effect of reservoir wall height, we constructed three types of sensors (50, 100, and 330 μm height) and compared their sensitivities. As shown in FIG. 9, as the reservoir height doubled, so did the sensitivity. When we increased the reservoir height to 330 μm, the sensitivity also increased 3-fold (shown only for 200 μm width), demonstrating 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 (indicated by 100T and 330T), the thicker sensor has a lower sensitivity of ~ 50%.

In summary, we have experimentally scrutinized a wide parameter range to understand and optimize sensor performance. We have various numbers of reservoir rings (1-5), ring width (w = 50-500 μm), reservoir height (50, 100, 330 μm), tip thickness (130 μm, 300 μm), and Sensors with different Young's moduli of about 1 MPa (PDMS) vs. about 10 MPa (NOA 65) and about 100 MPa (NOA 61) were machined. The results of these sensitivity tests demonstrated the following: i) Increased liquid reservoir height increases sensitivity. ii) We can improve sensitivity by adding more reservoir rings to the design as needed (eg, depending on the required continuous wear contact lens properties). iii) Rigidity (Young's modulus (E) x chip thickness (t) / width (w)) does not significantly alter sensitivity, however, it does, such as comfort and lens / corneal mechanical interactions. It needs to be optimized in consideration of other factors.

In another version of the auxetic metamaterial sensor for microfluidic strain sensing, the microfluidic channel network height increases in response to the tangential strain 1010 applied. Volume increase is achieved by Poisson's ratio correction through lithography patterning of the elastomer sensor. FIG. 10 shows the working principle of an autetic metamaterial for strain sensing through a cross section of a contact lens sensor. The ceiling of the microfluidic channel has a convex shape, i.e., curves towards the inside of the channel, as shown. This is achieved by patterning the ceiling film with either a circular or linear pattern as shown in FIG. These are the only patterns we have tested, but other patterns can be used to achieve the same effect. As shown in FIG. 10, when a tangential force is applied (ie, due to the IOP change), the ceiling is convex as opposed to the crushing observed when a flat ceiling is used. Due to the shaped ceiling, it is deformed outward. This variant towards the front of the sensor is shown in US Provisional Patent Application No. 62/5563666 filed September 9, 2017 (FIG. 14 thereof) (incorporated herein by reference). According to our COMSOL simulations such as, it causes an increase in channel height and thus an amplification of liquid reservoir volume expansion. This amplification increases the sensitivity of the sensor.

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

Biocompatible, electronic-free microfluidic mechanical metamaterials have made it possible to process highly sensitive and reliable strain sensors. The tangential strain sensing method developed by the present inventors is unique to IOP, as demonstrated by the experiments of the present inventors. We used this approach to monitor IOP in the porcine eye, demonstrating a detection limit of 1 mmHg (corresponding to 0.05% strain) and reliability over several weeks. Microfluidic strain sensors can measure eye strain due to shape changes in response to IOP within clinically relevant ranges.

Manufacturing We constructed sensors using photolithography and soft lithography techniques. First, a polydimethylsiloxane (PDMS) soft mold was processed. As the sensor material, a polyurethane-based Norland optical adhesive 65 (NOA65) was selected due to its transparency, flexibility, oleophobicity, and biocompatibility. A thin NOA65 film with the required characteristics was then made and joined together to make the sensor, as shown in FIG. For the purposes of the present invention, the inventors have developed a specific processing method for constructing an extremely thin (about 100 μm) microfluidic device. The gas permeability of polyurethane used in our devices is 6-8 orders of magnitude lower than that of metals used in wearable electronics.

We first cut the strain sensor into a desired shape and embedded a flat 100 μm strain sensor (FIG. 2A) in a PDMS contact lens. We have constructed our sensors flat, but they can also be constructed curved if a curved mold is used. We have developed a processing protocol that allows us to construct contact lenses with an 8-15 mm radius of curvature and a 10-14 radius as shown in FIG. 2B. We used dome-shaped plastic molds, where we poured PDMS over them to obtain a 10-100 μm silicone film with a desired radius of curvature, and we , (3-Aminopropyl) Triethoxysilane (APTES) chemical structure was used to bond our sensors onto a silicone film. Then more silicone was poured to completely embed the sensor in the silicone. Details are shown in FIG. Finally, we cured the silicone overnight at room temperature and then cut out the lens using a circular puncher. We have developed processes and techniques for constructing sensors as thin as 50 μm so that the entire contact lens sensor can be less than 150 μm.

With respect to the auxetic sensor version, the only difference in manufacture was in step 4 of FIG. 13, where we used a patterned film instead of a flat film as the bottom layer. The patterning was done as shown in FIG.

Modifications and Modifications 1) Microfluidic strain sensing principles 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 (eg, in wrist joint injury), speech recognition, fetal / infant monitoring, tremor disease, robotics, and the like.

2) Microfluidic strain sensing can be used for biological and biochemical sensing. For example, it can be used to monitor to measure the strain applied by cells on the surface. Mechanical cues play important roles in cellular processes such as cell differentiation, apoptosis, and motility. The cells sense and confer forces on 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 activated by mechanical signals. Understanding the mechanical cues generated by cells will be important for understanding cancer progression induced by mutations in cell mechanical signaling pathways. Our strain sensors will provide direct monitoring of direct cancer cell signaling under exposure to different physical and mechanical cues. Therefore, this will bring a new approach in cancer research. By using our sensors, new biomarkers can be discovered and new drug therapies can be implemented. These devices will also be useful in several other conditions, including the regulation of neuronal synaptic plasticity, as force is one of the key factors in the progression of synaptic plasticity.

To understand the cell's response to different conditions, two layers of microfluidic channels can be constructed as shown in FIG. As the cell grows, it can image the strain sensor on the bottom channel. This will provide tissue rigidity. The upper channel can also be manipulated by applying different fluxions that change the shear stress. In this design, the mechanical responses of cells can be observed while they are mechanically manipulated. This design will be used in biomarkers and drug development.

Cancer tissues exhibit more rigid properties as they progress. On average, cancer cells will have four times the stiffness of normal tissue. Understanding the earlier stiffness of cancer cells will lead to earlier cancer detection. The strain sensor can be incorporated into a patch that can be used externally on the skin. Specifically, it 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 the case of ovarian, liver and brain cancers. In particular, these patches can be implanted after severe tumor removal surgery to monitor the recurrence of the cancer. Combining a microfluidic-based strain sensor with a flexible silicon device will enable multiple measurements of 3D soft tissue in vivo. This signal can be transferred to a cloud-based system using wi-fi burial technology. Overall, strain sensors incorporating advanced electronics will provide continuous monitoring of tissues that are likely to have cancer recurrence.

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

4) Distances between fine geometric features on contact lenses 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 of a contact lens sensor (geoLenS) with geometric features similar to miLenS. FIG. 17 shows a top view and a side view of an exemplary geoLenS. The location and shape of the subtle features to be used for IOP determination is shown. In addition to the star shapes shown in the top and side views, other exemplary shapes are also provided. A combination 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 0.5-1 cm. θ indicates the angle between features located around the contact lens, which determines the number of features that will be angularly placed on the contact lens. θ can be from 10 ° (36 features in the periphery) to 180 ° (2 features in the periphery). A minimum of two features are needed on the contact lens. d ₁ , d ₂ , d ₃ , ... .. .. d _n represents the distance between successive features and can be 0.01-1 cm. 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 the side view can be 0.5 to 1 cm. The characteristic width w of the feature can be 0.001 to 0.5 cm.

As the IOP changes, the distance between the peripheral features, eg d ₁ , changes and can be used as a measure of the IOP change. Distances between central features, such as d ₂ or d ₃ , or the width w of any feature, can be used as reference measurements because they do not change in response to the IOP. The distance between opposing features in the perimeter (total distance is 2d) varies most in response to IOP changes. The distance of any one of the geoLenS features to a known feature of the eye (ie, the border of the iris) can be detected as a measure of IOP.

To test the feasibility of the concepts proposed above, we have machined contact lenses made from PDMS and having a thickness of about 250 μm. For testing purposes, the inventors are incorporated herein by reference to US Provisional Patent Application No. 62/556366 filed September 9, 2017 (in the left side of FIG. 19). ), A realistic eye model made from PDMS was machined.

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

As set forth in US Provisional Patent Application No. 62/556366 filed September 9, 2017 (on the right side of FIG. 19) (incorporated herein by reference), the inventors of the present invention. Marked on the contact lens and we placed it on the eye model we created. These marks acted as probes, allowing us to measure changes in distance between different locations on contact lenses as a function of applied pressure. We applied four levels of pressure in the ocular model, varying from 25 mbar to 100 mbar. We sampled four locations (three distance measurements) on contact lenses and the distance between these locations was filed September 9, 2017, US Provisional Patent Application No. 62 /. It is plotted as a function of the applied pressure, as shown in 556366 (in which FIG. 29) (incorporated herein by reference). Points located on the center of the contact lens are labeled as location "1" and the number is increased as the points are located farther from the center (eg, location "2"). The distance between the different marked points (eg, location "1" -location "2") was measured. In FIG. 20, the blue, red, and green lines indicate distances as a function of applied pressure for locations 1-2, location 2-4, and location 4-6, respectively. The corresponding linear fit is plotted in the same way. Overall, preliminary findings indicate that the distance between different locations on geoLenS follows a linear function of the applied pressure, which is within the measurable range.

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

6) The miLenS reservoir channel can have a meandering shape instead of a circular shape.

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

8) This device has low sensitivity to changes in atmospheric pressure. It can be used in vacuum, for example in space applications.

9) Images can be taken by a smartphone camera, by a special handheld camera, or by a wearable camera. Images can be taken from directly opposite the eye at a 45 ° angle, or at a 90 ° angle, or at any angle between 0 ° and 90 °.

10) A front or rear camera of the smartphone can be used for imaging.

11) Images can be collected by the patient at will, or automatically when the patient is reading something on the phone.

12) Image analysis can be performed by the microprocessor of the camera or transferred to a major server for further processing.

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

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

Additional Technical Notes The present invention relates to closed microfluidic networks for strain sensing applications. The device has a strain sensitivity of 2-15 mm interfacial movement per 1% strain, depending on the number of rings. Sensitivity can still be further increased by increasing the number of rings. It is robust enough to withstand pressure changes applied over a 24-hour period and has a lifespan of several months. These features make this attractive for applications where extreme strain levels less than 0.1% need to be measured over time periods longer than 2 hours. We have embedded a 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%. We have achieved this strain detection limit by designing a liquid reservoir network that includes multiple 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 a possible configuration is shown in FIG. FIG. 3 is a top view of the sensor showing when it is embedded in a contact lens. The sensor is filled with working liquid from the inlet using only capillary force. When the working liquid reaches the outlet, both the inlet and outlet are sealed using 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 its total length, creating a liquid / air interface. Both contact lenses and sensors are made from elastomers such as silicone and polyurethane, but can also be made from other materials such as silicone / hydrogel.

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

i) Elongation: When A ^* ) and B ^* ) are compared to A) and B), respectively, the total sensor length (1-1') may increase due to elongation. I understand. Similarly, the liquid reservoir network width (2-2') will also increase.

ii) Crushing: In the case of a single reservoir, the thin membrane above the liquid reservoir is due to the induced stress and due to the low rigidity of the membrane, as shown in B ^* ). Will crush. When multiple chambers with higher rigid membranes are used, crushing will not occur or will be significantly reduced, as shown in A ^* ).

iii) Liquid reservoir volume increase and consequent vacuum effect: As the liquid reservoir width is extended, its total volume will increase if membrane crushing can be prevented or significantly reduced. This volume increase can be amplified if the liquid reservoir consists of multiple chambers with a small width as shown in B ^* ). Amplification will be even greater when the auxetic pattern is created on the membrane of the small reservoir chamber. As the volume of the liquid reservoir increases, this causes a vacuum effect, which pulls the liquid / air interface position (3) towards the liquid reservoir. The movement of this interface in μm per μm change in mmHg is defined as 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 with respect to the interface position.

Another factor to consider for maximum sensitivity is 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 is especially for high E sensors, gliding and Will cause reduced sensitivity. According to our experiments and simulation results, the optimum 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 (17)

A microfluidic strain sensing device for monitoring changes in intraocular pressure.
(A) Contact lenses and
(B) Provided with a closed microfluidic network embedded in the contact lens.
The closed microfluidic network has a volume that is sensitive to the strain applied, and the closed microfluidic network further comprises.
(I) A gas reservoir containing gas and
(Ii) A liquid reservoir, wherein the liquid reservoir contains a liquid that changes its volume when the strain is applied.
(Iii) The sensing channel is provided with the sensing channel capable of holding the liquid in the sensing channel.
The sensing channel connects the gas reservoir on one end and the liquid reservoir on the other end.
The sensing channel will change fluidly in response to a variation in radius of curvature on the cornea or in response to mechanical extension and release of the cornea. Liquid-gas equilibrium pressure interface and equilibrium within the sensing channel. Established,
The liquid-gas equilibrium pressure interface and equilibrium are microfluidic strain sensing devices used to measure the intraocular pressure. The intraocular pressure monitoring device according to claim 1, wherein the liquid reservoir forms at least one ring, and the air reservoir is located inside the at least one ring. The intraocular pressure monitoring device according to claim 1, wherein the liquid reservoir volume is very sensitive to a tangential force to the eye as compared to a radial force to the eye wearing the contact lens. The liquid reservoir has a high stiffness in the radial direction and / or a channel width smaller than the stiffness in the tangential direction, resulting in the liquid reservoir being less sensitive to external forces, claim 1. The described intraocular pressure monitoring device. The intraocular pressure monitoring device according to claim 1, wherein the contact lens is a silicone contact lens, a hydrogel contact lens, or a combination thereof. The intraocular pressure monitoring device of claim 1, wherein the sensing channel has a strain sensitivity of about 4.5 mm interfacial movement per about 1% strain applied to the liquid reservoir. The intraocular pressure monitoring device according to claim 1, wherein the sensing channel has an inner diameter of about 1 to 10 mm. The intraocular pressure monitoring device according to claim 1, wherein the sensing channel has an inner diameter of 5 to 12 mm with a cross-sectional area of ^10-11 to ^10-8 m ² . The intraocular pressure monitoring device according to claim 1, wherein the liquid reservoir has one or more chambers. The intraocular pressure monitoring device of claim 1, wherein the liquid reservoir has one or more chambers with concentric rings. The intraocular pressure monitoring device of claim 1, wherein the liquid reservoir has one or more chambers with concentric rings, the concentric rings being connected at one or more locations. The intraocular pressure monitoring device of claim 1, wherein the liquid reservoir has one or more chambers with concentric rings, the sensitivity of which increases as the number of concentric rings increases. The intraocular pressure monitoring device according to claim 1, wherein the surface of the liquid reservoir is patterned. The intraocular pressure monitoring device according to claim 1, wherein the surface of the liquid reservoir ceiling has a convex shape, and the convex shape curves toward the reservoir channel floor. The intraocular pressure monitoring device of claim 1, wherein the contact lens has no actively controlled or electrical components. The intraocular pressure monitoring device according to claim 1, wherein the closed microfluidic network is transparent. The intraocular pressure monitoring device according to claim 1, wherein the closed microfluidic network is oleophobic.

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