Enhanced sensor for continuous biological monitor

Abstract

In brief, a sensor for a continuous biological monitor is provided for measuring a level of a target analyte of a patient. The sensor has a working lead and a reference lead, wherein the working lead has an analyte confinement layer that delivers one of more than 1000 analyte molecules from the patient to an enzyme layer. The enzyme layer has an enzyme trapped in a polyurethane cross-linked with an acrylic polyol. When free electrons are generated, a conductor transfers the electrons to the biological monitor. In some cases, the sensor may not be constructed using any expensive platinum.

Description

[0001] This application is a divisional application of the invention patent application with application number 201980024867.1, application date April 5, 2019, and invention name “An enhanced sensor for continuous biological monitor”.

[0002] Related Applications

[0003] This application claims priority to the following provisional applications: (1) U.S. Provisional Application No. 62 / 653,821, filed on April 6, 2018, and entitled “Continuous Glucose Monitoring Device”; (2) U.S. Provisional Application No. 62 / 796,832, filed on January 25, 2019, and entitled “Carbon Working Electrode for a Continuous Biological Sensor”; and (3) U.S. Provisional Application No. 62 / 796,842, filed on January 25, 2019, and entitled “Enhanced Membrane Layers for the Working Electrode of a Continuous Biological Sensor”; each of the provisional applications is incorporated herein by reference in its entirety. Background Art

[0004] Monitoring of glucose levels is critical for diabetics. A continuous glucose monitoring (CGM) sensor is a type of device that measures glucose from fluid sampled from just one area under the skin multiple times a day. A CGM device typically includes a small housing in which the electronics are located and which adheres to the patient's skin to be worn for a period of time. A small needle within the device delivers a subcutaneous sensor that is often electrochemical.

[0005] Glucose readings taken by the sensor can be tracked and analyzed by a monitoring device, such as by scanning the sensor with a custom receiver or by transmitting the signal to a smartphone or other device with an associated software application. Software features already included in CGM systems include viewing glucose levels over time, indicating glucose trends, and alerting patients to high and low glucose levels.

[0006] Medical patients often suffer from diseases or conditions that require measurement and reporting of biological conditions. For example, if a patient suffers from diabetes, it is important for the patient to have an accurate understanding of the glucose level in his or her system. Traditionally, diabetics monitor glucose levels by pricking their fingers with a small spear, forming a drop of blood, and then dipping a test strip into the blood. The test strip is placed in a handheld monitor that analyzes the blood and reports the measured glucose level to the patient in a visual manner. Based on this reported level, the patient makes important health decisions about what food to eat or how much insulin to inject. Although it is beneficial for patients to check glucose levels multiple times throughout the day, many patients fail to fully monitor their glucose levels due to pain and inconvenience. Therefore, patients may eat improperly or inject too much or too little insulin. Either way, the patient's quality of life will be reduced, and the risk of permanent damage to their health and body will also increase. Diabetes is a devastating disease that, if not properly controlled, may lead to terrible physiological symptoms such as kidney failure, skin ulcers or eye bleeding and ultimately blindness, pain, and frequent limb amputations.

[0007] As is well known, blood sugar levels can rise significantly or drop rapidly due to several known and unknown reasons, which complicates glucose monitoring for patients. Therefore, a single glucose measurement only provides a short snapshot of the instantaneous glucose level in the patient. This single measurement provides little information about how the patient's glucose usage changes over time or how the patient responds to a specific dose of insulin. Therefore, even if the patient adheres to a strict finger prick and test strip testing program, the patient may make wrong decisions about diet, exercise, and insulin injections. Of course, the patient's test strip test inconsistencies can exacerbate this situation. In order to give patients a more comprehensive understanding of their diabetes symptoms and obtain better treatment results, some diabetics are now using continuous glucose monitoring.

[0008] The CGM sensor is typically temporarily adhered to the patient's skin using an adhesive pad, and the CGM sensor is coupled to a small housing in which the electronics are located. The CGM sensor typically has a disposable applicator device that uses a small guide needle to deliver the CGM sensor subcutaneously to the patient. Once the CGM sensor is in place, the applicator is discarded and the electronics housing is attached to the sensor. While the electronics housing is reusable and can be used for a long time, the CGM sensor and applicator need to be replaced frequently, usually every few days.

[0009] It should be understood that continuous glucose monitoring can be performed at different time intervals depending on the patient's specific medical needs. For example, some continuous glucose monitors can be set to take multiple readings per minute, while in other cases, the continuous glucose monitor can be set to take readings every hour or so. It should be understood that a continuous glucose monitor can sense and report glucose readings at different time intervals, and the reading rate can change based on past measurements, time of day, or other criteria.

[0010] Electrochemical glucose sensors operate by using electrodes that typically detect an amperometric signal caused by the oxidation of the enzyme during the conversion of glucose to gluconolactone. The amperometric signal can then be correlated to the glucose concentration. A two-electrode (also called a two-pole) design uses a working electrode and a reference electrode, where the reference electrode provides a reference against which the working electrode is biased. The reference electrode essentially completes the electron flow in the electrochemical loop. A three-electrode (or tripole) design has a working electrode, a reference electrode, and a counter electrode. The counter electrode replenishes the ion loss on the reference electrode and is part of the ion loop.

[0011] Unfortunately, the current cost of using a continuous glucose monitor is too high for many patients who can benefit greatly from its use. As generally described above, a continuous glucose monitor has two main components. First, a housing for electronics, a processor, a memory, wireless communications, and a power source. The housing is typically reusable and can be reused over a long period of time (such as several months). This housing is then connected or communicated to a disposable CGM sensor that adheres to the patient's body, and the sensor is inserted subcutaneously into the patient's body using a guide needle. This sensor must sometimes be replaced every three days, and may be replaced at least every other week. Therefore, the cost of purchasing a new disposable sensor represents a significant financial burden for both patients and insurance companies. Because of this, a large number of patients who may benefit from continuous glucose monitoring cannot use such systems and are forced to rely on less reliable and painful finger puncture monitoring. Summary of the invention

[0012] In some embodiments, the continuous glucose monitoring sensor comprises a working electrode, a reference electrode and a counter electrode. The working electrode has a first wire with a first plane and an electrochemical element on the first plane. The reference electrode has a second wire with a second plane, and the counter electrode has a third wire with a third plane. The first wire, the second wire and the third wire are used as sensor wires for the working electrode, the reference electrode and the counter electrode. The second plane and the third plane face each other.

[0013] In another embodiment, a novel working electrode for use in a continuous biosensor is disclosed. The working electrode uses a plastic substrate coated with a specially formulated carbon-containing compound. This carbon-containing compound is an aqueous dispersion of a carbon material in an elastomeric material. The carbon compound is applied to a plastic substrate, and then further films and coatings are applied to form the working electrode. The working electrode can then be connected to one or more reference electrodes or counter electrodes to form the biosensor.

[0014] In one example, the plastic substrate can be polyethylene, polypropylene, polystyrene, polyvinyl chloride or polylactic acid, and can be formed into an elongated wire. The carbon material can be, for example, graphene, diamagnetic graphite, pyrolytic graphite, pyrolytic carbon, carbon black, carbon paste or carbon ink, which is aqueously dispersed in an elastomeric material such as polyurethane, silicone, acrylate or acrylic acid. Optionally, selected additives can be added to the carbon compound before the carbon compound is layered on the plastic wire. These additives can, for example, improve conductivity or sensitivity, or act as a catalyst for target analyte molecules.

[0015] In one particular application, the plastic substrate is formed into an elongated wire and then coated with a carbon compound having a carbon material aqueously dispersed in an elastomeric material. Additives may be added to the carbon compound to act as a hydrogen peroxide catalyst, such as phthalocyanine or Prussian blue. Furthermore, the additive may be in the form of a metal oxide to enhance electrical properties, wherein preferred metal oxides are formed with copper, nickel, Rh or Ir.

[0016] Advantageously, the working electrode can be constructed to be durable, strong, flexible and have excellent electrical and sensitivity characteristics. In addition, since the working electrode can be constructed without expensive and rare platinum, a more cost-effective working electrode can be provided. This platinum-free electrode will be able to provide patients with cheaper sensors, allowing more patients to obtain the substantial benefits of continuous monitoring and, in particular, continuous glucose monitoring. This also allows greater flexibility in the mechanical design and construction of the sensor. In addition, this design also allows for use with other analytes / enzymes besides glucose, many of which require carbon-based electrodes for optimal performance.

[0017] In another embodiment, a sensor for a continuous biomonitor is disclosed having a working electrode with: (1) a new interference layer for enhancing and stabilizing the interaction of hydrogen peroxide with the conductor layer and (2) an enhanced glucose limiting layer formed by physical hydrogen bonds. Although these inventive aspects can be used independently, they are combined to form a highly desirable new working electrode and new sensor. The new sensor is easier and cheaper to manufacture than previous devices and provides improved sensitivity, better linearity and enhanced accuracy. Compared to previous working sensors, the new interference layer more accurately regulates the flow of hydrogen peroxide from the enzyme membrane to its conductor and provides greater interaction between hydrogen peroxide and the surface of the conductor. The new sensor also has an external protective glucose limiting layer formed using physical hydrogen bonds rather than providing chemical crosslinking.

[0018] In one example of the interference layer, the interference compound is electrodeposited on a conductive substrate, and the enzyme layer is applied on the interference compound. The interference compound is: 1) non-conductive, 2) ion-passing, and 3) has selective permeability according to molecular weight. In addition, it is electrodeposited in a thin and conformal manner, so that the flow of hydrogen peroxide from the enzyme layer to the conductive substrate can be more accurately controlled. In a specific example, the interference material is prepared by mixing a monomer with a mild alkaline buffer and then electropolymerizing the mixture into a polymer. For example, the monomer can be 2-aminophenol, 3-aminophenol, 4-aminophenol, aniline, naphthol, phenylenediamine or its blend, and the monomer is mixed with the buffer and electropolymerized into a polymer. It should be understood that other monomers can be used. In a more specific example, the monomer is 2-aminophenol, and the buffer is phosphate buffered saline (PBS) at pH about 8. The monomer is mixed with the buffer and electropolymerized into a polymer poly-o-aminophenol (PoAP). Then PoAp is electro-deposited on the conductive substrate. The permselectivity of PoAP can be adjusted by the pH of the buffer, for example by adding sodium hydroxide (NaOH).

[0019] In an example of the glucose limiting layer, 1) hydrophilic bonding material, 2) hydrophobic bonding material and 3) solvent are mixed together to form a bonding gel. The bonding gel is then applied to the enzyme membrane layer and the gel is solidified. The hydrophilic material is usually selected to have a high molecular weight, is easy to distribute and provides strong hydrogen bonding. In a specific example, the hydrophilic bonding material is polyvinyl pyrrolidone ( Polyvinylpyrrolidone )(PVP). The hydrophobic material is selected to be biocompatible and have sufficient hardness while still providing appropriate interaction with the hydrophilic material and solvent. Polyurethanes and silicones have been found to be ideal hydrophobic materials. Finally, the solvent is selected to be polar, binary and volatile enough to meet the curing requirements.

[0020] Advantageously, both the novel interference layer and the novel glucose limiting layer can be economically manufactured to provide a more cost-effective working electrode. In addition, both novel membranes provide enhanced linearity and overall detection characteristics for the working electrode. In an example, the interference layer is non-electronically conductive, ion-permeable and has selective permeability for molecular weight, and the glucose limiting layer is a self-crosslinking formulation of polyacrylic acid and acrylic polyurethane.

[0021] In another embodiment of the invention, the working wire has an enzyme layer comprising an aqueous emulsion of a polyurethane and GOx blend, which is applied to the working wire and cured. The new enzyme layer has better stability and complete retention of GOx, more uniform dispersion and enables higher GOx loading and better overall sensor sensitivity. It should also be understood that other enzymes can replace GOx in order to measure other metabolic functions.

[0022] In another embodiment of the invention, the working wire has a carbon-enzyme layer of an aqueous emulsion comprising a polyurethane, carbon and GOx blend, which is applied and cured into a plastic substrate for the working wire. The new carbon-enzyme layer has better stability and complete retention of GOx, more uniform dispersion and can achieve higher GOx loading and better overall sensor sensitivity. In addition, the carbon-enzyme layer can directly generate free electrons proportional to the amount of glucose reacted, thereby eliminating any need for expensive platinum. It should also be understood that other enzymes can replace GOx in order to measure other metabolic functions. BRIEF DESCRIPTION OF THE DRAWINGS

[0023] These and other objects and advantages of the present disclosure will become apparent after reading the following detailed description and referring to the drawings and claims.

[0024] Figure 1 Various views of planar electrodes are shown according to some embodiments.

[0025] Figure 2A-2D Various views of a planar electrode with a triangular support core are shown according to some embodiments.

[0026] Figure 3 An electrochemical element mounted into a planar electrode is shown according to some embodiments.

[0027] Figure 4A is a not-to-scale illustration of a carbon-coated wire for a working electrode, according to some embodiments.

[0028] Figure 4B is a not-to-scale illustration of a carbon-coated wire for a working electrode, according to some embodiments.

[0029] Figure 4C is a not-to-scale illustration of a carbon-coated wire for a working electrode, according to some embodiments.

[0030] Figure 5 is a flow chart of general manufacturing steps for preparing a working electrode according to some embodiments.

[0031] Figure 6 is a not-to-scale cross-sectional block diagram of a prior art single-wire sensor.

[0032] Fig. 7A is a not-to-scale cross-sectional block diagram of a 2-wire sensor with an interference membrane layer, according to some embodiments.

[0033] Figure 7B is a not-to-scale cross-sectional block diagram of a 2-wire sensor having an interference membrane layer and a coated reference electrode according to some embodiments.

[0034] Figure 8 is a flow chart of general fabrication steps for preparing a 2-wire sensor with an interference membrane layer according to some embodiments.

[0035] Fig.9A is a not-to-scale cross-sectional block diagram of a 2-wire sensor having an interfering membrane layer and a glucose limiting layer, according to some embodiments.

[0036] Fig. 9B is a not-to-scale cross-sectional block diagram of a 2-wire sensor having an interfering membrane layer, a glucose limiting layer, and a coated reference electrode, according to some embodiments.

[0037] Fig.10 is a flow chart of general manufacturing steps for preparing and applying a glucose limiting layer according to some embodiments.

[0038] Fig.11 is a flow chart of general manufacturing steps for preparing and applying an enzyme layer according to some embodiments.

[0039] Fig.12 is a not-to-scale cross-sectional block diagram of a 2-wire sensor having an enzyme layer, an interference layer, a glucose limiting layer, and a non-platinum substrate, according to some embodiments.

[0040] Fig.13A is a not-to-scale cross-sectional block diagram of a 2-wire sensor with a carbon / GOx membrane for direct generation of electrons or peroxides and an interference layer, according to some embodiments.

[0041] Fig. 13B is a not-to-scale cross-sectional block diagram of a 1-wire sensor having a carbon / GOx membrane for direct generation of electrons or peroxides and an interference layer, according to some embodiments.

[0042] Fig.14A is a not-to-scale cross-sectional block diagram of a 2-wire sensor with a carbon / GOx membrane for direct generation of electrons or peroxides, according to some embodiments.

[0043] Fig. 14B is a not-to-scale cross-sectional block diagram of a 1-wire sensor with a carbon / GOx membrane for direct generation of electrons or peroxides, according to some embodiments.

[0044] Fig.15 is a not-to-scale cross-sectional block diagram of a 1-wire sensor with a working lead having a carbon / GOx film for direct electron or peroxide generation and an attached reference lead, according to some embodiments.

[0045] Fig.16 is a flow chart of general fabrication steps for preparing and applying carbon / GOx membranes for direct generation of electrons or peroxides, according to some embodiments. DETAILED DESCRIPTION

[0046] The present disclosure relates to the structure and process of sensors for use in continuous metabolic monitors (such as continuous glucose monitors). In particular, the apparatus and methods of the present invention describe novel membranes and substrates for use with working electrodes in continuous metabolic sensors. For patients who may benefit from the use of CGM, cost may be a prohibitive factor. Therefore, there is a great need in the market for low-cost sensors for continuous biological monitors. It should be understood that cost reduction can be obtained by reducing the manufacturing cost of the sensor itself, by increasing the length of time between sensor replacements, or by a combination of reducing costs and increasing service life. By reducing the cost of sensors for continuous monitoring, more patients can benefit from improved quality of life and enhanced continuous monitoring treatment effects.

[0047] Most CGM sensor designs are planar (flat substrate) or wire-based. The planar type is more suitable for 3-pole electrochemical designs because simple wire traces and small electrodes can be easily constructed. However, the planar type has physiological defects because the planar substrate has some directionality due to its geometry and also has sharp edges, which produces a more aggressive biological response to the device. Wire-based systems produce better physiological responses from patients than planar systems due to the smooth nature of their geometry, but they are mainly limited to a single wire that is easy to insert through a needle. This single-wire constraint caused by the space limitations of needle-based sensor delivery usually limits the design to a 2-pole electrochemical design. The 2-pole design has the additional disadvantage of making the reference electrode non-renewable, and therefore consumes electrode material to complete the electrochemical circuit, which limits the working life of the system.

[0048] The challenge of wire-based sensor design is to make electrical connections on the distal end. Single-wire configurations require in situ fabrication of working membranes and chemistries, and thus limit the methods and materials that can be used in such designs. For ease of fabrication, separate wires for the working, reference, and counter electrodes are ideal; however, this approach is limited by the inner diameter of the insertion needle.

[0049] Embodiments of the present invention disclose a wire-based 3-pole electrochemical design that addresses the deficiencies of the above-described designs. The working chemistry is prepared separately from the wire and then bonded to the underlying sensor wire. Because the components of the CGM device of the present invention can be prepared independently of each other, this allows for lower cost materials and methods. Moreover, because the wires are manufactured individually, 100% sensor quality testing is not required and quality testing can be performed on a sheet or batch basis, more cost-effective large-scale manufacturing can be achieved. Some embodiments of the disclosed wire-based system use carbon-based (such as graphene-based) electrodes manufactured in large-scale sheets with working chemicals, which are then attached to the working electrode.

[0050] Wire-based 3-pole electrode design

[0051] Figure 1 An embodiment of a wire-based 3-pole system 100 for a continuous glucose monitoring sensor using a split wire design is shown. In this embodiment, a portion of the wire (such as a half wire) is provided for a reference electrode 110 and a counter electrode 120, each electrode having a plane approximately passing through its diameter so that the wire has a semicircular cross-section. In some embodiments, the plane of the reference electrode 110 and the plane of the counter electrode 120 face each other. Each half-wire electrode, such as the reference electrode 110 and the counter electrode 120, can have a partial surface area, such as 82% of the surface area of ​​a full wire with the same diameter, while still allowing the reference electrode and counter electrode assembly to fit into a small diameter insertion needle 102 for insertion under the skin. In other words, the split wire configuration enables the reference electrode 110 and the counter electrode 120 to provide almost the same surface area as two full wire electrodes, but only occupy the space of one wire in the insertion needle 102 instead of the space of two full wires. Although half wires are depicted for reference electrode 110 and counter electrode 120 (where each wire has been split along its diameter along the length of the wire), other fractions of the wire may be utilized to form planar electrodes, such as, for example, 30% to 70% or 40% to 60%, typically defined by a chord stretched through the circular cross-section of the wire to yield a greater proportion of the full wire surface area.

[0052] The working electrode is also fabricated by forming a flat portion on the wire. Figure 1Two embodiments are shown, namely, 1-side working electrode 130 and 2-side working electrode 135, either of which can be used. 1-side working electrode 130 has a semicircular cross-section in which half of the cross-sectional area of ​​the wire has been removed, while 2-side working electrode 135 has a rectangular cross-section in which portions of the wire above and below the flat portion have been removed. The removed portions may be equal or one portion (i.e., the top or bottom) may be larger than the other. The flat portion of working electrode 130 or working electrode 135 is used to support an electrochemical element, which is a reactive component that senses glucose in the patient's interstitial fluid.

[0053] Figure 1 Insertion of the electrodes into the insertion needle 102 is also shown, where it can be seen that this 3-pole design of the sensor occupies the space equivalent to only two wires instead of three wires in the needle cavity. The working electrode (where the 2-side working electrode 135 is shown in this illustration) utilizes the space of one wire, and the reference electrode 110 and the counter electrode 120 together occupy the space of another wire. The diameter of the wire used for the reference electrode 110, the counter electrode 120, or the working electrode 135 can be, for example, 0.002 inches to 0.007 inches. The length or surface area of ​​the electrode portion itself can be customized according to the desired sensor sensitivity and the desired design specifications.

[0054] exist Figure 2A-2D Other embodiments of systems using planar electrodes in continuous glucose monitoring sensors are shown in FIG. Figure 2A ) and Design 210( Figure 2B ), the compact system is assembled from a supporting core wire 140 having a triangular cross section surrounded by a working electrode 135, a reference electrode 110, and a counter electrode 120 facing the plane of the triangular core wire 140. Each wire for the working electrode 135, the reference electrode 110, and the counter electrode 120 has a plane positioned to face the surface of the triangular core wire 140. Figure 2A-2B As shown, the cross-sections of the reference electrode 110 and the counter electrode 120 are approximately semicircular, while the cross-section of the working electrode 135 can be semicircular ( Figure 2A Design 200) or rectangular ( Figure 2B Design 210). Schematic diagram 220 ( Figure 2C ) and schematic diagram 230 ( Figure 2D ) provide a longitudinal cross-sectional view and a perspective view, respectively, of the end of the triangular support wire, which shows that this triangular design can be a completely self-inserting sensor. That is, the tip 142 of the triangular core wire 140 can be sharpened to a point or become pointed so that the sensor can be directly inserted without the need to use a needle to place the sensor in the subcutaneous tissue.

[0055] In these various embodiments, the flat surface of the electrode provides support for a fragile electrochemical material, such as a typically brittle carbon-based sheet. In one example, a support sheet (e.g., made of pyrrole or polyaniline) can be formed, and then the carbon material is deposited onto the support sheet. The support sheet provides a substrate to which the carbon bonds well, and should also be conductive to electrically couple the electrochemical (e.g., carbon / pyrrole) sheet to the electrode wires. The conductive sheet material can then be impregnated or coated with the sensing chemical by various stretch film or spin coating techniques.

[0056] The electrochemical material sheet can be prepared separately from the electrode wire and then as Figure 3 As shown, it is mounted on the flat surface of the electrode. Figure 3 In this example, a wire 310 having an insulator 320 surrounding a conductive core 330 has a portion of its end removed to form a flat surface 340. Carbon or carbon / graphene / pyrrole sheets 350 are cut to size and placed on the flat surface 340 of the flat electrode wire 310. For example, once the flat sheets are made with the sensing chemistry, these sheets 350 can be laser cut into small sections and then assembled onto the flat surface 340 of the wire 310.

[0057] The support sheet can be prepared, for example, by depositing a pyrrole layer to make electrical contact with the flat surface of the electrode. In other embodiments, electropolymerization of additional pyrrole can be used to connect the electrode metal to the sheet, or conductive adhesives or other electrical contact bonding methods can also be used to form the electrical contact.

[0058] In other embodiments, the electrochemical component can be formed in situ on the electrode rather than forming a sheet separate from the electrode. For example, alternative manufacturing methods for in situ generation of the sensing chemistry and membrane can include pad printing or screen printing, painting, or 3D printing directly on the plane of the wire.

[0059] The carbon material may be in the form of, for example, an ink or paste, and the carbon may include various allotropes such as, but not limited to, graphite, graphene, fullerenes, and / or nanotubes. Materials other than pure carbon may be used, including platinum black, carbon platinum paste, carbon gold paste, or other known working electrode surface materials (e.g., carbon, platinum, gold, palladium, rhodium, iridium) used alone or in combination. In some embodiments, high surface area nanoporous materials of graphene and / or other nanomaterials may be used to increase the number of active chemical sites available for reaction.

[0060] Carbon is less expensive than metals (e.g., gold and platinum) that are commonly used in biocompatible applications. However, due to the inherent brittleness of carbon materials, carbon-based electrodes are conventionally used in planar electrodes (such as fingersticks), where the carbon can be supported by a planar substrate without placing undue mechanical loads on the electrode. Embodiments of the present invention overcome the difficulties of using carbon-based materials on wire electrodes by providing the required mechanical support for the carbon material and by eliminating the typical need to manufacture the working chemistry in situ on the wire (although in situ manufacturing can be used).

[0061] After generating the sensing chemistry either separately or in situ on the electrodes, a final dip coating can be used to seal the entire system using the hoop strength created by shrinkage of the polymer after drying. This final polymer layer also serves as the biocompatible and glucose limiting membrane required to generate a linear glucose response and provides the biosafety required for implantable sensors.

[0062] Flat wire embodiments of the present invention can also be used to optimize electrochemical substrates so that they can be tuned for direct electron transfer chemistry by keeping the redox centers near the porous carbon surface or within the encapsulated polymer. One such embodiment uses aminophenols covalently bonded to a carbon electrode by electrografting and then linked to glucose oxidase (GOx) by diazo chemistry to provide direct electron transfer. Embodiments can be used directly with conductive polymers (e.g., PEDOT-PSS, polypyrrole, polyaniline, naphthol, phenylenediamine, etc.) formed in situ on porous carbon sheets, which can act with normal enzymes (glucose oxidase (GOx) or glucose dehydrogenase (GDH)) and / or enzymes with mediators to create hybrid enzyme systems that alter the need for high bias and thus reduce interference from all sources.

[0063] In some embodiments, the oxidoreductase can be immobilized on the electrode surface in a new way so that direct electron transfer can occur between the active side of the enzyme and the transducer. The main unique feature of such an embodiment of the amperometric glucose sensor is that its bias potential is in the range of 0V to -0.5V, ideally to about -0.1V. In contrast, conventional CGM sensors have a bias potential of typically +0.55V. There are two main methods for realizing the lower bias potential of the design of the present invention. The first method is to electropolymerize the conductive polymer with the oxidoreductase in situ. The sensing layer is formed by applying a potential cycle or a sequence of suitable potential pulses with an enzyme and a monomer / comonomer solution. The advantage of this method is that the film is formed only on the electrode surface due to the electrochemical initiation of the deposition process. The second method is to incorporate the redox mediator into a polymer or prepolymer. The polymer containing the redox mediator can be physically mixed with the enzyme and then deposited on the electrode by dip coating, spin coating or other coating methods. This can also be achieved by in situ polymerization of the prepolymer containing the redox mediator with other active prepolymers in the presence of an enzyme solution and an electrode. The resulting sensing layer on the electrode comprises the matrix enzyme inside the polymer network with the covalently attached redox mediator.

[0064] Carbon substrate

[0065] In some embodiments, a cost-effective platinum-free sensor is used in a continuous biomonitoring system. The embodiments provide significant cost reductions for the manufacture of working electrodes for such biosensors. Although embodiments for continuous glucose monitoring are primarily discussed, it should be understood that there are many other uses for biosensing that benefit from cost-reduced sensors and working electrodes.

[0066] Typically, sensors for continuous biomonitoring systems have a working electrode and a reference electrode. The working electrode and the reference electrode are constructed and arranged so that they can sense the concentration of an analyte in a patient's body, often by measuring the concentration or ion flow in blood or other body fluids, such as interstitial fluid (ISF). It should be understood that the sensor may include multiple working wires, multiple reference electrodes, and counter electrodes.

[0067] Typically, a working electrode needs to be constructed to meet three basic requirements. First, it must be strong enough to withstand insertion under the patient's skin and withstand vibration, shock, and movement during use. Second, it needs to be flexible enough to enter the skin along a curved path and allow some movement after insertion for patient comfort. And third, it needs to provide electrical properties to support consistent and accurate sensing. Therefore, known working electrodes typically use some form of platinum wire, i.e., solid platinum wire or a cheaper metal material (such as tantalum) coated with platinum. It is this reliance on and use of platinum that drives some of the high-cost current biosensors.

[0068] Advantageously, embodiments of the present disclosure eliminate the need for expensive and rare platinum to prepare working electrodes that not only have sufficient mechanical strength and flexibility but also have excellent electrical and sensing properties. In addition, embodiments of the working electrodes of the present invention are constructed from materials that are known to be safe for the human body. This also allows the use of alternative geometries of sensors and different styles of sensor manufacturing.

[0069] In one particularly cost-effective embodiment, the working electrode uses a plastic material as a substrate. The plastic material is strong enough to support insertion into the human body while having the flexibility required for insertion and patient comfort. This plastic substrate can be formed into elongated wires of many shapes to support the construction of different types of sensors. The plastic wire can then be coated with a specially formed carbon compound. Plastic wire has the additional advantage of improved fatigue performance compared to metal wires of the same size. Traditionally, elemental carbon paste electrodes are not considered for use on flexible working electrodes because carbon is very brittle and requires rigid support. Moreover, carbon paste electrodes are typically water soluble and therefore dissolve and degrade when inserted into a humid environment. And finally, carbon has a high electrical resistance compared to platinum metal and is therefore not practical for use as a conductor in a biosensor. However, the new form of carbon in the carrier compound used on the plastic wire as disclosed herein overcomes several disadvantages of elemental carbon.

[0070] In some embodiments, the carbon compound is prepared as an aqueous dispersion coating of a carbon material and an elastomeric material. For example, the carbon material can be in the form of graphene, diamagnetic graphite, pyrolytic graphite, pyrolytic carbon, carbon black, carbon paste, or carbon ink. In some cases, in order to support specific applications, other additives can be added to the carbon compound for enhancing electrical and reaction characteristics. For example, a hydrogen peroxide catalyst can be added to the carbon compound to support enhanced glucose level sensitivity. It should be understood that other sensing molecules can be used for other sensing applications.

[0071] The carbon compound as described above is then applied to the plastic wire. Most often this is done through a simple dipping process, although it should be understood that the coating can also be sprayed, extruded, deposited or even printed onto the plastic wire or directly 3D printed onto the substrate. The coated wire can then be processed into a working electrode using known methods by adding a membrane, connecting it to a reference electrode and adding a protective bio-coating.

[0072] Reference now Figure 4A , Figure 4B and Figure 4C , shows a carbon coated wire 400. The carbon coated wire 400 includes Figure 4A Wire 411 in Figure 4B The wire 412 and Figure 4CThe wires 413 in the figure are not to scale and are for illustration purposes only. Each of the carbon coated wires 400 has a plastic core 415 completely surrounded by a carbon compound 418. It should be understood that the plastic core 415 can be formed into many different elongated physical shapes. For example, Figure 4A As shown, the plastic core 415 may have a circular cross-section. Figure 4B As shown, the plastic core 415 may have a rectangular cross section or a square cross section. Figure 4C As shown, the plastic core 415 may have a triangular cross-section. It should be understood that many other cross-sectional shapes may be used.

[0073] Carbon compound 418 is formulated to have excellent electrical properties, appropriate mechanical properties (such as strength and flexibility) and cost-effective. For example, standard carbon conductive ink has a resistivity of about 23Ohm / mm2, while carbon compound 418 can be formulated to have a more ideal resistivity, such as 1-5Ohm / mm2. In this way, it has been found that carbon compound 418 has a resistivity that is one order of magnitude lower than standard carbon conductive ink, which significantly improves its utility and performance as a conductor for working wires. Carbon coating 418 is not only much cheaper than platinum, but it is also easier and more cost-effective to apply in the form of a coating. For example, carbon compound 418 can be used with low-cost immersion, spraying, extrusion, deposition or printing processes. It should be understood that, depending on the specific application, carbon coated wire 400 will be further processed to add membranes and protective coatings, and they will be connected to one or more reference electrodes or counter electrodes. It should be understood that the connection of the working wire to the reference wire can be accomplished in several ways. For example, the working and reference wires may be placed side-by-side, formed concentrically, twisted in a twisted relationship, layered, or formed in any other known physical relationship of a working wire and its associated reference wire.

[0074] In one example, the carbon coating can be formulated as follows. It should be understood that many other formulations fall within the teachings herein.

[0075] Formulation (wt%, total 100%)

[0076] An aqueous dispersion comprising:

[0077] 40%-60% polyurethane (i.e. Hauthaway HD4661,

[0078] 40%-60% acrylic polyol (Acquathane),

[0079] 0.5%-5% polyvinyl pyrrolidone).

[0080] 0.1%-0.5% Carbon Black

[0081] 0.05%-0.5% Graphene

[0082] 0.1%-0.5% pyrolytic graphite

[0083] 0%-10% additional water

[0084] Reference now Figure 5 , a process 500 for preparing a carbon working electrode is shown. The process 500 begins with selecting a plastic substrate material in step 522. This plastic substrate material is selected to have sufficient strength for insertion under the patient's skin, as well as flexibility for patient comfort and ease of manufacture. In addition, it should be understood that the plastic substrate should be biologically safe and generally non-electrically reactive. It should be understood that a wide range of materials meet the mechanical and functional requirements of the selected plastic substrate. For example, many organic polymers and thermoplastics can be used. For illustrative purposes only, the following specific plastic substrate materials can be used: polyethylene, polypropylene, polystyrene, polyvinyl chloride, and polylactic acid. It should be understood that a wide variety of materials can be used as plastic substrates.

[0085] The selected plastic substrate material is then formed into an elongated plastic wire in step 523. It should be understood that the wire can take many cross-sectional shapes, such as circular, square or triangular. Typically, these wires can be formed using well-known extrusion processes. The plastic substrate can also be formed into a ribbon wire, or in some cases manufactured by printing such as 3D printing.

[0086] In step 525, a carbon compound is prepared for application to a plastic substrate. The carbon compound has a carbon material aqueously dispersed in an elastomeric material. The elastomeric material is selected for its mechanical properties, such as strength and flexibility, and the carbon material is selected for favorable electrical properties. There are several acceptable elastomeric materials that can provide the desired properties, such as polyurethane, silicone, acrylate, or acrylic. It should be understood that other elastomeric materials may be substituted. Experimental results related to the present disclosure show that the carbon compound does not delaminate from the plastic wire after the elastomeric material is cured, compared to tantalum wire coated with platinum.

[0087] The carbon material in the carbon compound is selected for enhanced electrical properties. For example, as briefly discussed above, elemental carbon has too high an electrical resistance to be effectively used in a working electrode. However, by adding graphene, diamagnetic graphite, or pyrolytic carbon, the carbon compound can be formulated to have favorable electrical properties. In fact, the loading of the carbon material with the elastomeric material can be adjusted to produce a carbon compound with a desired resistance (e.g., 100 ohms / cm2 or less). In this way, a working electrode using such a carbon compound coating can be used so that the sensing system has a very ideal signal-to-noise ratio. Elemental carbon cannot achieve this signal-to-noise ratio due to the high electrical background from the high resistance of the carbon electrode.

[0088] Optionally, additional catalysts or materials may be added to the carbon compound in step 526 to enhance electrical or sensing properties. For example, metal oxides may be added to the carbon compound for reducing resistivity, thereby providing the working electrode with higher signal-to-noise capabilities compared to elemental carbon. For example, in some embodiments, metal oxides of nickel or copper may be used. In some embodiments, metal oxides of Rh and Ir, when added to the carbon compound, may enable the working conductor to operate at a lower bias voltage compared to a conductor formed with platinum. By operating at a lower bias voltage, the working conductor can operate with higher sensitivity and lower energy consumption.

[0089] In another example of an additive in step 526, a hydrogen peroxide catalyst may be added to the carbon compound. In one example, phthalocyanine or Prussian blue is added to the carbon compound, thereby substantially increasing the sensitivity of the working lead to hydrogen peroxide, which is very beneficial to the overall accuracy and sensitivity of the glucose monitor sensor. It should be understood that other hydrogen peroxide catalysts may be used. Moreover, for working leads intended for biosensing other than glucose, it should be understood that other sensing molecules and molecular catalysts may be used.

[0090] The coating may then be applied to the plastic wire substrate in step 527. Because carbon compounds are cheap and readily available, the plastic substrate may be dipped into the carbon compound. For other applications, the carbon compound may be sprayed onto the plastic wire, may be deposited using well-known deposition processes, co-extruded, or may be applied using a printing process such as pad printing. It may also be prepared to be 3D printed. It should be understood that any suitable application process may be used to coat or deposit the carbon compound onto the plastic substrate. The carbon compound coating is then cured prior to further processing.

[0091] Once the carbon coated working wire has been cured, it can be processed into a working electrode in step 528. In this way, membranes and protective coatings can be added and the working wire is connected to one or more reference electrodes or counter electrodes. The process for adding membranes, protective coatings and connecting other electrodes is well known and will not be described herein. For example, the working electrode and reference electrode can be placed side by side, layered, concentrically formed or twisted together. It should also be understood that some applications will use multiple working electrodes, multiple reference electrodes or counter electrodes.

[0092] Due to the aqueous nature of carbon-containing compounds, enzymes or other sensing molecules and chemicals can be included directly in the carbon-containing compounds, thereby increasing the efficiency of electron transfer and further improving the signal-to-noise ratio by removing additional layers and diffusion distances. Incorporating enzymes and other sensing chemicals into the sensor wire itself also further simplifies the manufacture of these sensors.

[0093] Reference now Figure 6 , a prior art single-wire sensor 600 for a continuous biomonitor is shown. Those skilled in the art will recognize that sensor 600 is a high-level diagram for instructional purposes only, and that a great deal of detail has been omitted to facilitate increased understanding. As will be appreciated, such prior art sensors will integrate the functionality of a working electrode with the functionality of a reference electrode on a single wire. It will be appreciated that a single-wire electrode may be constructed using multiple working electrode layers and multiple reference electrode layers. A single-wire electrode may also use or be supplemented with a counter electrode. Although sensor 600 is shown as a wire having concentrically formed layers, it will be appreciated that other physical embodiments may be used, such as layered, spiral, flat, and other well-known physical relationships.

[0094] The prior art sensor 600 has an elongated wire 605, which is often made of solid platinum or a platinum coating on a cheaper metal substrate or plastic substrate. It should be understood that other types of wires may be substituted. The wire 605 is wrapped with an electrical insulating layer 614a. During manufacturing, a strip 618 of the insulating layer 614a is removed, which exposes a portion 617 of the platinum wire that remains uninsulated. The removal of this strip 618 must be done very accurately and precisely because it affects the overall electrical sensitivity of the sensor 600. For example, this strip 618 can be approximately 20μm thick and must be cut to approximately 40μm, although other thicknesses and widths can be used depending on the overall structure of the sensor 600.

[0095] A layer of silver or silver chloride 611 is positioned around the electrically insulating layer 614, and a second layer of electrically insulating material 614b is disposed around the silver / silver chloride layer 611. During manufacturing, a portion 621 of the silver / silver chloride layer 611 needs to be exposed. Typically, this requires the precise removal of a small portion of the layer 614b using, for example, a laser ablation process. A second removal process may also be used at the connection end of the sensor 600 to expose a small portion of the silver / silver chloride layer 611 so that a more convenient electrical connection can be formed. Removing the insulating layer from the silver / silver chloride layer 611 is a precise operation as the layer may be only about 20 μm thick. This expensive removal operation adds a significant amount of cost and manufacturing risk to the preparation of the single-wire sensor 603.

[0096] In operation, glucose limiting membrane 607 substantially limits the amount of glucose that can reach enzyme membrane 608. By limiting the amount of glucose that can reach enzyme membrane 608, the linearity of overall reaction is improved. Glucose limiting membrane 607 also allows oxygen to enter enzyme membrane 608. The key chemical process for glucose detection occurs in enzyme membrane 608. Usually, enzyme membrane 608 has one or more glucose oxidases (GOx) dispersed in enzyme membrane 608. When a glucose molecule is combined with an oxygen molecule (O2) in the presence of glucose oxidase, a gluconic acid molecule and a hydrogen peroxide molecule (H2O2) are formed. Then, hydrogen peroxide is usually dispersed in enzyme membrane 608 and dispersed in ion conductive layer 609.

[0097] At least some of the hydrogen peroxide enters the window (strip 618) in the electrically insulating layer 614a where it contacts the exposed portion 617 of the platinum wire 605. The platinum surface promotes a reaction in which the hydrogen peroxide reacts to produce water and hydrogen ions that are released into the ion conductive layer 609, and two electrons are produced. Electrons are attracted to the platinum wire 605 by a bias voltage set across the platinum wire 605 and the silver / silver chloride layer 611. Positive ions from the silver / silver chloride layer 611 are released into the ion conductive layer 609 to complete the circuit. In this way, the magnitude of the current on the platinum wire is intended to be related to the number of hydrogen peroxide reactions, which is intended to be related to the number of oxidized glucose molecules. In this way, the current measurement on the platinum wire is intended to be associated with a specific glucose level in the patient's blood or ISF.

[0098] Unfortunately, because the platinum surface (portion 617) has been exposed during manufacturing, an oxide layer is formed in the window 618. This oxide layer contaminates the electrode and interferes with the efficiency of the exposed platinum in converting hydrogen peroxide. That is, the actual useful exposure area of ​​the exposed portion 617 of the platinum wire is greatly reduced due to oxidation contamination, which may also lead to unpredictable and undesirable sensitivity results. In order to overcome this defect, the single-wire sensor 603 must be calibrated in a complex manner. In addition, the bias voltage between the platinum wire 605 and the silver / silver chloride layer 611 must be set to be relatively high, for example, between 0.4V-1.0V. This high bias voltage is needed to attract electrons into the platinum wire, and it is used to attract contaminants from blood or ISF into the sensor. These contaminants such as acetaminophen and uric acid interfere with the chemical reaction, thereby causing erroneous and misleading glucose level readings. The manufacture of the single-wire sensor 603 is also expensive, partly due to the precise laser ablation required to expose the band 618 in the platinum wire 605, and the small portion 621 of the silver / silver chloride layer 611.

[0099] Embodiments of the present disclosure relate to cost-effective sensors for use in continuous biomonitoring systems. Embodiments provide significant cost reductions for the manufacture of working electrodes for such biosensors. Although embodiments for continuous glucose monitoring are primarily discussed, it should be understood that there are many other uses for biosensors that benefit from cost-reduced sensors and enhanced functional working electrodes.

[0100] Typically, sensors for continuous biomonitoring systems are constructed as two-wire continuous biosensors with a working electrode and a reference electrode. The working electrode and the reference electrode are constructed and arranged so that they can sense the concentration of an analyte in a patient's body, often by measuring a concentration molecule, such as glucose, in blood or other body fluids such as ISF. It should be understood that the sensor may include multiple working wires, multiple reference electrodes, and counter electrodes.

[0101] Interference layer

[0102] Reference now Fig. 7A , a sensor 700 for a continuous biomonitor is generally shown. The sensor 700 has a working electrode 703 that cooperates with a reference electrode 705 to provide an electrochemical reaction that can be used to determine the glucose level in the patient's blood or ISF. Although an electrode sensor 700 with one working electrode 703 and one reference electrode 705 is shown, it should be understood that some alternative sensors may use multiple working electrodes, multiple reference electrodes, and counter electrodes. It should also be understood that the sensor 700 may have different physical relationships between the working electrode 703 and the reference electrode 705. For example, the working electrode 703 and the reference electrode 705 may be arranged in layers, in a spiral, concentrically, or side by side. It should be understood that many other physical arrangements may be consistent with the disclosure herein.

[0103] The working electrode 703 has a conductive portion, which is shown as a conductive portion of the sensor 700 in the form of a conductive wire 710. This conductive wire 710 can be, for example, solid platinum, a platinum coating on a cheaper metal or plastic, or, as disclosed above, the conductive wire 710 can be a carbon compound coating on a plastic substrate. It should be understood that other electronic conductors can be used consistent with this disclosure. Like the working electrodes of the prior art, the working electrode 703 has a glucose limiting layer 707 that can be used to limit contamination and limit the amount of glucose received into the enzyme membrane 708.

[0104] In operation, glucose limiting membrane 707 substantially limits the amount of glucose that can reach enzyme membrane 708, for example, only allows 1 glucose molecule in about 1000 glucose molecules to pass. By strictly limiting the amount of glucose that can reach enzyme membrane 708, the linearity of the overall reaction is improved. Glucose limiting membrane 707 also allows oxygen to enter enzyme membrane 708. The key chemical process for glucose detection occurs in enzyme membrane 708. Usually, enzyme membrane 708 has one or more glucose oxidases (GOx) dispersed in enzyme membrane 708. When a glucose molecule is combined with an oxygen molecule (O2) in the presence of glucose oxidase, a gluconic acid molecule and a hydrogen peroxide molecule are formed. Then, hydrogen peroxide is usually both dispersed in enzyme membrane 708 and dispersed in interference membrane 709.

[0105] The interference membrane 709 is layered between the conductive wire 710 in the working electrode 703 and the enzyme membrane 708. As will be discussed in more detail below, in contrast to conventional insulating layers (e.g., Figure 6 The interference film 709 can be uniquely configured to have a more accurate regulation of the level of hydrogen peroxide molecules, enabling it to be transferred from the enzyme membrane layer 708 to the wider surface area of ​​the conductive wire 710. This interference film 709 can be electrodeposited on the conductive wire 710 in a very consistent and conformal manner, thereby reducing manufacturing costs and providing a more controllable and repeatable layer formation. The interference film 709 is non-conductive electron, but will pass negative ions at a preselected rate. In addition, the interference film 709 can be configured to have selective permeability to specific molecules. In an example, the interference film 709 is prepared and deposited in a manner that limits the passage of larger molecules, and the larger molecules can serve as pollutants that degrade the conductive layer 710, or can interfere with electrical detection and transfer processes.

[0106] Advantageously, the interference membrane 709 provides reduced manufacturing costs compared to known insulating layers, and enables more precisely regulated broad surface area for hydrogen peroxide molecules to pass to the underlying conductive layer 710. In addition, the formulation of the interference membrane 709 can be customized to allow for limiting or denying the passage of certain molecules to the underlying layer, for example, limiting or denying the passage of large molecules or specific target molecules.

[0107] Interference film 709 is a solid coating around platinum wire 710. In this way, the cost and uncertainty of providing a window through the insulating layer are avoided. Therefore, interference film 709 can be accurately coated or deposited on platinum wire 710 in a predictable and consistent way for hydrogen peroxide to pass. In addition, the allowable interaction area between hydrogen peroxide and the surface of platinum wire 710 is significantly increased because the interaction can occur anywhere along platinum wire 710. In this way, interference film 709 improves the interaction level between hydrogen peroxide molecules in the surface of platinum wire 710, so that the generation of electrons is fully amplified on the working electrode of the prior art. In this way, the interference film enables the sensor to operate at a higher electronic current, thereby reducing the sensitivity of the sensor to noise and interference from pollutants, and further enables the use of less complex and less accurate electronic devices in the housing. In a non-limiting example, the ability to operate with a higher electronic current allows the electronic devices of the sensor to use more standard operational amplifiers (op-amps), rather than the expensive precision op-amps required for the sensor system of the prior art. The resulting improved signal-to-noise ratio allows the realization of simplified filtering and streamlined calibration.

[0108] Additionally, during the manufacturing process, oxidation on the outer surface of the platinum wire 710 can be removed prior to depositing the interference film 709. Because the interference film 709 is used to seal the platinum wire 710, the oxidation level can be significantly reduced, again allowing for a larger interaction surface and further amplifying the glucose signal, thereby generating a higher electron flow and achieving a higher signal-to-noise ratio. In this way, the new interference layer prevents electrical interface fouling of the platinum by eliminating undesirable oxidation.

[0109] In some embodiments, the interference membrane 709 is non-conductive electron-conducting, but is conductive ionic. In practice, a particularly effective interference membrane can be constructed using, for example, poly-o-aminophenol (PoAP). PoAP can be deposited on the platinum wire 710 using an electrodeposition process to precisely control the thickness of the hydrogen peroxide that can pass through the interference membrane 709 to reach the platinum electrode 710. In addition, the pH value of the PoAP can be adjusted to set the ideal permeability selectivity of the interference membrane 709. For example, the pH can be advantageously adjusted to significantly block the passage of larger molecules such as acetaminophen, thereby reducing the pollutants that can reach the platinum wire 710. It should be understood that other materials, such as polyaniline, naphthol or polyethylene diamine, can be used.

[0110] The sensor 700 also has a reference electrode 705 that is separate from the working electrode 703. In this way, the manufacture of the working electrode is simplified and can be performed with consistency that helps to significantly improve stability and performance. The reference electrode 705 is constructed of silver or silver chloride 714.

[0111] Reference now Figure 7B, another sensor 701 for a continuous biomonitor is shown. Sensor 701 is similar to sensor 700 and will not be described in detail. Sensor 701 has a working electrode 703 that is the same as the working electrode described with reference to sensor 700. However, sensor 701 has a reference wire 725 that has a silver / silver chloride layer 726 surrounded by an ion limiting membrane 728. The application of this ion limiting membrane 728 on the silver / silver chloride layer 726 ideally controls the current sensitivity of the overall sensor device 701 by controlling the ion flow from the silver / silver chloride layer 726. In this way, the current sensitivity can be advantageously controlled and defined. As should be understood, this can also serve as a secondary method for controlling sensor sensitivity by controlling chloride release from the electrode surface.

[0112] Reference now Figure 8 , a general description of a process 800 for preparing and applying an interference film is shown. As shown in step 802, a conductive substrate is provided. This conductive substrate can be in the form of an elongated wire, but it should be understood that the conductive substrate can be provided in other forms, such as printed or in the form of a conductive pad. In some embodiments, the conductive substrate is a solid platinum wire, a cheaper wire that has been coated with platinum, or as disclosed herein, the conductive substrate can be a conductive carbon compound coated on a plastic substrate. It should be understood that other conductive substrates can be used.

[0113] As shown in step 804, an interfering membrane compound is now prepared. This compound is formulated to be 1) non-conductive; 2) ion-permeable; and 3) selectively permeable. In addition, the compound is specifically formulated to be electrodeposited in a thin and uniform layer, and the compound has a self-limiting thickness due to the electrically driven cross-linking nature. In this way, the compound can be applied using a simple and cost-effective manufacturing process in a manner that provides a well-controlled regulation of the passage of hydrogen peroxide molecules. In addition, the passage of hydrogen peroxide can occur over a much larger surface area than the working conductors of the prior art.

[0114] Generally, the above identified properties of the interfering membranes of the present invention can be formulated by mixing the monomers with a mild alkaline buffer and converting the monomers to more stable and usable polymers by applying an electropolymerization process. In one formulation:

[0115] a) Monomers: for example, 2-aminophenol, 3-aminophenol, 4-aminophenol, aniline, naphthol, phenylenediamine or a blend thereof.

[0116] b) Buffer: Phosphate buffered saline (PBS) adjusted to about pH 7.5 to about pH 10, such as pH 7.5 to pH 9, such as pH 8, for example by adding sodium hydroxide.

[0117] c) The monomers are mixed with a buffer and electropolymerized.

[0118] d) Production of polymers; for example poly-o-aminophenol (PoAP).

[0119] In the particular formulation set forth above, 2-aminophenol monomer is mixed with a PBS buffer that is mildly alkaline at pH 8. The pH of the PBS buffer is adjusted using an additive such as sodium hydroxide. It should be understood that the pH can be adjusted to produce alternative formulations consistent with this disclosure. For example, the pH of the compound can be adjusted so that the permeation selectivity of the resulting PoAP can be modified. More particularly, the PoAp can be formulated to have a defined cutoff molecular weight. That is, by adjusting the pH of the formulation, the PoAP can be modified to substantially limit the passage of molecules having a molecular weight greater than the cutoff molecular weight. Thus, the PoAP is modified as needed to limit the molecular weight of contaminants reaching the platinum wire. It should also be understood that other monomers can be selected and that these alternative monomers can provide desired functional properties at different pHs. The 2-aminophenol and PBS mixture is electropolymerized into poly-o-aminophenol (PoAP).

[0120] Optionally, as shown in block 805, oxides or oxide layers may be removed from the surface of the conductive platinum substrate. As previously described, these oxides or oxide layers significantly limit the surface area available for the hydrogen peroxide to react with the platinum. By removing these oxides or oxide layers, such as by chemical etching or physical polishing, a less contaminated platinum wire may be provided for coating. In this manner, the surface area of ​​platinum available for hydrogen peroxide interaction is significantly increased, thereby increasing the overall electrical sensitivity of the sensor.

[0121] Then, as shown in block 807, the interfering compound is applied to the conductive substrate. In one particular application, the interfering compound is electrodeposited on the conductive substrate, which deposits the compound in a thin and uniform layer. In addition, the electrodeposition process promotes chemical crosslinking of the polymer as the PoAP is deposited. It should be understood that other processes can be used to apply the polymer to the conductive substrate.

[0122] As described above, the interfering membrane has a compound of self-limiting thickness. The overall allowable thickness of the membrane can be adjusted according to the ratio between the monomer and the buffer and the specific electrical properties used for the electropolymerization process. Moreover, the interfering membrane with specific selective permeability characteristics can be formulated by adjusting the pH. It should also be understood that the cyclic voltammetry (CV) process can be used to electrodeposit interfering membrane compounds, such as PoAP. The CV process is usually defined by the following characteristics: (1) a scanning window with a lower voltage limit and an upper voltage limit, (2) a starting point and direction within the scanning window, (3) the elapsed time of each cycle, and (4) the number of cycles completed. Those skilled in the art will understand that these four factors can provide almost unlimited alternatives in the precise application of interfering membrane compounds. In one example, the following range was found to be effective for the CV process applying PoAP:

[0123] Scan window: -1.0V to 2.0V

[0124] Starting point: -0.5V to 0.5VV

[0125] Rate: xy cycles per minute

[0126] Cycles: 5-50

[0127] As shown in step 811, an enzyme layer comprising glucose oxidase is then applied, and then a glucose limiting layer is applied as shown in 818. As described above, this glucose limiting layer can be used to limit the number of glucose molecules allowed to enter the enzyme layer.

[0128] Finally, an insulator may be applied to the reference wire, as shown in block 821. In many cases, the reference wire will be a silver / silver oxide wire, and the insulator will be an ion confinement layer that does not conduct electrons.

[0129] Glucose limiting layer

[0130] Reference now Fig.9A , a sensor 900 for use in a continuous biomonitor is shown. The sensor 900 has a working electrode 903 and a reference electrode 905. The silver / silver chloride reference electrode 914, the conductive layer 910, the interference membrane 909, and the enzyme layer 908 are similar to those discussed previously with reference to the sensor 700, and therefore will not be discussed in detail. It should be understood that there are several alternatives to these layers consistent with this disclosure.

[0131] Sensor 900 has glucose limiting membrane layer 907.As described, glucose limiting membrane 907 can use simple and cheap manufacturing technology to manufacture, and provide the more uniform glucose limiting membrane with more accurate regulation to glucose molecule.In this way, can more accurately and more uniformly limit and control the level of glucose molecule that allows to enter enzyme layer, and make gained calculation and result can be more linear and more accurate.Construct glucose limiting membrane 907 to provide the thin conformal layer of physical crosslinking material, described conformal layer is easy to arrange and provide excellent uniformity, glucose molecule control and linearity result.In a specific instance, physical crosslinking material uses hydrogen bond.Importantly, glucose limiting layer does not rely on chemical crosslinking.

[0132] Although glucose limiting layer 907 of the present invention is shown with sensor 900, it should be understood that glucose limiting layer 907 may also be advantageously used on other sensors, such as prior art sensor 600. It should be understood that glucose limiting layer of the present invention may be widely used on other types of biosensors.

[0133] like Fig.9A The glucose limiting layer 907 can be formulated to provide a uniform layer that more evenly and accurately delivers glucose molecules to the enzyme layer 908 than a typical prior art glucose limiting layer, which results in a more stable, consistent, and accurate generation of free electrons. Because the sensor 900 passes glucose and generates electrons more evenly and uniformly, the sensor is more accurate, less sensitive to noise, more stable, and easier to calibrate.

[0134] Reference now Fig. 9B , another sensor 901 for a continuous biomonitor is shown. Sensor 901 is similar to sensor 900 and therefore will not be described in detail. Sensor 901 has a working electrode 903 that is the same as the working electrode described with reference to sensor 900. However, sensor 901 has a reference lead 925 that has a silver / silver chloride layer 926 surrounded by an ion limiting membrane 928. The application of this ion limiting membrane 928 on the silver / silver chloride layer 926 ideally controls the current sensitivity of the overall sensor device 901 by controlling the ion flow from the silver / silver chloride layer. In this way, the current sensitivity can be advantageously controlled and defined.

[0135] Reference now Fig.10, generally describes the process 1000 for producing the glucose limiting layer.Before providing details and examples, generally describes the process.First, as shown in step 1002, select a hydrophilic bonding material.And, as shown in step 1004, select a hydrophobic bonding material, and as shown in step 1007, select a solvent.Hydrophilic bonding material, hydrophobic bonding material and solvent are mixed together with a desired ratio, which produces a bonding gel as shown in step 1011.Then this bonding gel can be applied to the enzyme layer on the working wire 1018.Then as shown in 1021, the gel is solidified to form a strong and flexible hydrogen bond structure.It should be understood that other materials can be used, and other types of physical crosslinks can be formed.

[0136] When selecting hydrophilic bonding material 1002, it is desirable to identify a hydrophilic bonding material with a relatively high molecular weight, such as 1 million to 5 million. It has been found that the hydrophilic bonding material with a molecular weight of 1 million to 3 million is particularly effective. As understood, the molecular weight of a polymer is the sum of the atomic weights of all atoms in a molecule. Therefore, the selected hydrophilic bonding material is typically a fairly large polymer. In addition, the hydrophilic body material is selected so that it is easy to distribute in a standard manufacturing process, and has the ability to form a strong hydrogen bond. Although in some embodiments, the hydrophilic bonding material has a relatively high molecular weight, is easy to distribute, and has strong hydrogen bonding ability, it should be understood that, according to specific application, other characteristics may become important. For example, polyvinyl alcohol, polyacrylic acid or polyvinyl pyrrolidone (PVP) can be used as the hydrophilic bonding material of the glucose limiting layer. In a specific example, the PVP of pharmaceutical grade form has a molecular weight of approximately 1.3 million. It should be understood that other polymers with similar or other desirable properties can be found.

[0137] Then select hydrophobic bonding material 1004. In particular, hydrophobic materials are selected based on ideal biocompatibility and the ratio between hard segments and soft segments. Typically, hydrophobic materials are formed by segments of small monomers cross-linked to much larger polymer parts. A higher ratio of soft segments allows hydrophobic bonding materials to have a higher degree of interaction with solvents and hydrophilic bonding materials; however, a higher ratio of soft segments also reduces the hydrophobic properties of the material because small segments tend to be hydrophilic. As for hard segments, a higher ratio of hard segments provides stronger physical properties, which are often measured as Shore hardness using a durometer. In this way, it is possible to select a hydrophobic material with an appropriate level of interaction with solvents and hydrophilic materials, and with sufficient hardness to be effectively used as a protective coating. In some embodiments, polyurethane can be used as a hydrophobic bonding material, which has the desired characteristics of providing sufficient hardness and the ideal interaction with hydrophilic bonding materials (e.g., PVP) and selected solvents. In addition, silicone can also be used as a hydrophobic bonding material. It should also be understood that other types of hydrophobic bonding materials can be selected according to the needs of a specific application.

[0138] The third material in step 1007 is a solvent. Usually, a polar, binary and volatile solvent is selected to meet the needs of solidification. First, the solvent should have sufficiently strong polarity characteristics to help correctly align hydrophilic bonding materials and hydrophobic bonding materials. Secondly, because the solvent must dissolve hydrophilic materials and hydrophobic materials simultaneously, a solvent with favorable solubility characteristics for each of the selected bonding materials should be selected. It should be understood that there are alternative ternary solvents. Finally, the volatility of the solvent supporting the required solidification characteristics should be selected. For example, some applications may need to be completed in a short period of time, so a fast flash solvent is required. In other cases, the less volatile solvent is alternative. In an example, a mixture of heavy organic compounds and alcohols can provide an ideal solvent for the glucose restriction layer. In a specific example, the heavy organic compound can be tetrahydrofuran (THF) or dimethylformamide (DMF), and the alcohol can be ethanol. It should be understood that other compounds that can provide ideal solvent characteristics can be used.

[0139] In step 1011, hydrophilic bonding material, hydrophobic bonding material and solvent are mixed together to form bonding gel. The viscosity of bonding gel can be adjusted by adjusting the ratio of solvent and bonding material. Then in step 1018, bonding gel can be applied to the enzyme layer. Bonding gel is easy to use, and various manufacturing processes can be used to immerse, spray, deposit or pad print. Then in step 1021, bonding gel is solidified, which can be completed in ambient air by using additional heating or by using additional vacuum. It should be understood that other processes can be used to speed up or slow down the curing process. Along with the bonding material solidification, hydrophobic and hydrophilic physical crosslinking, and particularly form hydrogen bond. Compared with the previous chemical bonding layer, the hydrogen bonding layer of gained enables glucose molecules to be highly desirable uniform and uniformly passed through.

[0140] Enzyme layer

[0141] As discussed with reference to sensor 600, sensor 700, and sensor 900, the working wire for each sensor has a corresponding enzyme layer 608, enzyme layer 708, and enzyme layer 908. As is well known, the enzyme layer promotes the chemical interaction between glucose and glucose oxidase (GOx), which produces hydrogen peroxide (H2O2). The hydrogen peroxide further reacts with the conductive platinum substrate, which produces a measurable free electron current, wherein the measured current level is proportional to the glucose level in the bloodstream or another body fluid such as ISF. In order to prepare a useful membrane for a glucose sensor, GOx is often stabilized with glutaraldehyde, imidates (dimethyl adipimidate, dimethylsuberimidate), hydroxysuccinimide, and its derivatives. Typically, about 0.6% glutaraldehyde preparations are mixed with GOx, and the mixture is then applied to the working wire. It should be understood that other ratios can be used and other additives can be present in the mixture. It is well known that polyethylenimine can also be used as a stabilizer for GOx, but it has similar disadvantages as glutaraldehyde.

[0142] Unfortunately, even with proper mixing, GOx cannot be evenly dispersed in glutaraldehyde, resulting in portions of the enzyme layer having higher GOx concentrations and portions having lower GOx concentrations. This uneven distribution of GOx causes an uneven interaction between glucose and GOx, which results in uneven production of hydrogen peroxide. That is, given a constant glucose level, different portions of the enzyme layer will produce more or less hydrogen peroxide molecules, thereby producing more or less free electrons. In this way, the measured glucose level can vary based on where the glucose molecules arrive and react on the enzyme layer. This uncertainty and variability can lead to erroneous blood glucose readings due to uneven GOx dispersion.

[0143] Furthermore, commercial GOx is derived from bacterial or fungal sources and is therefore known to be cytotoxic, that is, harmful to cells. Even when GOx is stabilized with glutaraldehyde, some GOx can still migrate within the layer and penetrate from the enzyme layer into the subject. Because none of the known protective layers for the enzyme layer completely entrap GOx, there is a risk that at least some GOx may be exposed to the subject's cells.

[0144] To address deficiencies in known glutaraldehyde-stabilized GOx, embodiments of enzyme layers are provided that provide substantially improved GOx retention and uniform distribution. Fig.11 As shown, the enzyme layer is prepared using process 1100. An aqueous polyurethane emulsion is prepared as shown in step 1103. It should be understood that the amount of water mixed with the polyurethane can be adjusted according to the requirements of the specific application. Fig.11 In the description of , polyurethane should be used, but it should be understood that other emulsions can be substituted, such as aqueous silicone dispersions. As shown in step 1105, the aqueous polyurethane emulsion is mixed with an aqueous acrylic polyol emulsion. The acrylic polyol acts as a self-crosslinking agent to produce a highly stable and compact structure with the polyurethane that can completely entrap GOx. The combination of the polyurethane emulsion from step 1103 and the acrylic polyol emulsion from step 1105 produces a base emulsion in step 1107. The ratio of polyurethane to acrylic polyol can be adjusted according to the requirements of a specific application; however, in one example, approximately equal amounts of each compound are mixed together in step 1107 to form a base emulsion. In some embodiments, GOx is blended with polyurethane at a ratio of about 1 part GOx to 60 parts polyurethane by volume. It should be understood that other ratios can be used depending on the specific application. It should also be understood that other enzymes can be used if other metabolic functions other than glucose levels are to be tested.

[0145] As shown in box 1110, other optional additives may be added. For example, one or more hydrophiles may be added to the emulsion mixture to promote better mixing or provide a more appropriate application viscosity. Examples of hydrophiles that can be used in the formulation of the enzyme layer include PVP, PEO, and Si-PEO. It should be understood that Si-PEO includes silane and PDMS PEO. It should be understood that other hydrophiles may be used. Although acrylic polyols can provide self-crosslinking when they are cured, other crosslinking polymers may be added for additional crosslinking. For GOx, such crosslinking agents may include, for example, glutaraldehyde, imidoesters, hydroxysuccinimide ( Hydroxysuccinimide ), carbodilite, melamine, epoxy resin and polyaziridine.

[0146] The polyurethane / GOx blend is applied to the working electrode in step 1109, for example by spraying, dipping, depositing or printing 1121. The blend is cured in step 1125, at which time the layers cross-link to provide a stable GOx dispersion.

[0147] Advantageously, the polyurethane / GOx blend is safe as an aqueous emulsion, easy to handle and apply, and provides a uniform distribution of GOx. In addition, because the cross-linked polymer is completely stabilized and trapped within the layer, GOx cannot move from the enzyme layer into the subject, thereby eliminating safety concerns. Moreover, because the polyurethane / GOx blend is more stable than the enzyme layer of the prior art, it has a longer usable shelf life and exhibits the ability to support higher loadings. In the case of higher GOx loadings, the polyurethane / GOx blend has higher sensitivity and achieves a higher signal-to-noise ratio.

[0148] sensor

[0149] Reference now Fig.12 , shows a sensor 1200 according to some embodiments. The sensor 1200 is constructed using a working electrode 1203 and a reference electrode 1205. As previously described, the reference electrode 1205 is typically silver chloride or silver 1214. The working electrode 1203 has a Fig.9A , Fig. 9B and Fig.10 The glucose limiting layer 1207. The working electrode 1203 also has Fig.11 The enzyme layer 1208 in contact with the glucose limiting layer 1207. The enzyme layer 1208 is also as shown in the reference Fig. 7A , Figure 7B and Figure 8 The interference film 1209 is in contact with the interference film 1209. The interference film 1209 is made of Figures 4A-4C and Figure 5 The substrate 1210 supports the fully described substrate 1210. As shown, the substrate 1210 has a plastic substrate portion 1210b, which has a plastic substrate portion 1210b as shown in reference Figures 4A-4C and Figure 5 The carbon coating 1210a is shown in the figure. It should be understood that, as shown, the carbon coating 1210a includes a hydrogen peroxide catalyst, such as phthalocyanine or Prussian blue.

[0150] Advantageously, sensor 700 provides an electrical signal with a higher signal-to-noise ratio, is cheaper to manufacture, and is safer for patients to wear than previous sensors.

[0151] Enzyme layer with direct electron or peroxide generation

[0152] Reference now Fig.13A, a sensor 1300 is shown. The sensor 1300 is shown as a two-wire sensor having a working electrode 1303 and a reference electrode 1305. As described with reference to previous sensors, the reference electrode 1305 is typically silver chloride or silver 1314. It should be understood that the construction of the sensor 1300 can be any of the common structures used for two-wire sensors.

[0153] The working wire 1303 has a glucose limiting layer 1307. The glucose limiting layer can have a known structure, but as shown in the figure, the glucose limiting layer 1307 is a reference Fig.9A and Fig. 9B As previously described, the glucose limiting layer is provided to limit and control the number of glucose molecules that can pass from the patient's blood or ISF to the enzyme layer, thereby improving the linearity of the overall sensor response. The glucose limiting layer still allows oxygen to enter the enzyme layer. Interference membrane 1309 can be positioned below glucose limiting layer 1307. In one example, interference membrane 1309 is a reference Fig. 7A and Figure 7B The interference membrane. Therefore, the interference layer 1309 has selective permeability to reject the passage of larger molecules. In this way, large molecules (such as acetaminophen) or other contaminants can be prevented from reaching the enzyme layer.

[0154] The sensor 1300 has a plastic material substrate 1312, such as a plastic wire. This plastic substrate material is selected to have sufficient strength for insertion under the patient's skin, as well as flexibility for patient comfort and ease of manufacture. In addition, it should be understood that the plastic substrate should be biologically safe and generally non-electrically reactive. It should be understood that a wide range of materials meet the mechanical and functional requirements of the selected plastic substrate. For example, many organic polymers and thermoplastics can be used. For illustrative purposes only, the following specific plastic substrate materials can be used: polyethylene, polypropylene, polystyrene, polyvinyl chloride, and polylactic acid. It should be understood that a wide variety of materials can be used as plastic substrates. This plastic substrate can be formed into elongated wires of many shapes to support the construction of different types of sensors. Typically, these plastic wires are formed using a well-known extrusion process.

[0155] Plastic substrate 1312 supports carbon-enzyme layer 1310. Typically, the novel carbon-enzyme layer 1310 is prepared as a coating of an aqueous dispersion of a carbon material, an elastomeric material, a cross-linking agent, and GOx. For example, the carbon material may be in the form of graphite, graphene, diamagnetic graphite, pyrolytic carbon, carbon black, carbon paste, or carbon ink. In some cases, other additives may be added to the carbon compound for enhancing electrical and reaction properties in order to support specific applications.

[0156] There are several acceptable elastomeric materials that will provide the desired properties, such as polyurethane, silicone, acrylate, or acrylic. It should be understood that other elastomeric materials may be substituted. Moreover, in some embodiments, the carbon compound does not delaminate from the plastic wire after the elastomeric material is cured, as compared to the platinum coating of the tantalum wire. In one example, an aqueous polyurethane emulsion is selected to be mixed with an aqueous acrylic polyol dispersion as a cross-linking agent. It should be understood that alternative or additional cross-linking agents and other additives may be used.

[0157] As described with reference to sensor 1300, GOx enzyme is used because sensor 1300 is related to detecting glucose levels. It should be understood that other enzymes such as lactate dehydrogenase (lactic acid), hydroxybutyrate (ketone) can be used for other metabolic sensors. It should also be understood that if enzymes other than GOx are used, the selection and ratio of materials in the carbon-enzyme layer may require additional modifications.

[0158] The carbon-enzyme coating 1310 is applied to a plastic substrate 1312 and allowed to cure. As the aqueous dispersion cures on the plastic substrate, it crosslinks to a flexible but strong coating of the substrate in which the GOx is uniformly dispersed and completely entrapped. In addition, the combination of selected carbon materials provides favorable structural, mechanical, and electrical properties of the carbon-enzyme layer.

[0159] In operation, the sensor 1300 allows glucose and oxygen to pass through the glucose limiting layer 1307 and the interference membrane 1309 into the carbon-enzyme layer 1310. The interference membrane 1309 blocks larger molecules that can contaminate or interfere with the chemical and electrical processes. Once the glucose and oxygen enter the carbon-enzyme layer 1310, they react to form hydrogen peroxide, which then interacts with the carbon to produce free electrons. These free electrons can then be conducted through the carbon-enzyme layer 1310 to the electronics of the sensor 1300 as shown by arrows 1315.

[0160] Advantageously, sensor 1300 does not use any platinum and has the GOx enzyme evenly dispersed and completely retained within the carbon-enzyme layer. In this way, sensor 1300 reduces the risk of potential safety issues using GOx and provides highly desirable sensitivity levels and high signal-to-noise performance.

[0161] refer to Fig. 13B , shows sensor 1301. Sensor 1301 is similar to sensor 1300 and therefore will not be described in detail. Sensor 1301 is a single-wire sensor in which a reference electrode 1305 is attached to a working electrode 1303. It should be understood that this physical configuration can be achieved through various printing processes, extrusion processes, and deposition processes.

[0162] Reference now Fig.14A, shows sensor 1400.Sensor 1400 is similar to sensor 1300, so will not be described in detail.As shown in the figure, sensor 1400 has a glucose limiting layer 1407 similar to glucose limiting layer 1307, a carbon-enzyme layer 1410 similar to carbon-enzyme layer 1310, a plastic substrate 1412 similar to plastic substrate 1312 and a reference electrode 1405 similar to reference electrode 1305, so these will not be described in detail.As shown in the figure, sensor 1400 does not have an interference membrane.In some cases, an interference membrane will not be needed, because one of the main purposes of the interference membrane is to prevent macromolecular contaminants from reaching and contaminating the platinum wire.Because sensor 1300 does not use platinum wire, the need for interference membrane can be reduced.

[0163] refer to Fig. 14B , sensor 1401 is shown. Sensor 1401 is similar to sensor 1400 and therefore will not be described in detail. Sensor 1401 is a single-wire sensor in which a reference electrode 1405 having silver chloride or silver 1414 is attached to a working electrode 1403. It should be understood that this physical configuration can be achieved through various printing processes and deposition processes.

[0164] Reference now Fig.15 , a single-wire sensor 1501 is shown. Sensor 1501 has a working wire 1505 physically attached to a reference wire 1507, wherein both working wire 1505 and reference wire 1507 have a semicircular cross-section with their planes facing each other. In some cases, an insulating member 1509 may be positioned at the planar interface between working wire 1505 and reference wire 1507. In some embodiments, working wire 1505 may be a reference wire. Fig. 13B or Fig. 14B The working conductor described. Fig.15 , another single-wire sensor 1511 is shown. The single-wire sensor 1511 has a reference wire 1515 attached to an insulating substrate 1513 having a triangular cross-section. In one example, the substrate 1513 can be an extruded plastic wire. A reference electrode 1517 and a counter electrode 1519 are also attached to the substrate 1513. In some embodiments, the working wire 1515 can be a reference Fig. 13B or Fig. 14B The working wire, reference wire 1515 , reference electrode 1517 and counter electrode 1519 have a plane facing substrate 1513 .

[0165] Reference now Fig.16 , describes a method for preparing and applying Fig.13A , Fig. 13B , Fig.14A and Fig. 14BFlow 1600 of the carbon-enzyme layer shown. As discussed with reference to sensor 600, sensor 700 and sensor 900, the working wire for each sensor has a corresponding enzyme layer 608, enzyme layer 708 and enzyme layer 908. As is well known, the enzyme layer promotes the chemical interaction between glucose and glucose oxidase (GOx), which produces hydrogen peroxide (H2O2). The hydrogen peroxide further reacts with the conductive platinum substrate, which produces a measurable free electron current, wherein the measured current level is proportional to the glucose level in the blood flow or in the ISF. In order to prepare a useful membrane for a glucose sensor, GOx is often stabilized with glutaraldehyde. Typically, about 0.6% glutaraldehyde preparation is mixed with GOx, and the mixture is then applied to the working wire. It should be understood that other ratios can be used and other additives may be present in the mixture. As is known, polyaziridine, imino ester, hydroxysuccinimide, carbodilite, melamine, epoxy resin, benzoyl peroxide, dicumyl peroxide can also be used as stabilizers for GOx, but they have similar disadvantages as glutaraldehyde.

[0166] Unfortunately, even with proper mixing, GOx cannot be evenly dispersed in glutaraldehyde, resulting in portions of the enzyme layer having higher GOx concentrations and portions having lower GOx concentrations. This uneven distribution of GOx causes an uneven interaction between glucose and GOx, which results in uneven production of hydrogen peroxide. That is, given a constant glucose level, different portions of the enzyme layer will produce more or less hydrogen peroxide molecules, thereby producing more or less free electrons. In this way, the measured glucose level can vary based on where the glucose molecules arrive and react on the enzyme layer. This uncertainty and variability can lead to erroneous blood glucose readings due to uneven GOx dispersion.

[0167] Furthermore, GOx is known to be cytotoxic, that is, harmful to cells. Even when GOx is stabilized with glutaraldehyde, some GOx can still migrate within the layer and penetrate from the enzyme layer into the subject. Because none of the known protective layers for the enzyme layer completely entrap GOx, there is a risk that at least some GOx may be exposed to the subject's cells.

[0168] Furthermore, most known sensors use either solid platinum or a substrate coated with platinum wire. Either way, platinum is expensive and results in a higher sensor price. Platinum also oxidizes easily, which can lead to instability and low signal-to-noise ratios. Figure 4A , Figure 4B , Figure 4C and Figure 5As discussed, a carbon compound can be applied to a plastic substrate, thereby eliminating the need for any platinum in the working lead. As described with reference to process 500, a carbon coating can be constructed to be flexible, strong, have desirable electrical properties, and applied in a manner that does not delaminate from the plastic substrate.

[0169] To address the deficiencies in known glutaraldehyde-stabilized GOx and the expense of platinum, a new carbon-enzyme layer is provided that provides substantially improved GOx capture and uniform distribution, wherein the carbon in the same layer provides strength, flexibility, and appropriate electrical properties. Furthermore, it has been found that the carbon-enzyme layer provides direct free electron generation, which not only eliminates the need for any ion-conducting layer or platinum wire, but also provides enhanced stability and substantially improved signal-to-noise ratio.

[0170] like Fig.16 As shown, the carbon-enzyme layer is prepared using process 1600. As shown in step 1603, an aqueous polyurethane emulsion is prepared. It should be understood that the amount of water mixed with the polyurethane can be adjusted according to the requirements of a specific application. Although polyurethane has been shown to perform well, it should be understood that other emulsions can be substituted, such as aqueous silicone dispersions. As shown in step 1605, the aqueous polyurethane emulsion is mixed with an aqueous acrylic polyol emulsion. Acrylic polyols act as self-crosslinking agents to produce highly stable and compact structures that can completely entrap GOx with polyurethane. The combination of the polyurethane emulsion in step 1603 and the acrylic polyol emulsion in step 1605 produces a base emulsion in step 1607. The ratio of polyurethane to acrylic polyol can be adjusted according to the requirements of a specific application; however, in one example, approximately equal amounts of each compound are mixed together in step 1607 to form a base emulsion. In one embodiment, GOx is blended with polyurethane at a ratio of about 1 part GOx to 60 parts polyurethane by volume. It should be understood that other ratios can be used according to specific applications. It should also be understood that other enzymes may be used in step 1609, such as ketones, lactate, or other metabolic catalysts, if other metabolic functions besides glucose levels are to be tested.

[0171] As shown in box 1610, other optional additives may be added. For example, one or more hydrophiles may be added to the emulsion mixture to promote better mixing or provide a more appropriate application viscosity. Well-known hydrophiles include PVP, PEO, and Si-PEO. It should be understood that Si-PEO includes silane and PDMS PEO. It should be understood that other hydrophiles may be used. Although acrylic polyols may provide self-crosslinking when they are cured, other crosslinking polymers may be added for additional crosslinking. For GOx, such crosslinking agents may include, for example, glutaraldehyde and polyethylenimine.

[0172] In step 1612, a blend of carbon materials is also mixed into the emulsion. The carbon materials and ratios are selected based on the application and functional requirements. For example, carbon (graphite) can be added to increase the strength of the resulting layer, while graphene, pyrolytic graphite, or a blend of graphene and pyrolytic graphite can be added to provide improved electrical properties. The ratio of graphite to be added will be selected to provide sufficient strength for the resulting carbon-enzyme layer, but still allow the layer to be flexible enough to avoid delamination from the plastic substrate or being too brittle to break. Moreover, the amount of graphene and pyrolytic graphite can be adjusted to set the ideal resistance for the carbon-enzyme layer. It should be understood that other forms of carbon can be substituted.

[0173] When the final mixture is prepared in step 1612, additional water or one or more hydrophiles (such as PVP) may be added in step 1616 to obtain appropriate viscosity and fluid properties (e.g., to thin the mixture) to promote uniform dispersion of GOx and enable the selected application technique.

[0174] Although a specific order of adding the components of the polyurethane / GOx / C blend in step 1612 has been shown, it should be understood that the order can be changed without affecting the resulting layers. In step 1621, the polyurethane / GOx / C blend is applied to the working electrode, such as by spraying, dipping, deposition, or printing. The blend is cured in step 1625, at which time the layers are cross-linked to provide a stable GOx dispersion.

[0175] Advantageously, the polyurethane / GOx / C blend is safe as an aqueous emulsion, easy to handle and apply, and provides a uniform distribution of GOx. In addition, because the cross-linked polymer is completely stabilized and trapped within the layer, GOx cannot migrate from the enzyme layer into the subject, thereby reducing safety concerns. Moreover, because the polyurethane / GOx / C blend is more stable than previous enzyme layers, it has a longer usable shelf life and exhibits the ability to support higher loadings. At higher GOx loadings, the polyurethane / GOx / C blend has higher sensitivity and achieves higher signal-to-noise ratios than known devices.

[0176] In one example of the carbon-enzyme emulsion applied in step 1621, the carbon-enzyme emulsion comprises the following:

[0177] 0.5 to 2 parts of polyurethane emulsion;

[0178] 0.5 to 2 parts of acrylic polyol emulsion;

[0179] 0.5 to 2 parts of carbon, comprising:

[0180] 0.5 to 1 part of graphite;

[0181] 0.0 to 1 part of graphene; and

[0182] 0.0 to 2 parts of pyrolytic graphite;

[0183] 0.0 to 3 parts of water and hydrophilic substances; and

[0184] 0.01 parts to 0.1 parts GOx.

[0185] Although the process 1600 has been discussed with reference to GOx enzymes, it should be understood that other enzymes may be substituted depending on the specific metabolic function to be monitored. For example, the following enzymes may be used in the process 1600. It should be understood that other enzymes and metabolic functions may be used, such as:

[0186] Enzyme metabolic function

[0187] Lactate dehydrogenase Lactate

[0188] HBD

[0189] Reference has been made in detail to the embodiments of the disclosed invention, one or more examples of which have been shown in the accompanying drawings. Each example has been provided by way of illustration of the present technology, rather than by way of limitation. In fact, although this specification has been described in detail with respect to specific embodiments of the present invention, it should be understood that those skilled in the art, after understanding the foregoing, can easily think of substitutions, variations and equivalents to these embodiments. For example, a feature shown or described as part of one embodiment can be used together with another embodiment to produce yet another embodiment. Therefore, it is intended that this subject matter covers all such modifications and variations within the scope of the attached claims and their equivalents. Without departing from the scope of the present invention more specifically set forth in the attached claims, a person of ordinary skill in the art may practice these and other modifications and variations of the present invention. In addition, a person of ordinary skill in the art should understand that the foregoing description is by way of example only, and is not intended to limit the present invention.

Claims

1. A method for manufacturing an enzyme membrane for a working electrode of a continuous biomonitor, the method comprising: preparing a waterborne polyurethane emulsion; preparing an acrylic polyol emulsion; mixing a polyurethane emulsion and an acrylic polyol emulsion to prepare a base emulsion; applying an enzyme to the base emulsion to produce an enzyme / base emulsion dispersion, the enzyme being selected based on the biological function to be monitored; applying the enzyme / base emulsion dispersion to the working electrode; and The applied enzyme / base emulsion dispersion is allowed to solidify.

2. The method of claim 1, wherein the enzyme is glucose oxidase and the biological function is glucose level.

3. The method of claim 1 further comprising adding a hydrophile to the enzyme / base emulsion dispersion prior to applying.

4. The method of claim 3, wherein the hydrophile is polyvinylpyrrolidone (PVP), polyethylene oxide (PEO) or silane-PEO (Si-PEO).

5. The method of claim 1 further comprising adding a cross-linking agent to the enzyme / base emulsion dispersion prior to applying.

6. A method according to claim 5, wherein the cross-linking agent is glutaraldehyde or polyethylenimine.

7. The method of claim 1 further comprising adding an imino ester, hydroxysuccinimide, carbodiimide, melamine, epoxy resin, benzoyl peroxide, or dicumyl peroxide to the enzyme / base emulsion dispersion prior to applying.

8. The method of claim 1, wherein the applying comprises dipping, spraying, depositing, printing, or pad printing.

9. The method of claim 1, wherein the curing comprises moving air, applying heat, or applying a vacuum.

10. The method of claim 1, wherein the ratio of the polyurethane emulsion to the acrylic polyol emulsion is about 1 to 1 by volume.

11. The method according to claim 1, wherein the enzyme is lactate dehydrogenase, lactate oxidase or hydroxybutyrate dehydrogenase.

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