Patent Grant
10631768
Electronic and electromechanical systems that are implanted in the body of animals, such as sensors and their associated electronic circuits that function to amplify the sensor signals and transmit them to a nearby receiver, require a power source. Currently, these systems are powered externally by batteries. The smallest batteries are, however, much larger than the implantable sensors and their associated signal amplifier circuits. For this reason, the size of autonomous packages that include a sensor, an amplifier-transmitter, and a power source are generally defined by the battery. Batteries cannot be made as small as the sensors or amplifiers because the batteries require cases and seals, the miniaturization of which is difficult and prohibitively expensive.
Known fuel cells are also much larger than available sensors because they require a case and a seal, and usually also a membrane, which is difficult to miniaturize and seal. Biological fuel cells, also known as biofuel cells, have been described in the past fifty years. However, only a few of these biofuel cells could be operated under physiological conditions. Physiological conditions include, for instance, a pH of about 7.2 to 7.4, a temperature of near 37° C., and a chloride concentration of about 0.14 M. Furthermore, known biofuel cells having higher power densities require ion-conducting separation membranes.
Accordingly, there still exists a need for the development of a system having millimeter to sub-millimeter dimensions that can function under physiological conditions to provide power to a sensor.
Generally, embodiments of the invention relate to self-powered analyte determining devices (e.g., electrochemical analyte monitoring systems) that include a working electrode, a counter electrode, and an optional resistance value, where the working electrode includes analyte sensing components and the self-powered analyte determining device spontaneously passes a current directly proportional to analyte concentration in the absence of an external power source. Also provided are systems and methods of using the, for example electrochemical, analyte sensors in analyte monitoring.
These and other objects, advantages, and features of the invention will become apparent to those persons skilled in the art upon reading the details of the invention as more fully described below.
The invention is best understood from the following detailed description when read in conjunction with the accompanying drawings. It is emphasized that, according to common practice, the various features of the drawings are not to-scale. On the contrary, the dimensions of the various features are arbitrarily expanded or reduced for clarity. Included in the drawings are the following figures:
Before the present invention is described, it is to be understood that this invention is not limited to particular embodiments described, as such may, of course, vary. It is also to be understood that the terminology used herein is for the purpose of describing particular embodiments only, and is not intended to be limiting, since the scope of the present invention will be limited only by the appended claims.
Where a range of values is provided, it is understood that each intervening value, to the tenth of the unit of the lower limit unless the context clearly dictates otherwise, between the upper and lower limits of that range is also specifically disclosed. Each smaller range between any stated value or intervening value in a stated range and any other stated or intervening value in that stated range is encompassed within the invention. The upper and lower limits of these smaller ranges may independently be included or excluded in the range, and each range where either, neither or both limits are included in the smaller ranges is also encompassed within the invention, subject to any specifically excluded limit in the stated range. Where the stated range includes one or both of the limits, ranges excluding either or both of those included limits are also included in the invention.
Unless defined otherwise, all technical and scientific terms used herein have the same meaning as commonly understood by one of ordinary skill in the art to which this invention belongs. Although any methods and materials similar or equivalent to those described herein can be used in the practice or testing of the present invention, some potential and preferred methods and materials are now described. All publications mentioned herein are incorporated herein by reference to disclose and describe the methods and/or materials in connection with which the publications are cited. It is understood that the present disclosure supercedes any disclosure of an incorporated publication to the extent there is a contradiction.
It must be noted that as used herein and in the appended claims, the singular forms “a”, “an”, and “the” include plural referents unless the context clearly dictates otherwise.
The publications discussed herein are provided solely for their disclosure prior to the filing date of the present application. Nothing herein is to be construed as an admission that the present invention is not entitled to antedate such publication by virtue of prior invention. Further, the dates of publication provided may be different from the actual publication dates which may need to be independently confirmed.
Generally, embodiments of the invention relate to self-powered analyte determining devices (e.g., electrochemical analyte monitoring systems) that include a working electrode, a counter electrode, and an optional resistance value, where the working electrode includes analyte sensing components and the self-powered analyte determining device spontaneously passes a current directly proportional to analyte concentration in the absence of an external power source. Also provided are systems and methods of using the, for example electrochemical, analyte sensors in analyte monitoring.
The self-powered analyte sensor uses compounds available in a variety of biological systems to provide a low level of power and also detect a level of an analyte. In general, the self-powered analyte sensor described herein spontaneously passes a current proportional to an analyte concentration, such as glucose, in the absence of an external power supply, such as a battery. As such, a self-powered analyte sensor that is partially implanted under the skin of a subject or fully implanted under the skin of a subject can be coupled to a radio-frequency powered measurement circuit to remotely provide analyte levels.
Non-self-powered continuous analyte sensors have fairly long equilibration times, and therefore need to be continuously powered after inserting a portion thereof under the skin of a subject. That is, if the power supply of a continuous analyte sensor is removed and it is turned “off”, then a fairly long wait time of approximately 10 to 15 minutes must pass before a reliable analyte level reading is detected following reconnection of the power supply. In contrast, a self-powered analyte sensor, as described herein, has the advantage of providing a low level of continuous power in the absence of an external power supply while also detecting an analyte level thereby requiring little to no equilibration time between analyte measurement intervals.
With reference to the schematic illustration of
In one embodiment of the invention, an enzyme used for the electrooxidation of glucose can be GOx disposed on the working electrode with redox polymer and cross-linker, and the electrocatalyst for the electroreduction of oxygen can be redox polymer, redox polymer with enzyme, or platinum-carbon cloth disposed on the counter electrode. In the operation of the cell, via its wiring, electrons are collected from glucose-reduced GOx at the working electrode, where glucose is electrooxidized to δ-gluconolactone (see Eq. 1 below), and electrons are delivered to the reox polymer, redox polymer with enzyme, or platinum-carbon cloth at the counter electrode, where oxygen is electroreduced to water (see Eq. 2 below). The overall reaction is that of Eq. 3 below.
D-glucose→δ-gluconolactone+2H++2e− (Eq. 1)
O2+4H++4e−→2H2O (Eq. 2)
2D-glucose+O2→2δ-gluconolactone+2H2O (Eq. 3)
The reactions represented by Eqs. 1-3 above are now described in relation to the schematic illustration of
A working electrode of the self-powered analyte sensor effectuates the electrooxidation of the reductant, such as the electrooxidation of glucose to gluconolactone as schematically illustrated in
A counter electrode of the self-powered analyte sensor effectuates the four-electron electroreduction of O2 to water under physiological conditions as schematically illustrated in
The physical dimensions of the self-powered analyte sensor, as well as its operational parameters (e.g., output power and voltage) are, at least in part, a function of the components that form the sensor. Merely by way of example, the voltage drop across the resistance, R, of the self-powered analyte sensor can range from about 0 volts to about 1.2 volts, and typically, from about 0 volts to about 0.3 volts.
The output current of an implanted self-powered analyte sensor, can be limited by electrode kinetics, ohmic resistance, and/or mass transport. When the electrode kinetics is fast and the self-powered analyte sensor is small enough for the ohmic resistance of the electrolytic solution to be low, the output current is determined by the mass transport of the reactants to the electrodes. In particular, output current is generally determined by mass transport of a single limiting reactant, for example glucose. The mass-transport-limited current density increases upon raising the concentration of the reactants, particularly the limiting reactant, or upon increasing the permeability of any overlying membrane.
The self-powered analyte sensor further includes an optional resistance disposed between the working electrode and the counter electrode and having an R value that is chosen from within a range of R values for which an electrochemical analyte sensor current output at an analyte concentration at the top of the physically relevant range of the analyte is independent of R. Such an R value can be calculated, as exemplified in
In some embodiments, the R value is chosen such that the potential drop across the resistor, at an analyte concentration at the top of its physically relevant range, is less than 75% of the open circuit potential. In other embodiments, the R value is chosen such that the potential drop across the resistor, at an analyte concentration at the top of its physically relevant range, is less than 50% of the open circuit potential. In some embodiments, the R value is chosen such that the potential drop across the resistor, at an analyte concentration at the top of its physically relevant range, is less than 25% of the open circuit potential. In certain embodiments, the R value is less than about 40 MΩ, including less than about 20 MΩ, less than about 10 MΩ, and less than about 1 MΩ.
Thus, in accordance with an embodiment of the invention, the self-powered analyte sensor is usually assembled without the case, the seal, and the compartment-separating membrane required by conventional batteries. As a result, the self-powered analyte sensor can be manufactured with a footprint smaller than about 3 mm2, including smaller than about 1 mm2, such as smaller than about 0.3 mm2. The self-powered analyte sensor, with dimensions on the millimeter to sub-millimeter level, can power sensor components, as further described herein. Further, the self-powered analyte sensor can be constructed so that none of its components, including its catalytic components, will be dissolved or leached while residing in the body.
Electrochemical Sensors
Embodiments of the invention relate to methods and devices for detecting at least one analyte, including glucose, in body fluid. Embodiments relate to the continuous and/or automatic in vivo monitoring of the level of one or more analytes using a continuous analyte monitoring system that includes an analyte sensor at least a portion of which is to be positioned beneath a skin surface of a user for a period of time and/or the discrete monitoring of one or more analytes using an in vitro blood glucose (“BG”) meter and an analyte test strip. Embodiments include combined or combinable devices, systems and methods and/or transferring data between an in vivo continuous system and an in vivo system. In some embodiments, the systems, or at least a portion of the systems, are integrated into a single unit.
A self-powered analyte sensor may be an in vivo sensor or an in vitro sensor (i.e., a discrete monitoring test strip). Such a sensor can be formed on a substrate, e.g., a substantially planar substrate. In certain embodiments, such a sensor is a wire, e.g., a working electrode wire inner portion with one or more other electrodes associated (e.g., on, including wrapped around) therewith. The sensor also includes at least one counter electrode.
Accordingly, embodiments include self-powered analyte monitoring devices and systems that include an analyte sensor at least a portion of which is positionable beneath the skin of the user for the in vivo detection, of an analyte, including glucose, lactate, and the like, in a body fluid. Embodiments include wholly implantable self-powered analyte sensors and analyte sensors in which only a portion of the sensor is positioned under the skin and a portion of the sensor resides above the skin, e.g., for contact to a sensor control unit (which may include a transmitter), a receiver/display unit, transceiver, processor, etc. The sensor may be, for example, subcutaneously positionable in a patient for the continuous or periodic monitoring of a level of an analyte in a patient's interstitial fluid. For the purposes of this description, continuous monitoring and periodic monitoring will be used interchangeably, unless noted otherwise. The sensor response may be correlated and/or converted to analyte levels in blood or other fluids. In certain embodiments, an analyte sensor may be positioned in contact with interstitial fluid to detect the level of glucose, which detected glucose may be used to infer the glucose level in the patient's bloodstream. Self-powered analyte sensors may be insertable into a vein, artery, or other portion of the body containing fluid. Embodiments of the analyte sensors of the subject invention having a plasticizer may be configured for monitoring the level of the analyte over a time period which may range from seconds, minutes, hours, days, weeks, to months, or longer.
Of interest are self-powered analyte sensors, such as glucose sensors, that are capable of in vivo detection of an analyte for about one hour or more, e.g., about a few hours or more, e.g., about a few days or more, e.g., about three or more days, e.g., about five days or more, e.g., about seven days or more, e.g., about several weeks or at least one month or more. Future analyte levels may be predicted based on information obtained, e.g., the current analyte level at time t0, the rate of change of the analyte, etc. Predictive alarms may notify the user of a predicted analyte levels that may be of concern in advance of the user's analyte level reaching the future level. This provides the user an opportunity to take corrective action.
Analytes that may be monitored include, but are not limited to, acetyl choline, amylase, bilirubin, cholesterol, chorionic gonadotropin, creatine kinase (e.g., CK-MB), creatine, creatinine, DNA, fructosamine, glucose, glutamine, growth hormones, hormones, ketone bodies, lactate, peroxide, prostate-specific antigen, prothrombin, RNA, thyroid stimulating hormone, and troponin. The concentration of drugs, such as, for example, antibiotics (e.g., gentamicin, vancomycin, and the like), digitoxin, digoxin, drugs of abuse, theophylline, and warfarin, may also be monitored. In those embodiments that monitor more than one analyte, the analytes may be monitored at the same or different times.
The analyte monitoring system 400 includes a sensor 401, a data processing unit 402 connectable to the sensor 401, and a primary receiver unit 404 which is configured to communicate with the data processing unit 402 via a communication link 403. In certain embodiments, the primary receiver unit 104 may be further configured to transmit data to a data processing terminal 405 to evaluate or otherwise process or format data received by the primary receiver unit 404. The data processing terminal 405 may be configured to receive data directly from the data processing unit 402 via a communication link which may optionally be configured for bi-directional communication. Further, the data processing unit 402 may include a transmitter or a transceiver to transmit and/or receive data to and/or from the primary receiver unit 404 and/or the data processing terminal 405 and/or optionally the secondary receiver unit 406.
Also shown in
Only one self-powered sensor 401, data processing unit 402 and data processing terminal 405 are shown in the embodiment of the analyte monitoring system 400 illustrated in
The analyte monitoring system 400 may be a continuous monitoring system, or semi-continuous, or a discrete monitoring system. In a multi-component environment, each component may be configured to be uniquely identified by one or more of the other components in the system so that communication conflict may be readily resolved between the various components within the analyte monitoring system 400. For example, unique IDs, communication channels, and the like, may be used.
In certain embodiments, the sensor 401 is physically positioned in or on the body of a user whose analyte level is being monitored. The sensor 401 may be configured to at least periodically sample the analyte level of the user and convert the sampled analyte level into a corresponding signal for transmission by the data processing unit 402. The data processing unit 402 is coupleable to the self-powered sensor 401 so that both devices are positioned in or on the user's body, with at least a portion of the self-powered analyte sensor 401 positioned transcutaneously. The data processing unit may include a fixation element such as adhesive or the like to secure it to the user's body. An optional mount (not shown) attachable to the user and mateable with the unit 402 may be used. For example, a mount may include an adhesive surface. The data processing unit 402 performs data processing functions, where such functions may include but are not limited to, filtering and encoding of data signals, each of which corresponds to a sampled analyte level of the user, for transmission to the primary receiver unit 404 via the communication link 403. In one embodiment, the self-powered sensor 401 or the data processing unit 402 or a combined self-powered sensor/data processing unit may be wholly implantable under the skin layer of the user.
In certain embodiments, the primary receiver unit 404 may include an analog interface section including and RF receiver and an antenna that is configured to communicate with the data processing unit 402 via the communication link 403, and a data processing section for processing the received data from the data processing unit 402 including data decoding, error detection and correction, data clock generation, data bit recovery, etc., or any combination thereof.
In operation, the primary receiver unit 404 in certain embodiments is configured to synchronize with the data processing unit 402 to uniquely identify the data processing unit 402, based on, for example, an identification information of the data processing unit 402, and thereafter, to periodically receive signals transmitted from the data processing unit 402 associated with the monitored analyte levels detected by the sensor 401.
Referring again to
The data processing terminal 405 may include an infusion device such as an insulin infusion pump or the like, which may be configured to administer insulin to patients, and which may be configured to communicate with the primary receiver unit 404 for receiving, among others, the measured analyte level. Alternatively, the primary receiver unit 404 may be configured to integrate an infusion device therein so that the primary receiver unit 404 is configured to administer insulin (or other appropriate drug) therapy to patients, for example, for administering and modifying basal profiles, as well as for determining appropriate boluses for administration based on, among others, the detected analyte levels received from the data processing unit 402. An infusion device may be an external device or an internal device (wholly implantable in a user).
In certain embodiments, the data processing terminal 405, which may include an insulin pump, may be configured to receive the analyte signals from the data processing unit 402, and thus, incorporate the functions of the primary receiver unit 404 including data processing for managing the patient's insulin therapy and analyte monitoring. In certain embodiments, the communication link 403 as well as one or more of the other communication interfaces shown in
As can be seen in the embodiment of
In certain embodiments, the test strip interface 601 includes a glucose level testing portion to receive a blood (or other body fluid sample) glucose test or information related thereto. For example, the interface may include a test strip port to receive a glucose test strip. The device may determine the glucose level of the test strip, and optionally display (or otherwise notice) the glucose level on the output 610 of the primary receiver unit 404. Any suitable test strip may be employed, e.g., test strips that only require a very small amount (e.g., one microliter or less, e.g., 0.5 microliter or less, e.g., 0.1 microliter or less), of applied sample to the strip in order to obtain accurate glucose information, e.g. FreeStyle® blood glucose test strips from Abbott Diabetes Care, Inc. Glucose information obtained by the in vitro glucose testing device may be used for a variety of purposes, computations, etc. For example, the information may be used to calibrate sensor 401, confirm results of the sensor 401 to increase the confidence thereof (e.g., in instances in which information obtained by sensor 401 is employed in therapy related decisions), etc.
In further embodiments, the data processing unit 402 and/or the primary receiver unit 404 and/or the secondary receiver unit 405, and/or the data processing terminal/infusion section 405 may be configured to receive the blood glucose value wirelessly over a communication link from, for example, a blood glucose meter. In further embodiments, a user manipulating or using the analyte monitoring system 400 (
Additional detailed descriptions are provided in U.S. Pat. Nos. 5,262,035; 5,264,104; 5,262,305; 5,320,715; 5,593,852; 6,175,752; 6,650,471; 6,746,582, and in application Ser. No. 10/745,878 filed Dec. 26, 2003 entitled “Continuous Glucose Monitoring System and Methods of Use”, each of which is incorporated herein by reference.
The sensor may be wholly implantable in a user or may be configured so that only a portion is positioned within (internal) a user and another portion outside (external) a user. For example, the sensor 700 may include a portion positionable above a surface of the skin 710, and a portion positioned below the skin. In such embodiments, the external portion may include contacts (connected to respective electrodes of the second portion by traces) to connect to another device also external to the user such as a transmitter unit. While the embodiment of
A first insulation layer such as a first dielectric layer 805 is disposed or layered on at least a portion of the first conducting layer 801. A second conducting layer 803 may provide the counter electrode 803. It may be disposed on at least a portion of the first insulation layer 805. Finally, a second insulation layer may be disposed or layered on at least a portion of the second conducting layer 803. In this manner, the sensor 800 may be layered such that at least a portion of each of the conducting layers is separated by a respective insulation layer (for example, a dielectric layer). The embodiment of
In certain embodiments, some or all of the electrodes 801, 803 may be provided on the same side of the substrate 804 in the layered construction as described above, or alternatively, may be provided in a co-planar manner such that two or more electrodes may be positioned on the same plane (e.g., side-by side (e.g., parallel) or angled relative to each other) on the substrate 804. For example, co-planar electrodes may include a suitable spacing there between and/or include dielectric material or insulation material disposed between the conducting layers/electrodes. Furthermore, in certain embodiments one or more of the electrodes 801, 803 may be disposed on opposing sides of the substrate 804. In such embodiments, contact pads may be one the same or different sides of the substrate. For example, an electrode may be on a first side and its respective contact may be on a second side, e.g., a trace connecting the electrode and the contact may traverse through the substrate.
As noted above, analyte sensors may include an analyte-responsive enzyme to provide a sensing component or sensing layer. Some analytes, such as oxygen, can be directly electrooxidized or electroreduced on a sensor, and more specifically at least on a working electrode of a sensor. Other analytes, such as glucose and lactate, require the presence of at least one electron transfer agent and/or at least one catalyst to facilitate the electrooxidation or electroreduction of the analyte. Catalysts may also be used for those analytes, such as oxygen, that can be directly electrooxidized or electroreduced on the working electrode. For these analytes, each working electrode includes a sensing layer (see for example sensing layer 808 of
The sensing layer includes one or more components constructed to facilitate the electrochemical oxidation or reduction of the analyte. The sensing layer may include, for example, a catalyst to catalyze a reaction of the analyte and produce a response at the working electrode, an electron transfer agent to transfer electrons between the analyte and the working electrode (or other component), or both. The sensing layer and the working electrode also function as the anode of the power generating component of the self-powered analyte sensor, thereby providing the dual-function of power generation and analyte level detection.
A variety of different sensing layer configurations may be used. In certain embodiments, the sensing layer is deposited on the conductive material of a working electrode. The sensing layer may extend beyond the conductive material of the working electrode. In some cases, the sensing layer may also extend over other electrodes.
A sensing layer that is in direct contact with the working electrode may contain an electron transfer agent to transfer electrons directly or indirectly between the analyte and the working electrode, and/or a catalyst to facilitate a reaction of the analyte. For example, a glucose, lactate, or oxygen electrode may be formed having a sensing layer which contains a catalyst, including glucose oxidase, glucose dehydrogenase, lactate oxidase, or laccase, respectively, and an electron transfer agent that facilitates the electrooxidation of the glucose, lactate, or oxygen, respectively.
In other embodiments the sensing layer is not deposited directly on the working electrode. Instead, the sensing layer 808 may be spaced apart from the working electrode, and separated from the working electrode, e.g., by a separation layer. A separation layer may include one or more membranes or films or a physical distance. In addition to separating the working electrode from the sensing layer the separation layer may also act as a mass transport limiting layer and/or an interferent eliminating layer and/or a biocompatible layer.
In certain embodiments which include more than one working electrode, one or more of the working electrodes may not have a corresponding sensing layer, or may have a sensing layer which does not contain one or more components (e.g., an electron transfer agent and/or catalyst) needed to electrolyze the analyte. Thus, the signal at this working electrode may correspond to background signal which may be removed from the analyte signal obtained from one or more other working electrodes that are associated with fully-functional sensing layers by, for example, subtracting the signal.
In certain embodiments, the sensing layer includes one or more electron transfer agents. Electron transfer agents that may be employed are electroreducible and electrooxidizable ions or molecules having redox potentials that are a few hundred millivolts above or below the redox potential of the standard calomel electrode (SCE). The electron transfer agent may be organic, organometallic, or inorganic. Examples of organic redox species are quinones and species that in their oxidized state have quinoid structures, such as Nile blue and indophenol. Examples of organometallic redox species are metallocenes including ferrocene. Examples of inorganic redox species are hexacyanoferrate (III), ruthenium hexamine etc. Additional examples include those described in U.S. Pat. No. 6,736,957 and U.S. Patent Publication Nos. 2004/0079653 and 2006/0201805, the disclosures of which are incorporated herein by reference in their entirety.
In certain embodiments, electron transfer agents have structures or charges which prevent or substantially reduce the diffusional loss of the electron transfer agent during the period of time that the sample is being analyzed. For example, electron transfer agents include but are not limited to a redox species, e.g., bound to a polymer which can in turn be disposed on or near the working electrode. The bond between the redox species and the polymer may be covalent, coordinative, or ionic. Although any organic, organometallic or inorganic redox species may be bound to a polymer and used as an electron transfer agent, in certain embodiments the redox species is a transition metal compound or complex, e.g., osmium, ruthenium, iron, and cobalt compounds or complexes. It will be recognized that many redox species described for use with a polymeric component may also be used, without a polymeric component.
One type of polymeric electron transfer agent contains a redox species covalently bound in a polymeric composition. An example of this type of mediator is poly(vinylferrocene). Another type of electron transfer agent contains an ionically-bound redox species. This type of mediator may include a charged polymer coupled to an oppositely charged redox species. Examples of this type of mediator include a negatively charged polymer coupled to a positively charged redox species such as an osmium or ruthenium polypyridyl cation. Another example of an ionically-bound mediator is a positively charged polymer including quaternized poly(4-vinyl pyridine) or poly(1-vinyl imidazole) coupled to a negatively charged redox species such as ferricyanide or ferrocyanide. In other embodiments, electron transfer agents include a redox species coordinatively bound to a polymer. For example, the mediator may be formed by coordination of an osmium or cobalt 2,2′-bipyridyl complex to poly(1-vinyl imidazole) or poly(4-vinyl pyridine).
Suitable electron transfer agents are osmium transition metal complexes with one or more ligands, each ligand having a nitrogen-containing heterocycle such as 2,2′-bipyridine, 1,10-phenanthroline, 1-methyl, 2-pyridyl biimidazole, or derivatives thereof. The electron transfer agents may also have one or more ligands covalently bound in a polymer, each ligand having at least one nitrogen-containing heterocycle, such as pyridine, imidazole, or derivatives thereof. One example of an electron transfer agent includes (a) a polymer or copolymer having pyridine or imidazole functional groups and (b) osmium cations complexed with two ligands, each ligand containing 2,2′-bipyridine, 1,10-phenanthroline, or derivatives thereof, the two ligands not necessarily being the same. Some derivatives of 2,2′-bipyridine for complexation with the osmium cation include but are not limited to 4,4′-dimethyl-2,2′-bipyridine and mono-, di-, and polyalkoxy-2,2′-bipyridines, including 4,4′-dimethoxy-2,2′-bipyridine. Derivatives of 1,10-phenanthroline for complexation with the osmium cation include but are not limited to 4,7-dimethyl-1,10-phenanthroline and mono, di-, and polyalkoxy-1,10-phenanthrolines, such as 4,7-dimethoxy-1,10-phenanthroline. Polymers for complexation with the osmium cation include but are not limited to polymers and copolymers of poly(1-vinyl imidazole) (referred to as “PVI”) and poly(4-vinyl pyridine) (referred to as “PVP”). Suitable copolymer substituents of poly(1-vinyl imidazole) include acrylonitrile, acrylamide, and substituted or quaternized N-vinyl imidazole, e.g., electron transfer agents with osmium complexed to a polymer or copolymer of poly(1-vinyl imidazole).
Embodiments may employ electron transfer agents having a redox potential ranging from about −200 mV to about +200 mV versus the standard calomel electrode (SCE). The sensing layer may also include a catalyst which is capable of catalyzing a reaction of the analyte. The catalyst may also, in some embodiments, act as an electron transfer agent. One example of a suitable catalyst is an enzyme which catalyzes a reaction of the analyte. For example, a catalyst, including a glucose oxidase, glucose dehydrogenase (e.g., pyrroloquinoline quinone (PQQ), dependent glucose dehydrogenase, flavine adenine dinucleotide (FAD) dependent glucose dehydrogenase, or nicotinamide adenine dinucleotide (NAD) dependent glucose dehydrogenase), may be used when the analyte of interest is glucose. A lactate oxidase or lactate dehydrogenase may be used when the analyte of interest is lactate. Laccase may be used when the analyte of interest is oxygen or when oxygen is generated or consumed in response to a reaction of the analyte.
The sensing layer may also include a catalyst which is capable of catalyzing a reaction of the analyte. The catalyst may also, in some embodiments, act as an electron transfer agent. One example of a suitable catalyst is an enzyme which catalyzes a reaction of the analyte. For example, a catalyst, including a glucose oxidase, glucose dehydrogenase (e.g., pyrroloquinoline quinone (PQQ), dependent glucose dehydrogenase or oligosaccharide dehydrogenase, flavine adenine dinucleotide (FAD) dependent glucose dehydrogenase, nicotinamide adenine dinucleotide (NAD) dependent glucose dehydrogenase), may be used when the analyte of interest is glucose. A lactate oxidase or lactate dehydrogenase may be used when the analyte of interest is lactate. Laccase may be used when the analyte of interest is oxygen or when oxygen is generated or consumed in response to a reaction of the analyte.
In certain embodiments, a catalyst may be attached to a polymer, cross linking the catalyst with another electron transfer agent, which, as described above, may be polymeric. A second catalyst may also be used in certain embodiments. This second catalyst may be used to catalyze a reaction of a product compound resulting from the catalyzed reaction of the analyte. The second catalyst may operate with an electron transfer agent to electrolyze the product compound to generate a signal at the working electrode. Alternatively, a second catalyst may be provided in an interferent-eliminating layer to catalyze reactions that remove interferents.
In certain embodiments, the sensing layer functions at a gentle oxidizing potential, e.g., a potential of about +40 mV vs. Ag/AgCl. This sensing layer uses, for example, an osmium (Os)-based mediator constructed for low potential operation and includes a plasticizer. Accordingly, in certain embodiments the sensing element is a redox active component that includes (1) Osmium-based mediator molecules that include (bidente) ligands, and (2) glucose oxidase enzyme molecules. These two constituents are combined together with a cross-linker.
A mass transport limiting layer (not shown), e.g., an analyte flux modulating layer, may be included with the sensor to act as a diffusion-limiting barrier to reduce the rate of mass transport of the analyte, for example, glucose or lactate, into the region around the working electrodes. The mass transport limiting layers are useful in limiting the flux of an analyte to a working electrode in an electrochemical sensor so that the sensor is linearly responsive over a large range of analyte concentrations and is easily calibrated. Mass transport limiting layers may include polymers and may be biocompatible. A mass transport limiting layer may provide many functions, e.g., biocompatibility and/or interferent-eliminating, etc.
In certain embodiments, a mass transport limiting layer is a membrane composed of crosslinked polymers containing heterocyclic nitrogen groups, such as polymers of polyvinylpyridine and polyvinylimidazole. In some embodiments, a plasticizer is combined with the mass transport limiting layer or membrane. Embodiments also include membranes that are made of a polyurethane, or polyether urethane, or chemically related material, or membranes that are made of silicone, and the like.
A membrane may be formed by crosslinking in situ a polymer, modified with a zwitterionic moiety, a non-pyridine copolymer component, and optionally another moiety that is either hydrophilic or hydrophobic, and/or has other desirable properties, in an alcohol-buffer solution. In certain embodiments, the membrane formulation further includes a plasticizer. The modified polymer may be made from a precursor polymer containing heterocyclic nitrogen groups. For example, a precursor polymer may be polyvinylpyridine or polyvinylimidazole. Optionally, hydrophilic or hydrophobic modifiers may be used to “fine-tune” the permeability of the resulting membrane to an analyte of interest. Optional hydrophilic modifiers, such as poly(ethylene glycol), hydroxyl or polyhydroxyl modifiers, may be used to enhance the biocompatibility of the polymer or the resulting membrane.
A membrane may be formed in situ by applying an alcohol-buffer solution of a crosslinker and a modified polymer over an enzyme-containing sensing layer and allowing the solution to cure for about one to two days or other appropriate time period. The crosslinker-polymer solution may be applied to the sensing layer by placing a droplet or droplets of the solution on the sensor, by dipping the sensor into the solution, or the like. Generally, the thickness of the membrane is controlled by the concentration of the solution, by the number of droplets of the solution applied, by the number of times the sensor is dipped in the solution, or by any combination of these factors. A membrane applied in this manner may have any combination of the following functions: (1) mass transport limitation, i.e., reduction of the flux of analyte that can reach the sensing layer, (2) biocompatibility enhancement, or (3) interferent reduction.
The substrate may be formed using a variety of non-conducting materials, including, for example, polymeric or plastic materials and ceramic materials. Suitable materials for a particular sensor may be determined, at least in part, based on the desired use of the sensor and properties of the materials.
In some embodiments, the substrate is flexible. For example, if the sensor is configured for implantation into a patient, then the sensor may be made flexible (although rigid sensors may also be used for implantable sensors) to reduce pain to the patient and damage to the tissue caused by the implantation of and/or the wearing of the sensor. A flexible substrate often increases the patient's comfort and allows a wider range of activities. Suitable materials for a flexible substrate include, for example, non-conducting plastic or polymeric materials and other non-conducting, flexible, deformable materials. Examples of useful plastic or polymeric materials include thermoplastics such as polycarbonates, polyesters (e.g., Mylar™ and polyethylene terephthalate (PET)), polyvinyl chloride (PVC), polyurethanes, polyethers, polyamides, polyimides, or copolymers of these thermoplastics, such as PETG (glycol-modified polyethylene terephthalate).
In other embodiments, the sensors are made using a relatively rigid substrate to, for example, provide structural support against bending or breaking. Examples of rigid materials that may be used as the substrate include poorly conducting ceramics, such as aluminum oxide and silicon dioxide. One advantage of an implantable sensor having a rigid substrate is that the sensor may have a sharp point and/or a sharp edge to aid in implantation of a sensor without an additional insertion device.
It will be appreciated that for many sensors and sensor applications, both rigid and flexible sensors will operate adequately. The flexibility of the sensor may also be controlled and varied along a continuum by changing, for example, the composition and/or thickness of the substrate.
In addition to considerations regarding flexibility, it is often desirable that implantable sensors should have a substrate which is physiologically harmless, for example, a substrate approved by a regulatory agency or private institution for in vivo use.
The sensor may include optional features to facilitate insertion of an implantable sensor. For example, the sensor may be pointed at the tip to ease insertion. In addition, the sensor may include a barb which assists in anchoring the sensor within the tissue of the patient during operation of the sensor. However, the barb is typically small enough so that little damage is caused to the subcutaneous tissue when the sensor is removed for replacement.
An implantable sensor may also, optionally, have an anticlotting agent disposed on a portion of the substrate which is implanted into a patient. This anticlotting agent may reduce or eliminate the clotting of blood or other body fluid around the sensor, particularly after insertion of the sensor. Blood clots may foul the sensor or irreproducibly reduce the amount of analyte which diffuses into the sensor. Examples of useful anticlotting agents include heparin and tissue plasminogen activator (TPA), as well as other known anticlotting agents.
The anticlotting agent may be applied to at least a portion of that part of the sensor that is to be implanted. The anticlotting agent may be applied, for example, by bath, spraying, brushing, or dipping. The anticlotting agent is allowed to dry on the sensor. The anticlotting agent may be immobilized on the surface of the sensor or it may be allowed to diffuse away from the sensor surface. Typically, the quantities of anticlotting agent disposed on the sensor are far below the amounts typically used for treatment of medical conditions involving blood clots and, therefore, have only a limited, localized effect.
Insertion Device
An insertion device can be used to subcutaneously insert the self-powered analyte sensor into the patient. The insertion device is typically formed using structurally rigid materials, such as metal or rigid plastic. Exemplary materials include stainless steel and ABS (acrylonitrile-butadiene-styrene) plastic. In some embodiments, the insertion device is pointed and/or sharp at the tip to facilitate penetration of the skin of the patient. A sharp, thin insertion device may reduce pain felt by the patient upon insertion of the self-powered analyte sensor. In other embodiments, the tip of the insertion device has other shapes, including a blunt or flat shape. These embodiments may be particularly useful when the insertion device does not penetrate the skin but rather serves as a structural support for the sensor as the sensor is pushed into the skin.
Sensor Control Unit
The sensor control unit can be integrated in the sensor, part or all of which is subcutaneously implanted or it can be configured to be placed on the skin of a patient. The sensor control unit is optionally formed in a shape that is comfortable to the patient and which may permit concealment, for example, under a patient's clothing. The thigh, leg, upper arm, shoulder, or abdomen are convenient parts of the patient's body for placement of the sensor control unit to maintain concealment. However, the sensor control unit may be positioned on other portions of the patient's body. One embodiment of the sensor control unit has a thin, oval shape to enhance concealment. However, other shapes and sizes may be used.
The particular profile, as well as the height, width, length, weight, and volume of the sensor control unit may vary and depends, at least in part, on the components and associated functions included in the sensor control unit. In general, the sensor control unit includes a housing typically formed as a single integral unit that rests on the skin of the patient. The housing typically contains most or all of the electronic components of the sensor control unit.
The housing of the sensor control unit may be formed using a variety of materials, including, for example, plastic and polymeric materials, particularly rigid thermoplastics and engineering thermoplastics. Suitable materials include, for example, polyvinyl chloride, polyethylene, polypropylene, polystyrene, ABS polymers, and copolymers thereof. The housing of the sensor control unit may be formed using a variety of techniques including, for example, injection molding, compression molding, casting, and other molding methods. Hollow or recessed regions may be formed in the housing of the sensor control unit. The electronic components of the sensor control unit and/or other items, including a battery or a speaker for an audible alarm, may be placed in the hollow or recessed areas.
The sensor control unit is typically attached to the skin of the patient, for example, by adhering the sensor control unit directly to the skin of the patient with an adhesive provided on at least a portion of the housing of the sensor control unit which contacts the skin or by suturing the sensor control unit to the skin through suture openings in the sensor control unit.
When positioned on the skin of a patient, the sensor and the electronic components within the sensor control unit are coupled via conductive contacts. The one or more working electrodes, counter electrode, and optional temperature probe are attached to individual conductive contacts. For example, the conductive contacts are provided on the interior of the sensor control unit. Other embodiments of the sensor control unit have the conductive contacts disposed on the exterior of the housing. The placement of the conductive contacts is such that they are in contact with the contact pads on the sensor when the sensor is properly positioned within the sensor control unit.
Sensor Control Unit Electronics
The sensor control unit also typically includes at least a portion of the electronic components that measure the sensor current and the analyte monitoring device system. The electronic components of the sensor control unit typically include a power supply for operating the sensor control unit, a sensor circuit for obtaining signals from the sensor, a measurement circuit that converts sensor signals to a desired format, and a processing circuit that, at minimum, obtains signals from the sensor circuit and/or measurement circuit and provides the signals to an optional transmitter. In some embodiments, the processing circuit may also partially or completely evaluate the signals from the sensor and convey the resulting data to the optional transmitter and/or activate an optional alarm system if the analyte level exceeds a threshold. The processing circuit often includes digital logic circuitry.
The sensor control unit may optionally contain a transmitter for transmitting the sensor signals or processed data from the processing circuit to a receiver/display unit; a data storage unit for temporarily or permanently storing data from the processing circuit; a temperature probe circuit for receiving signals from and operating a temperature probe; a reference voltage generator for providing a reference voltage for comparison with sensor-generated signals; and/or a watchdog circuit that monitors the operation of the electronic components in the sensor control unit.
Moreover, the sensor control unit may also include digital and/or analog components utilizing semiconductor devices, including transistors. To operate these semiconductor devices, the sensor control unit may include other components including, for example, a bias control generator to correctly bias analog and digital semiconductor devices, an oscillator to provide a clock signal, and a digital logic and timing component to provide timing signals and logic operations for the digital components of the circuit.
As an example of the operation of these components, the sensor circuit and the optional temperature probe circuit provide raw signals from the sensor to the measurement circuit. The measurement circuit converts the raw signals to a desired format, using for example, a current-to-voltage converter, current-to-frequency converter, and/or a binary counter or other indicator that produces a signal proportional to the absolute value of the raw signal. This may be used, for example, to convert the raw signal to a format that can be used by digital logic circuits. The processing circuit may then, optionally, evaluate the data and provide commands to operate the electronics.
Calibration
Sensors may be configured to require no system calibration or no user calibration. For example, a sensor may be factory calibrated and need not require further calibrating. In certain embodiments, calibration may be required, but may be done without user intervention, i.e., may be automatic. In those embodiments in which calibration by the user is required, the calibration may be according to a predetermined schedule or may be dynamic, i.e., the time for which may be determined by the system on a real-time basis according to various factors, including, but not limited to, glucose concentration and/or temperature and/or rate of change of glucose, etc.
In addition to a transmitter, an optional receiver may be included in the sensor control unit. In some cases, the transmitter is a transceiver, operating as both a transmitter and a receiver. The receiver may be used to receive calibration data for the sensor. The calibration data may be used by the processing circuit to correct signals from the sensor. This calibration data may be transmitted by the receiver/display unit or from some other source such as a control unit in a doctor's office. In addition, the optional receiver may be used to receive a signal from the receiver/display units to direct the transmitter, for example, to change frequencies or frequency bands, to activate or deactivate the optional alarm system and/or to direct the transmitter to transmit at a higher rate.
Calibration data may be obtained in a variety of ways. For instance, the calibration data may simply be factory-determined calibration measurements which can be input into the sensor control unit using the receiver or may alternatively be stored in a calibration data storage unit within the sensor control unit itself (in which case a receiver may not be needed). The calibration data storage unit may be, for example, a readable or readable/writeable memory circuit.
Calibration may be accomplished using an in vitro test strip (or other reference), e.g., a small sample test strip such as a test strip that requires less than about 1 microliter of sample (for example FreeStyle® blood glucose monitoring test strips from Abbott Diabetes Care). For example, test strips that require less than about 1 nanoliter of sample may be used. In certain embodiments, a sensor may be calibrated using only one sample of body fluid per calibration event. For example, a user need only lance a body part one time to obtain sample for a calibration event (e.g., for a test strip), or may lance more than one time within a short period of time if an insufficient volume of sample is firstly obtained. Embodiments include obtaining and using multiple samples of body fluid for a given calibration event, where glucose values of each sample are substantially similar. Data obtained from a given calibration event may be used independently to calibrate or combined with data obtained from previous calibration events, e.g., averaged including weighted averaged, etc., to calibrate. In certain embodiments, a system need only be calibrated once by a user, where recalibration of the system is not required.
Alternative or additional calibration data may be provided based on tests performed by a doctor or some other professional or by the patient. For example, it is common for diabetic individuals to determine their own blood glucose concentration using commercially available testing kits. The results of this test is input into the sensor control unit either directly, if an appropriate input device (e.g., a keypad, an optical signal receiver, or a port for connection to a keypad or computer) is incorporated in the sensor control unit, or indirectly by inputting the calibration data into the receiver/display unit and transmitting the calibration data to the sensor control unit.
Other methods of independently determining analyte levels may also be used to obtain calibration data. This type of calibration data may supplant or supplement factory-determined calibration values.
In some embodiments of the invention, calibration data may be required at periodic intervals, for example, every eight hours, once a day, or once a week, to confirm that accurate analyte levels are being reported. Calibration may also be required each time a new sensor is implanted or if the sensor exceeds a threshold minimum or maximum value or if the rate of change in the sensor signal exceeds a threshold value. In some cases, it may be necessary to wait a period of time after the implantation of the sensor before calibrating to allow the sensor to achieve equilibrium. In some embodiments, the sensor is calibrated only after it has been inserted. In other embodiments, no calibration of the sensor is needed.
Analyte Monitoring Device
In some embodiments of the invention, the analyte monitoring device includes a sensor control unit and a self-powered analyte sensor. In some embodiments, the processing circuit of the sensor control unit is able to determine a level of the analyte and activate an alarm system if the analyte level exceeds a threshold. The sensor control unit, in these embodiments, has an alarm system and may also include a display, such as an LCD or LED display.
A threshold value is exceeded if the datapoint has a value that is beyond the threshold value in a direction indicating a particular condition. For example, a datapoint which correlates to a glucose level of 200 mg/dL exceeds a threshold value for hyperglycemia of 180 mg/dL, because the datapoint indicates that the patient has entered a hyperglycemic state. As another example, a datapoint which correlates to a glucose level of 65 mg/dL exceeds a threshold value for hypoglycemia of 70 mg/dL because the datapoint indicates that the patient is hypoglycemic as defined by the threshold value. However, a datapoint which correlates to a glucose level of 75 mg/dL would not exceed the same threshold value for hypoglycemia because the datapoint does not indicate that particular condition as defined by the chosen threshold value.
An alarm may also be activated if the sensor readings indicate a value that is beyond a measurement range of the sensor. For glucose, the physiologically relevant measurement range is typically about 30 to 500 mg/dL, including about 40-300 mg/dL and about 50-250 mg/dL, of glucose in the interstitial fluid.
The alarm system may also, or alternatively, be activated when the rate of change or acceleration of the rate of change in analyte level increase or decrease reaches or exceeds a threshold rate or acceleration. For example, in the case of a subcutaneous glucose monitor, the alarm system might be activated if the rate of change in glucose concentration exceeds a threshold value which might indicate that a hyperglycemic or hypoglycemic condition is likely to occur.
A system may also include system alarms that notify a user of system information such as battery condition, calibration, sensor dislodgment, sensor malfunction, etc. Alarms may be, for example, auditory and/or visual. Other sensory-stimulating alarm systems may be used including alarm systems which heat, cool, vibrate, or produce a mild electrical shock when activated.
Drug Delivery System
The subject invention also includes sensors used in sensor-based drug delivery systems. The system may provide a drug to counteract the high or low level of the analyte in response to the signals from one or more sensors. Alternatively, the system may monitor the drug concentration to ensure that the drug remains within a desired therapeutic range. The drug delivery system may include one or more (e.g., two or more) sensors, a processing unit such as a transmitter, a receiver/display unit, and a drug administration system. In some cases, some or all components may be integrated in a single unit. A sensor-based drug delivery system may use data from the one or more sensors to provide necessary input for a control algorithm/mechanism to adjust the administration of drugs, e.g., automatically or semi-automatically. As an example, a glucose sensor may be used to control and adjust the administration of insulin from an external or implanted insulin pump.
Each of the various references, presentations, publications, provisional and/or non-provisional U.S. Patent Applications, U.S. patents, non-U.S. Patent Applications, and/or non-U.S. patents that have been identified herein, is incorporated herein in its entirety by this reference.
Other aspects, advantages, and modifications within the scope of the invention will be apparent to those skilled in the art to which the invention pertains. Various modifications, processes, as well as numerous structures to which the embodiments of the invention may be applicable will be readily apparent to those of skill in the art to which the invention is directed upon review of the specification. Various aspects and features of the invention may have been explained or described in relation to understandings, beliefs, theories, underlying assumptions, and/or working or prophetic examples, although it will be understood that the invention is not bound to any particular understanding, belief, theory, underlying assumption, and/or working or prophetic example. Although various aspects and features of the invention may have been described largely with respect to applications, or more specifically, medical applications, involving diabetic humans, it will be understood that such aspects and features also relate to any of a variety of applications involving non-diabetic humans and any and all other animals. Further, although various aspects and features of the invention may have been described largely with respect to applications involving partially implanted sensors, such as transcutaneous or subcutaneous sensors, it will be understood that such aspects and features also relate to any of a variety of sensors that are suitable for use in connection with the body of an animal or a human, such as those suitable for use as fully implanted in the body of an animal or a human. Finally, although the various aspects and features of the invention have been described with respect to various embodiments and specific examples herein, all of which may be made or carried out conventionally, it will be understood that the invention is entitled to protection within the full scope of the appended claims.
The following examples are put forth so as to provide those of ordinary skill in the art with a complete disclosure and description of how to make and use the present invention, and are not intended to limit the scope of what the inventors regard as their invention nor are they intended to represent that the experiments below are all or the only experiments performed. Efforts have been made to ensure accuracy with respect to numbers used (e.g. amounts, temperature, etc.) but some experimental errors and deviations should be accounted for. Unless indicated otherwise, parts are parts by weight, molecular weight is weight average molecular weight, temperature is in degrees Centigrade, and pressure is at or near atmospheric.
A self-powered glucose sensor was constructed using either a redox polymer or platinum-carbon cloth (Pt—C) as an oxygen reducing species on the counter electrode and glucose oxidase (GOx) as glucose oxidizing species on the working electrode, which also acts as the analyte responsive enzyme of the glucose sensor. Fabrication of the self-powered analyte sensors is described below.
Working electrode (glucose oxidizing anode): 30 nL of a solution including 8.2 mg/mL GOx, 6.8 mg/mL redox polymer (a poly(vinylpyridine) derivative containing covalently bound Os complexes), and 5 mg/mL PEG 400 crosslinker, all dissolved in 10 mM HEPES, was deposited on the working electrode of a continuous glucose monitor sensor, such as a Navigator® electrode, to form an active area of about 0.1 mm2. This electrode was overlaid with the flux-limiting membrane, approximately 50 micron thick coating of the membrane polymer (a poly(vinylpyridine) derivative) crosslinked with tri-glycidyl glycerol.
Counter electrode (oxygen reducing cathode): 150 nL of a solution including 5.2 mg/mL redox polymer dissolved in PBS, was deposited on the entirety of the working electrode of a continuous glucose monitor sensor, such as a Navigator® electrode, to form an active area of about 0.9 mm2. This electrode was overlaid with a flux-limiting membrane, an approximately 50 micron thick coating of the membrane polymer (a poly(vinylpyridine) derivative), crosslinked with tri-glycidyl glycerol.
A 4.67 Mohm resistor was placed between these two electrodes, and the resulting spontaneous current was measured, as the sensor pair was exposed to varying glucose concentrations and at varying time points. The output of this circuit was compared to that from the identical working electrode in 3-electrode mode (paired through a potentiostat with an Ag/AgCl reference and a carbon counter electrode), as well as to the same working electrode paired in self-powered mode with a Pt/carbon counter electrode. The results are shown in
Glucose aliquots were added to produce concentrations of 4, 12, 20, and 28 mM, as evident from the humps in
Calibration curves from the data shown in
For many applications, it is desirable to intermittently address or query a sensor for current values, while leaving the sensor unpowered in the intervening intervals, since this can result in enormous savings in power consumption, and therefore overall size of the analyte sensor. Unfortunately, it is not feasible to intermittently power the current generation of a standard continuous glucose sensors since they have a relatively long response time. This is illustrated in
As shown in
The oxygen reducing counter electrode of a self-powered sensor can also be constructed of an oxygen reducing enzyme, in addition to the other materials previously discussed, such as Pt, Pt black, and redox polymer. The enzyme, in this case bilirubin oxidase (BOD) can accept electrons directly from oxygen, then transmit them to the counter electrode through a cross-linked redox polymer. Fabrication of the self-powered analyte sensors is described below.
Working electrode (glucose oxidizing anode): 30 nL of a solution including 8.2 mg/mL GOx, 6.8 mg/mL redox polymer (a poly(vinylpyridine) derivative containing covalently bound Os complexes), and 5 mg/mL PEG 400 cross-linker, all dissolved in 10 mM HEPES, was deposited on the working electrode of a continuous glucose sensor, such as a Navigator® sensor, to form an active area of about 0.1 mm2. This electrode was overlaid with a flux-limiting membrane, an approximately 50 micron thick coating of the membrane polymer (a poly(vinylpyridine) derivative)cross-linked with tri-glycidyl glycerol.
Counter electrode (oxygen reducing cathode): 150 nL of a solution including 12.6 mg/mL BOD, 5.2 mg/mL redox polymer, and 5.4 mg/mL PEG 400 cross-linker, all dissolved in PBS, was deposited on the entirety of the working electrode of a second continuous glucose sensor, such as a Navigator® sensor, to form an active area of about 0.9 mm2. This electrode was overlaid with a flux-limiting membrane, an approximately 50 micron thick coating of the membrane polymer (a poly(vinylpyridine) derivative) cross-linked with tri-glycidyl glycerol.
A 4.67 Mohm resistor was placed between these two electrodes, and the resulting spontaneous current was measured, as the sensor pair was exposed to varying glucose concentrations. The output of this circuit was compared to that from the identical working electrode in 3-electrode mode (paired through a potentiostat with an Ag/AgCl reference and a carbon counter electrode), as well as to the same working electrode paired in self-powered mode with a Pt/carbon cloth counter electrode. The results are shown in
A self-powered glucose sensor was constructed using either a redox polymer or platinum (Pt) as an oxygen reducing species on the counter electrode and glucose dehydrogenase complexed with flavin adenine dinucleotide (FADGDH) on the working electrode, which also acts the analyte responsive enzyme for the sensor. Fabrication of the self-powered analyte sensors is described below.
Working electrode (anode): 30 nL of a solution including 8.2 mg/mL FADGDH, 6.8 mg/mL redox polymer (a poly(vinylpyridine) derivative containing covalently bound Os complexes), and 5 mg/mL PEG 400 crosslinker, all dissolved in 10 mM HEPES, was deposited on the working electrode of a continuous glucose monitor sensor, such as a Navigator® electrode, to form an active area of about 0.1 mm2. This electrode was overlaid with the flux-limiting membrane, an approximately 50 micron thick coating of the membrane polymer (a poly(vinylpyridine) derivative) crosslinked with tri-glycidyl glycerol.
Counter electrode (cathode): 150 nL of a solution including 12.6 mg/mL BOD, 5.2 mg/mL redox polymer, and 5.4 mg/mL PEG 400 cross-linker, all dissolved in PBS, was deposited on the entirety of the working electrode of a second continuous glucose sensor, such as a Navigator® sensor, to form an active area of about 0.9 mm2. This electrode was overlaid with a flux-limiting membrane, an approximately 50 micron thick coating of the membrane polymer (a poly(vinylpyridine) derivative) cross-linked with tri-glycidyl glycerol.
A 4.67 Mohm resistor was placed between these two electrodes, and the resulting spontaneous current was measured, as the sensor pair was exposed to varying glucose concentrations and at varying time points. The output of this circuit was compared to that from the identical working electrode in 3-electrode mode (paired through a potentiostat with an Ag/AgCl reference and a carbon counter electrode), as well as to the same working electrode paired in self-powered mode with a Pt/carbon counter electrode. The results are shown in
Glucose aliquots were added to produce concentrations of 4, 12, 20, and 28 mM. Calibration curves from the three sensors are shown in
In order to have a linear response over the concentration range of analytical interest, the correct value of R must be chosen for the resistor which couples together the working electrode and the counter electrode. This is illustrated by the following experiment. A self powered sensor was made from a GOx and redox polymer working electrode, and a Pt-carbon cloth counter electrode, coupled together with a variety of resistors, ranging in resistance from 0 MΩ to 100 MΩ. This sensor was then immersed in 30 mM (540 mg/dL) glucose (the top of the analytical range of interest) in PBS buffer at 37° C., and the resulting spontaneous steady state current was measured for sensors configured with different R values. The results are shown in
Note that the current is essentially independent of resistance for resistance values of 10 MΩ or less, then falls off gradually at higher potentials, as the potential drop becomes a substantial fraction of the open circuit potential. Therefore, to correctly configure the sensor, a value of R should be selected that does not reduce the current output from its maximum value, as illustrated by the plateau in the graph of
Another way to elucidate the correct range for R is in terms of the potential across the resistor. In the sensor described by the graph above, the open circuit potential (shown) is 474 mV. When the potential drop across the resistor is more than half of the open circuit potential, the output of the sensor begins to decline. The potential drop across the resistor is, for example, less than 75% of the open circuit potential, such as less than 50% of the open circuit potential, and less than 25% of the open circuit potential in certain embodiments.
An in vivo self-powered glucose sensor used for subcutaneous implantation was constructed using the two carbon electrodes on a continuous glucose monitoring electrode of a Navigator® sensor: the working electrode, the counter electrode, and a resistor disposed in between the working electrode and the counter electrode and having an R value of 4.85 Mohm.
The working electrode was constructed to include redox polymer, glucose oxidase, and cross-linker sensing layer. The counter electrode was constructed to include a layer of platinum catalyst formed by depositing 50 nL 20 mg/ml platinum-carbon black (HiSPEC 8000, Alfa Aesar) suspension in 80% ethanol-20% water. The sensor was then dip-coated with a flux-limiting polymer membrane. After sterilization, the sensor was subcutaneously implanted with an insertion device. The data was collected with a transmitter that sent via radio-frequency, the self-powered sensor current signal to a standard receiver. Reference glucose values were obtained through frequent measurement with FreeStyle® glucose strips. The results in
The preceding merely illustrates the principles of the invention. It will be appreciated that those skilled in the art will be able to devise various arrangements which, although not explicitly described or shown herein, embody the principles of the invention and are included within its spirit and scope. Furthermore, all examples and conditional language recited herein are principally intended to aid the reader in understanding the principles of the invention and the concepts contributed by the inventors to furthering the art, and are to be construed as being without limitation to such specifically recited examples and conditions. Moreover, all statements herein reciting principles, aspects, and embodiments of the invention as well as specific examples thereof, are intended to encompass both structural and functional equivalents thereof. Additionally, it is intended that such equivalents include both currently known equivalents and equivalents developed in the future, i.e., any elements developed that perform the same function, regardless of structure. The scope of the present invention, therefore, is not intended to be limited to the exemplary embodiments shown and described herein. Rather, the scope and spirit of present invention is embodied by the appended claims.
| Number | Name | Date | Kind |
|---|---|---|---|
| 3260656 | Ross, Jr. | Jul 1966 | A |
| 3653841 | Klein | Apr 1972 | A |
| 3719564 | Lilly, Jr. et al. | Mar 1973 | A |
| 3774243 | Ng et al. | Nov 1973 | A |
| 3776832 | Oswin et al. | Dec 1973 | A |
| 3811950 | Avampato et al. | May 1974 | A |
| 3837339 | Aisenberg et al. | Sep 1974 | A |
| 3861397 | Rao et al. | Jan 1975 | A |
| 3972320 | Kalman | Aug 1976 | A |
| 3979274 | Newman | Sep 1976 | A |
| 4008717 | Kowarski | Feb 1977 | A |
| 4016866 | Lawton | Apr 1977 | A |
| 4055175 | Clemens et al. | Oct 1977 | A |
| 4059406 | Fleet | Nov 1977 | A |
| 4076596 | Connery et al. | Feb 1978 | A |
| 4098574 | Dappen | Jul 1978 | A |
| 4100048 | Pompei et al. | Jul 1978 | A |
| 4117202 | Beck | Sep 1978 | A |
| 4151845 | Clemens | May 1979 | A |
| 4168205 | Danninger et al. | Sep 1979 | A |
| 4172770 | Semersky et al. | Oct 1979 | A |
| 4178916 | McNamara | Dec 1979 | A |
| 4206755 | Klein | Jun 1980 | A |
| 4224125 | Nakamura et al. | Sep 1980 | A |
| 4240438 | Updike et al. | Dec 1980 | A |
| 4240889 | Yoda et al. | Dec 1980 | A |
| 4247297 | Berti et al. | Jan 1981 | A |
| 4271119 | Columbus | Jun 1981 | A |
| 4318784 | Higgins et al. | Mar 1982 | A |
| 4340458 | Lerner et al. | Jul 1982 | A |
| 4356074 | Johnson | Oct 1982 | A |
| 4365637 | Johnson | Dec 1982 | A |
| 4366033 | Richter et al. | Dec 1982 | A |
| 4375399 | Havas et al. | Mar 1983 | A |
| 4384586 | Christiansen | May 1983 | A |
| 4392933 | Nakamura et al. | Jul 1983 | A |
| 4401122 | Clark, Jr. | Aug 1983 | A |
| 4404066 | Johnson | Sep 1983 | A |
| 4407959 | Tsuji et al. | Oct 1983 | A |
| 4418148 | Oberhardt | Nov 1983 | A |
| 4420564 | Tsuji et al. | Dec 1983 | A |
| 4427770 | Chen et al. | Jan 1984 | A |
| 4431004 | Bessman et al. | Feb 1984 | A |
| 4436094 | Cerami | Mar 1984 | A |
| 4436812 | Endoh et al. | Mar 1984 | A |
| 4440175 | Wilkins | Apr 1984 | A |
| 4444892 | Malmros | Apr 1984 | A |
| 4450842 | Zick et al. | May 1984 | A |
| 4461691 | Frank | Jul 1984 | A |
| 4469110 | Slama | Sep 1984 | A |
| 4477314 | Richter et al. | Oct 1984 | A |
| 4483924 | Tsuji et al. | Nov 1984 | A |
| 4484987 | Gough | Nov 1984 | A |
| 4522690 | Venkatasetty | Jun 1985 | A |
| 4524114 | Samuels et al. | Jun 1985 | A |
| 4526661 | Steckhan et al. | Jul 1985 | A |
| 4534356 | Papadakis | Aug 1985 | A |
| 4538616 | Rogoff | Sep 1985 | A |
| 4543955 | Schroeppel | Oct 1985 | A |
| 4545382 | Higgins et al. | Oct 1985 | A |
| 4552840 | Rifler | Nov 1985 | A |
| 4560534 | Kung et al. | Dec 1985 | A |
| 4571292 | Liu et al. | Feb 1986 | A |
| 4573994 | Fischell et al. | Mar 1986 | A |
| 4581336 | Malloy et al. | Apr 1986 | A |
| 4595011 | Phillips | Jun 1986 | A |
| 4595479 | Kimura et al. | Jun 1986 | A |
| 4619754 | Niki et al. | Oct 1986 | A |
| 4633878 | Bombardieri | Jan 1987 | A |
| 4637403 | Garcia et al. | Jan 1987 | A |
| 4650547 | Gough | Mar 1987 | A |
| 4655880 | Liu | Apr 1987 | A |
| 4655885 | Hill et al. | Apr 1987 | A |
| 4656197 | Yoshimura et al. | Apr 1987 | A |
| 4671288 | Gough | Jun 1987 | A |
| 4679562 | Luksha | Jul 1987 | A |
| 4680268 | Clark, Jr. | Jul 1987 | A |
| 4682602 | Prohaska | Jul 1987 | A |
| 4684537 | Graetzel et al. | Aug 1987 | A |
| 4685463 | Williams | Aug 1987 | A |
| 4703756 | Gough et al. | Nov 1987 | A |
| 4711245 | Higgins et al. | Dec 1987 | A |
| 4717673 | Wrighton et al. | Jan 1988 | A |
| 4721601 | Wrighton et al. | Jan 1988 | A |
| 4726378 | Kaplan | Feb 1988 | A |
| 4750496 | Reinhart et al. | Jun 1988 | A |
| 4757022 | Shults et al. | Jul 1988 | A |
| 4758323 | Davis et al. | Jul 1988 | A |
| 4759371 | Franetzki | Jul 1988 | A |
| 4759828 | Young et al. | Jul 1988 | A |
| 4764416 | Ueyama et al. | Aug 1988 | A |
| 4776944 | Janata et al. | Oct 1988 | A |
| 4781798 | Gough | Nov 1988 | A |
| 4784736 | Lonsdale et al. | Nov 1988 | A |
| 4795707 | Niiyama et al. | Jan 1989 | A |
| 4805624 | Yao et al. | Feb 1989 | A |
| 4806468 | Wagner et al. | Feb 1989 | A |
| 4813424 | Wilkins | Mar 1989 | A |
| 4815469 | Cohen et al. | Mar 1989 | A |
| 4820399 | Senda et al. | Apr 1989 | A |
| 4820636 | Hill et al. | Apr 1989 | A |
| 4822337 | Newhouse et al. | Apr 1989 | A |
| 4830959 | McNeil et al. | May 1989 | A |
| 4832797 | Vadgama et al. | May 1989 | A |
| 4840893 | Hill et al. | Jun 1989 | A |
| 4848351 | Finch | Jul 1989 | A |
| 4871351 | Feingold | Oct 1989 | A |
| 4871440 | Nagata et al. | Oct 1989 | A |
| 4890620 | Gough | Jan 1990 | A |
| 4894137 | Takizawa et al. | Jan 1990 | A |
| 4894253 | Heineman et al. | Jan 1990 | A |
| 4897162 | Lewandowski et al. | Jan 1990 | A |
| 4897173 | Nankai et al. | Jan 1990 | A |
| 4909908 | Ross et al. | Mar 1990 | A |
| 4911794 | Parce et al. | Mar 1990 | A |
| 4919141 | Zier et al. | Apr 1990 | A |
| 4919767 | Vadgama et al. | Apr 1990 | A |
| 4923586 | Katayama et al. | May 1990 | A |
| 4927516 | Yamaguchi et al. | May 1990 | A |
| 4935105 | Churchouse | Jun 1990 | A |
| 4935345 | Guilbeau et al. | Jun 1990 | A |
| 4936956 | Wrighton | Jun 1990 | A |
| 4938860 | Wogoman | Jul 1990 | A |
| 4942127 | Wada et al. | Jul 1990 | A |
| 4945045 | Forrest et al. | Jul 1990 | A |
| 4950378 | Nagata | Aug 1990 | A |
| 4953552 | DeMarzo | Sep 1990 | A |
| 4968400 | Shimomura et al. | Nov 1990 | A |
| 4970145 | Bennetto et al. | Nov 1990 | A |
| 4974929 | Curry | Dec 1990 | A |
| 4986271 | Wilkins | Jan 1991 | A |
| 4994167 | Shults et al. | Feb 1991 | A |
| 5034192 | Wrighton et al. | Jul 1991 | A |
| 5037527 | Hayashi et al. | Aug 1991 | A |
| 5070535 | Hochmair et al. | Dec 1991 | A |
| 5078854 | Burgess | Jan 1992 | A |
| 5082550 | Rishpon et al. | Jan 1992 | A |
| 5082786 | Nakamoto | Jan 1992 | A |
| 5089112 | Skotheim et al. | Feb 1992 | A |
| 5094951 | Rosenberg | Mar 1992 | A |
| 5096560 | Takai et al. | Mar 1992 | A |
| 5096836 | Macho et al. | Mar 1992 | A |
| 5101814 | Palti | Apr 1992 | A |
| 5108564 | Szuminsky et al. | Apr 1992 | A |
| 5109850 | Blanco et al. | May 1992 | A |
| 5120420 | Nankai et al. | Jun 1992 | A |
| 5120421 | Glass et al. | Jun 1992 | A |
| 5126034 | Carter et al. | Jun 1992 | A |
| 5126247 | Palmer et al. | Jun 1992 | A |
| 5130009 | Marsoner et al. | Jul 1992 | A |
| 5133856 | Yamaguchi et al. | Jul 1992 | A |
| 5140393 | Hijikihigawa et al. | Aug 1992 | A |
| 5141868 | Shanks et al. | Aug 1992 | A |
| 5161532 | Joseph | Nov 1992 | A |
| 5165407 | Wilson et al. | Nov 1992 | A |
| 5168046 | Hamamoto et al. | Dec 1992 | A |
| 5174291 | Schoonen et al. | Dec 1992 | A |
| 5185256 | Nankai et al. | Feb 1993 | A |
| 5192415 | Yoshioka et al. | Mar 1993 | A |
| 5192416 | Wang et al. | Mar 1993 | A |
| 5198367 | Aizawa et al. | Mar 1993 | A |
| 5200051 | Cozzette et al. | Apr 1993 | A |
| 5202261 | Musho et al. | Apr 1993 | A |
| 5205920 | Oyama et al. | Apr 1993 | A |
| 5206145 | Cattell | Apr 1993 | A |
| 5208154 | Weaver et al. | May 1993 | A |
| 5217595 | Smith et al. | Jun 1993 | A |
| 5227042 | Zawodzinski et al. | Jul 1993 | A |
| 5229282 | Yoshioka et al. | Jul 1993 | A |
| 5250439 | Musho et al. | Oct 1993 | A |
| 5262035 | Gregg et al. | Nov 1993 | A |
| 5262305 | Heller et al. | Nov 1993 | A |
| 5264103 | Yoshioka et al. | Nov 1993 | A |
| 5264104 | Gregg et al. | Nov 1993 | A |
| 5264105 | Gregg et al. | Nov 1993 | A |
| 5264106 | McAleer et al. | Nov 1993 | A |
| 5271815 | Wong | Dec 1993 | A |
| 5272060 | Hamamoto et al. | Dec 1993 | A |
| 5278079 | Gubinski et al. | Jan 1994 | A |
| 5286362 | Hoenes et al. | Feb 1994 | A |
| 5286364 | Yacynych et al. | Feb 1994 | A |
| 5288636 | Pollmann et al. | Feb 1994 | A |
| 5293546 | Tadros et al. | Mar 1994 | A |
| 5310885 | Maier et al. | May 1994 | A |
| 5320715 | Berg | Jun 1994 | A |
| 5320725 | Gregg et al. | Jun 1994 | A |
| 5326449 | Cunningham | Jul 1994 | A |
| 5337747 | Neftel | Aug 1994 | A |
| 5352348 | Young et al. | Oct 1994 | A |
| 5356786 | Heller et al. | Oct 1994 | A |
| 5364797 | Olson et al. | Nov 1994 | A |
| 5368028 | Palti | Nov 1994 | A |
| 5372133 | Hogen Esch | Dec 1994 | A |
| 5378628 | Gratzel et al. | Jan 1995 | A |
| 5380422 | Negishi et al. | Jan 1995 | A |
| 5382346 | Uenoyama et al. | Jan 1995 | A |
| 5387327 | Khan | Feb 1995 | A |
| 5390671 | Lord et al. | Feb 1995 | A |
| 5391250 | Cheney, II et al. | Feb 1995 | A |
| 5393903 | Gratzel et al. | Feb 1995 | A |
| 5395504 | Saurer | Mar 1995 | A |
| 5411647 | Johnson et al. | May 1995 | A |
| 5413690 | Kost et al. | May 1995 | A |
| 5422246 | Koopal et al. | Jun 1995 | A |
| 5437973 | Vadgama et al. | Aug 1995 | A |
| 5437999 | Diebold et al. | Aug 1995 | A |
| 5443701 | Willner et al. | Aug 1995 | A |
| 5478751 | Oosta et al. | Dec 1995 | A |
| 5478754 | Brandt et al. | Dec 1995 | A |
| 5494562 | Maley et al. | Feb 1996 | A |
| 5496453 | Uenoyama et al. | Mar 1996 | A |
| 5497772 | Schulman et al. | Mar 1996 | A |
| 5501956 | Wada et al. | Mar 1996 | A |
| 5506144 | Sundrehagen | Apr 1996 | A |
| 5507288 | Booker et al. | Apr 1996 | A |
| 5508171 | Walling et al. | Apr 1996 | A |
| 5514253 | Davis et al. | May 1996 | A |
| 5520787 | Hanagan et al. | May 1996 | A |
| 5525511 | D'Costa | Jun 1996 | A |
| 5526120 | Jina et al. | Jun 1996 | A |
| 5531878 | Vadgama et al. | Jul 1996 | A |
| 5552027 | Birkle et al. | Sep 1996 | A |
| 5556524 | Albers | Sep 1996 | A |
| 5565085 | Ikeda et al. | Oct 1996 | A |
| 5567302 | Song et al. | Oct 1996 | A |
| 5568806 | Cheney, II et al. | Oct 1996 | A |
| 5569186 | Lord et al. | Oct 1996 | A |
| 5575895 | Ikeda et al. | Nov 1996 | A |
| 5580527 | Bell et al. | Dec 1996 | A |
| 5582184 | Erickson et al. | Dec 1996 | A |
| 5582697 | Ikeda et al. | Dec 1996 | A |
| 5582698 | Flaherty et al. | Dec 1996 | A |
| 5586553 | Halili et al. | Dec 1996 | A |
| 5589326 | Deng et al. | Dec 1996 | A |
| 5593852 | Heller et al. | Jan 1997 | A |
| 5596150 | Arndt et al. | Jan 1997 | A |
| 5609749 | Yamauchi et al. | Mar 1997 | A |
| 5617851 | Lipkovker | Apr 1997 | A |
| 5628890 | Carter et al. | May 1997 | A |
| 5639672 | Burd et al. | Jun 1997 | A |
| 5650062 | Ikeda et al. | Jul 1997 | A |
| 5651869 | Yoshioka et al. | Jul 1997 | A |
| 5660163 | Schulman et al. | Aug 1997 | A |
| 5665222 | Heller et al. | Sep 1997 | A |
| 5667653 | Schneider et al. | Sep 1997 | A |
| 5670031 | Hintsche et al. | Sep 1997 | A |
| 5680858 | Hansen et al. | Oct 1997 | A |
| 5682233 | Brinda | Oct 1997 | A |
| 5695623 | Michel et al. | Dec 1997 | A |
| 5695947 | Guo et al. | Dec 1997 | A |
| 5708247 | McAleer et al. | Jan 1998 | A |
| 5711861 | Ward et al. | Jan 1998 | A |
| 5711862 | Sakoda et al. | Jan 1998 | A |
| 5720862 | Hamamoto et al. | Feb 1998 | A |
| 5727548 | Hill et al. | Mar 1998 | A |
| 5741211 | Renirie et al. | Apr 1998 | A |
| 5741688 | Oxenboll et al. | Apr 1998 | A |
| 5746217 | Erickson et al. | May 1998 | A |
| 5770028 | Maley et al. | Jun 1998 | A |
| 5791344 | Schulman et al. | Aug 1998 | A |
| 5804048 | Wong et al. | Sep 1998 | A |
| 5804401 | Gardiol et al. | Sep 1998 | A |
| 5807747 | Wallworth et al. | Sep 1998 | A |
| 5820570 | Erickson et al. | Oct 1998 | A |
| 5820622 | Gross et al. | Oct 1998 | A |
| 5830341 | Gilmartin | Nov 1998 | A |
| 5834224 | Ruger et al. | Nov 1998 | A |
| 5837454 | Cozzette et al. | Nov 1998 | A |
| 5837546 | Allen et al. | Nov 1998 | A |
| 5842983 | Abel et al. | Dec 1998 | A |
| 5846702 | Deng et al. | Dec 1998 | A |
| 5846744 | Athey et al. | Dec 1998 | A |
| 5857983 | Douglas et al. | Jan 1999 | A |
| 5879311 | Duchon et al. | Mar 1999 | A |
| 5888787 | Chen et al. | Mar 1999 | A |
| 5906921 | Ikeda et al. | May 1999 | A |
| 5951492 | Douglas et al. | Sep 1999 | A |
| 5951493 | Douglas et al. | Sep 1999 | A |
| 5972199 | Heller et al. | Oct 1999 | A |
| 6004441 | Fujiwara et al. | Dec 1999 | A |
| 6033866 | Guo et al. | Mar 2000 | A |
| 6054039 | Shieh | Apr 2000 | A |
| 6063039 | Cunningham et al. | May 2000 | A |
| 6071391 | Gotoh et al. | Jun 2000 | A |
| 6103033 | Say et al. | Aug 2000 | A |
| 6120676 | Heller et al. | Sep 2000 | A |
| 6134461 | Say et al. | Oct 2000 | A |
| 6143164 | Heller et al. | Nov 2000 | A |
| 6174734 | Ito et al. | Jan 2001 | B1 |
| 6175752 | Say et al. | Jan 2001 | B1 |
| 6206835 | Spillman, Jr. et al. | Mar 2001 | B1 |
| 6262264 | Buck, Jr. et al. | Jul 2001 | B1 |
| 6284125 | Hodges et al. | Sep 2001 | B1 |
| 6294062 | Buck, Jr. et al. | Sep 2001 | B1 |
| 6294281 | Heller | Sep 2001 | B1 |
| 6338790 | Feldman et al. | Jan 2002 | B1 |
| 6352824 | Buck, Jr. et al. | Mar 2002 | B1 |
| 6399293 | Pachl et al. | Jun 2002 | B1 |
| 6436255 | Yamamoto et al. | Aug 2002 | B2 |
| 6442413 | Silver | Aug 2002 | B1 |
| 6500571 | Liberatore et al. | Dec 2002 | B2 |
| 6531239 | Heller | Mar 2003 | B2 |
| 6599407 | Taniike et al. | Jul 2003 | B2 |
| 6650471 | Doi | Nov 2003 | B2 |
| 6736957 | Forrow et al. | May 2004 | B1 |
| 6740215 | Nakaminami et al. | May 2004 | B1 |
| 6746582 | Heller et al. | Jun 2004 | B2 |
| 6766817 | da Silva | Jul 2004 | B2 |
| 6885196 | Taniike et al. | Apr 2005 | B2 |
| 6908535 | Rankin et al. | Jun 2005 | B2 |
| 6918404 | da Silva | Jul 2005 | B2 |
| 7018755 | Ikegami et al. | Mar 2006 | B2 |
| 7066586 | da Silva | Jun 2006 | B2 |
| 7125382 | Zhou et al. | Oct 2006 | B2 |
| 7171252 | Scarantino et al. | Jan 2007 | B1 |
| 7238442 | Heller | Jul 2007 | B2 |
| 7297112 | Zhou et al. | Nov 2007 | B2 |
| 7410709 | Ladisch et al. | Aug 2008 | B2 |
| 7485212 | Willner et al. | Feb 2009 | B2 |
| 7736309 | Miller et al. | Jun 2010 | B2 |
| 9668684 | Feldman | Jun 2017 | B2 |
| 10004441 | Feldman | Jun 2018 | B2 |
| 20020172992 | Heller | Nov 2002 | A1 |
| 20030148169 | Willner et al. | Aug 2003 | A1 |
| 20030152823 | Heller | Aug 2003 | A1 |
| 20040064133 | Miller et al. | Apr 2004 | A1 |
| 20040079653 | Karinka et al. | Apr 2004 | A1 |
| 20040122489 | Mazar et al. | Jun 2004 | A1 |
| 20040219664 | Heller et al. | Nov 2004 | A1 |
| 20040236200 | Say et al. | Nov 2004 | A1 |
| 20040245101 | Willner et al. | Dec 2004 | A1 |
| 20050261562 | Zhou et al. | Nov 2005 | A1 |
| 20050261563 | Zhou et al. | Nov 2005 | A1 |
| 20060003398 | Heller et al. | Jan 2006 | A1 |
| 20060201805 | Forrow et al. | Sep 2006 | A1 |
| 20060269826 | Katz et al. | Nov 2006 | A1 |
| 20060270923 | Brauker et al. | Nov 2006 | A1 |
| 20070038054 | Zhou et al. | Feb 2007 | A1 |
| 20070042377 | Gao et al. | Feb 2007 | A1 |
| 20070197890 | Boock et al. | Aug 2007 | A1 |
| 20070203407 | Hoss et al. | Aug 2007 | A1 |
| 20070232877 | He | Oct 2007 | A1 |
| 20080044721 | Heller et al. | Feb 2008 | A1 |
| 20080093230 | Diamond et al. | Apr 2008 | A1 |
| 20090000961 | Heller et al. | Jan 2009 | A1 |
| 20090018425 | Ouyang et al. | Jan 2009 | A1 |
| 20090299672 | Zhang | Dec 2009 | A1 |
| 20100213057 | Feldman | Aug 2010 | A1 |
| 20100213082 | Feldman et al. | Aug 2010 | A1 |
| 20100326842 | Mazza et al. | Dec 2010 | A1 |
| 20100331771 | Mazza et al. | Dec 2010 | A1 |
| 20160045147 | Ouyang | Feb 2016 | A1 |
| Number | Date | Country |
|---|---|---|
| 301808 | Nov 1917 | DE |
| 2903216 | Aug 1979 | DE |
| 227029 | Sep 1985 | DE |
| 4126692 | Feb 1993 | DE |
| 4314417 | Nov 1994 | DE |
| 4344646 | Jun 1995 | DE |
| 0048090 | Mar 1982 | EP |
| 0078636 | May 1983 | EP |
| 0096288 | Dec 1983 | EP |
| 0125139 | Nov 1984 | EP |
| 0136362 | Apr 1985 | EP |
| 0170375 | Feb 1986 | EP |
| 0080304 | May 1986 | EP |
| 0184909 | Jun 1986 | EP |
| 0206218 | Dec 1986 | EP |
| 0230472 | Aug 1987 | EP |
| 0241309 | Oct 1987 | EP |
| 0245073 | Nov 1987 | EP |
| 0278647 | Aug 1988 | EP |
| 0286084 | Oct 1988 | EP |
| 0359831 | Mar 1990 | EP |
| 0368209 | May 1990 | EP |
| 0390390 | Oct 1990 | EP |
| 0400918 | Dec 1990 | EP |
| 0453283 | Oct 1991 | EP |
| 0470290 | Feb 1992 | EP |
| 0255291 | Jun 1992 | EP |
| 0127958 | Apr 1996 | EP |
| 0745843 | Dec 1996 | EP |
| 0781406 | May 1998 | EP |
| 1394171 | May 1975 | GB |
| 2073891 | Oct 1981 | GB |
| 2154003 | Oct 1985 | GB |
| 2204408 | Nov 1988 | GB |
| 54-41191 | Apr 1979 | JP |
| 55-10581 | Jan 1980 | JP |
| 55-10583 | Jan 1980 | JP |
| 55-10584 | Jan 1980 | JP |
| 55-12406 | Jan 1980 | JP |
| 55-78242 | Jun 1980 | JP |
| 56-163447 | Dec 1981 | JP |
| 57-12359 | Jan 1982 | JP |
| 57-70448 | Apr 1982 | JP |
| 60-173457 | Sep 1985 | JP |
| 60-173458 | Sep 1985 | JP |
| 60-173459 | Sep 1985 | JP |
| 61-90050 | May 1986 | JP |
| 62-85855 | Apr 1987 | JP |
| 62-114747 | May 1987 | JP |
| 63-58149 | Mar 1988 | JP |
| 63-128252 | May 1988 | JP |
| 63-139246 | Jun 1988 | JP |
| 63-294799 | Dec 1988 | JP |
| 63-317758 | Dec 1988 | JP |
| 1-114746 | May 1989 | JP |
| 1-114747 | May 1989 | JP |
| 1-134244 | May 1989 | JP |
| 1-156658 | Jun 1989 | JP |
| 2-62958 | Mar 1990 | JP |
| 2-120655 | May 1990 | JP |
| 2-287145 | Nov 1990 | JP |
| 2-310457 | Dec 1990 | JP |
| 3-26956 | Feb 1991 | JP |
| 3-28752 | Feb 1991 | JP |
| 3-202764 | Sep 1991 | JP |
| 5-72171 | Mar 1993 | JP |
| 5-196595 | Aug 1993 | JP |
| 1281988 | Jan 1987 | SU |
| 198505119 | Nov 1985 | WO |
| 198906802 | Jul 1989 | WO |
| 198908713 | Sep 1989 | WO |
| 199005300 | May 1990 | WO |
| 199104704 | Apr 1991 | WO |
| 199109139 | Jun 1991 | WO |
| 199213271 | Aug 1992 | WO |
| 199420602 | Sep 1994 | WO |
| 199427140 | Nov 1994 | WO |
| 199502817 | Jan 1995 | WO |
| 199700441 | Jan 1997 | WO |
| 199718464 | May 1997 | WO |
| 199719344 | May 1997 | WO |
| 199742882 | Nov 1997 | WO |
| 199742883 | Nov 1997 | WO |
| 199742886 | Nov 1997 | WO |
| 199742888 | Nov 1997 | WO |
| 199743962 | Nov 1997 | WO |
| 199835225 | Aug 1998 | WO |
| 199908106 | Feb 1999 | WO |
| 199930152 | Jun 1999 | WO |
| 200206828 | Jan 2002 | WO |
| Entry |
|---|
| (Author Unknown), “Hemoglobin Function” Ch. 3, pp. 37-60 (Date Unknown). |
| Abruna, H. D., et al., “Rectifying Interfaces Using Two-Layer Films of Electrochemically Polymerized Vinylpyridine and Vinylbipyridine Complexes of Ruthenium and Iron on Electrodes”, Journal of the American Chemical Society. vol. 103. No. 1, 1981, pp. 1-5. |
| Albery, W. J., et al., “Amperometric Enzyme Electrodes Part II: Conducting Salts as Electrode Materials for the Oxidation of Glucose Oxidase”, Journal of Electroanalytical Chemistry. vol. 194, 1985, pp. 223-235. |
| Albery, W. J., et al., “Amperometric Enzyme Electrodes”, Philosophical Transactions of The Royal Society of London, vol. 316, 1987, pp. 107-119. |
| Alcock, S. J., et al., “Continuous Analyte Monitoring to Aid Clinical Practice”, IEEE Engineering in Medicine and Biology Magazine, 1994, pp. 319-325. |
| Alkire, et al., “Current Distribution in a Tubular Electrode under Liminar Flow: One Electrode Reaction,” J. Electrochem. Soc.: Electrochemical Science and Technology, 1977 vol. 124, No. 7, pp. 1043-1049. |
| Anderson, L. B., et al., “Thin-Layer Electrochemistry: Steady-State Methods of Studying Rate Processes”, Journal of Electroanalytical Chemistry. vol. 10, 1965, pp. 295-305. |
| Aoki et al., “Effect of Quaternization on Electron Diffusion Coefficients for Redox Hyrdrogels Based on Poly(4-vinylpyridine),” The Journal of Physical Chemistry, 1995, vol. 99, No. 14, pp. 5012-5110. |
| Aoki et al., “Electron Diffusion Coefficients in Hydrogels Fonned of Cross-Linked Redox Polymers,” The Journal of Physical Chemistry, 1993, vol. 97, pp. 11014-11019. |
| Bartlett, P. N., et al., “Covalent Binding of Electron Relays to Glucose Oxidase”, Journal of the Chemical Society. Chemical Communications, 1987, pp. 1603-1604. |
| Bartlett, P. N., et al., “Modification of Glucose Oxidase by Tetrathiafulvalene”, Journal of the Chemical Society. Chemical Communications, 1990, pp. 1135-1136. |
| Bartlett, P. N., et al., “Strategies for the Development of Amperometric Enzyme Electrodes”, Biosensors. vol. 3, 1987/88, pp. 359-379. |
| Barton et al., “Electroreduction of 02 to Water on the ‘Wired’ Laccase Cathode,” J. Phys. Chem. B 2001, vol. 105, pp. 11917-11921. |
| Barton et al., “Electroreduction of 02 to Water at 0.6 V (SHE) at pH 7 on the ‘Wired’ Pleurotus Ostreatus Laccase Cathode,” Biosensors and Bioelectronics, 2002, vol. 17, pp. 1071-1074. |
| Barton et al., “The ‘Wired’ Laccase Cathode: High Current Density Electroreduction of 02 to Water at +0.7 V (NHE) at pH 5,” J. Am. Chem. Soc. 2001, vol. 123, pp. 5802-5803. |
| Bier, et al., “An Enzymatic Amplification Cycle for High Sensitive Immunoassay”, Anal. Chim, Acta, 1996, vol. 328, No. 1, pp. 27-32. |
| Bier, et al., “High Sensitive Competitive Immunodetection of 2, 4-Dichlorophenoxyacetic Acid Using Enzymatic Amplification with Electrochemical Detection”, J. Anal. Chem., 1996, vol. 354, No. 7-8, pp. 861-865. |
| Binyamin et al., “Mechanical and Electrochemical Characteristics of Composites of Wired Glucose Oxidase and Hydrophilic Graphite,” Journal of the Electrochemical Society. 2000, vol. 147, No. 7, pp. 2780-2783. |
| Binyamin et al., “Stabilization of Wired Glucose Oxidase Anodes Rotating at 1000 rpm at 37° C.,” Journal of the Electrochemical Society, 1999 vol. 146, No. 8, pp. 2965-2967. |
| Blauch et al., “Effects of Long-Range Electron Transfer on Charge Transport in Static Assemblies of Redox Centers,” The Journal of Physical Chemistry. 1993, vol. 97, No. 24, pp. 6444-6448. |
| Bobbioni-Harsch, E., et al., “Lifespan of Subcutaneous Glucose Sensors and Their Performances During Dynamic Glycaemia Changes in Rats”, Journal of Biomedical Engineering. vol. 15, 1993, pp. 457-463. |
| Brandt, J., et al., “Covalent Attachment of Proteins to Polysaccharide Carriers by Means of Benzoquinone”, Biochimica et Biophysica Acta. vol. 386, 1975, pp. 196-202. |
| Brownlee, M., et al., “A Glucose-Controlled Insulin-Delivery System: Semisynthetic Insulin Bound to Lectin”, Science. vol. 206, 1979, 1190-1191. |
| Cass, A. E., et al., “Ferricinum Ton as an Electron Acceptor for Oxido-Reductases”, Journal of Electroanalytical Chemistry. vol. 190, 1985, pp. 117-127. |
| Cass, A. E., et al., “Ferrocene-Medicated Enzyme Electrode for Amperometric Determination of Glucose”, Analytical Chemistry. vol. 56. No. 4, 1984, 667-671. |
| Castner, J. F., et al., “Mass Transport and Reaction Kinetic Parameters Determined Electrochemically for Immobilized Glucose Oxidase”, Biochemistry. vol. 23 No. 10, 1984, 2203-2210. |
| Chen et al., “A Miniature Biofuel Cell,” Journal of the American Chemical Society, 2001, vol. 123, No. 35, pp. 8630-8631. |
| Chen et al., “In Situ Assembled Mass-Transport Controlling Micromembranes and Their Application in Implanted Amperometric Glucose Sensors,” Analytical Chemistry. Aug. 15, 2000, vol. 72, No. 16, pp. 3757-3763. |
| Chen, C.Y. et al., “A Biocompatible Needle-Type Glucose Sensor Based on Platinum Electroplated Carbon Electrode”, Applied Biochemistry and Biotechnology, 36:211-226 (1992). |
| Chen, C.Y. et al., “Amperometric Needle-Type Glucose Sensor based on a Modified Platinum Electrode with Diminished Response to Interfering Materials”, Analytica Chimica Acta, 265:5-14 (1992). |
| Claremont, D. J., et al., “Biosensors for Continuous In Vivo Glucose Monitoring”, Proceedings of the Annual International Conference of the IEEE Engineering in Medicine and Biology Society. vol. 10, 1988. |
| Clark Jr., L. C., et al., “Differential Anodic Enzyme Polarography for the Measurement of Glucose”, Oxygen Transport to Tissue: Instrumentation, Methods, and Physiology, 1973, pp. 127-133. |
| Clark Jr., L. C., et al., “Electrode Systems for Continuous Monitoring in Cardiovascular Surgery”, Annals New York Academy of Sciences, 1962, pp. 29-45. |
| Clarke, W. L., et al., “Evaluating Clinical Accuracy of Systems for Self-Monitoring of Blood Glucose”, Diabetes Care. vol. 10. No. 5, 1987, pp. 622-628. |
| Colon et al., “Cobalt Polypyridyl Complexes as Redox Mediators for Lipoamide Dehydrogenase,” Electroanalysis, 1998, vol. 10, No. 9, pp. 621-627. |
| Cook “Enzyme Mechanism for Isotope Effects” CRC Press, 1991, pp. 351-352. |
| Csoregi E., et al., “On-Line Glucose Monitoring by Using Microdialysis Sampling and amperometric Detection Based on ‘Wired’ Glucose Oxidase in Carbon Paste”, Mikrochimica Acta. vol. 121, 1995, pp. 31-40. |
| Csoregi, E., et al., “Design and Optimization of a Selective Subcutaneously Implantable Glucose Electrode Based on ‘Wired’ Glucose Oxidase”, Analytical Chemistry. vol. 67. No. 7, 1995, pp. 1240-1244. |
| Csoregi, E., et al., “Design, Characterization, and One-Point in Vivo Calibration of a Subcutaneously Implanted Glucose Electrode”, Analytical Chemistry. vol. 66 No. 19, 1994, pp. 3131-3138. |
| Davis, G., “Electrochemical Techniques for the Development of Amperometric Biosensors”, Biosensors. vol. 1, 1985, pp. 161-178. |
| De Lumley-Woodyear et al., “Polyacrylamide-Based Redox Polymer for Connecting Redox Centers of Enzymes to Electrodes,” Analytical Chemistry, 1995, vol. 67, No. 8, pp. 1332-1338. |
| Degani, Y., et al., “Direct Electrical Communication Between Chemically Modified Enzymes and Metal Electrodes. 1. Electron Transfer from Glucose Oxidase to Metal Electrodes via Electron Relays, Bound Covalently to the Enzyme”, The Journal of Physical Chemistry. vol. 91. No. 6, 1987, pp. 1285-1289. |
| Degani, Y., et al., “Direct Electrical Communication Between Chemically Modified Enzymes and Metal Electrodes. 2. Methods for Bonding Electron-Transfer Relays to Glucose Oxidase and D-Amino-Acid Oxidase”, Journal of the American Chemical Society. vol. 110. No. 8, 1988, pp. 2615-2620. |
| Degani, Y., et al., “Electrical Communication Between Redox Centers of Glucose Oxidase and Electrodes via Electrostatically and Covalently Bound Redox Polymers”, Journal of the American Chemical Society. vol. 111, 1989, pp. 2357-2358. |
| Denisevich, P., et al., “Unidirectional Current Flow and Charge State Trapping at Redox Polymer Interfaces on Bilayer Electrodes: Principles, Experimental Demonstration, and Theory”, Journal of the American Chemical Society. vol. 103, 1981, pp. 4727-4737. |
| Dicks, J. M., et al., “Ferrocene Modified Polypyrrole with Immobilised Glucose Oxidase and its Application in Amperometric Glucose Microbiosensors”, Annales de Bioloie Clinique. vol. 47, 1989, pp. 607-619. |
| Ellis, C. D., et al., “Selectivity and Directed Charge Transfer through an Electroactive Metallopolymer Film”, Journal of the American Chemical Society. vol. 103. No. 25, 1981, pp. 7480-7483. |
| Engstrom, R. C., “Electrochemical Pretreatment of Glassy Carbon Electrodes”, Analytical Chemistry. vol. 54.No. 13, 1982, pp. 2310-2314. |
| Engstrom, R. C., et al., “Characterization of Electrochemically Pretreated Glassy Carbon Electrodes”, Analytical Chemistry. vol. 56. No. 2, 1984, pp. 136-141. |
| Fischer, H., et al., “Intramolecular Electron Transfer Medicated by 4,4-Bypyridine and Related Bridging Groups”, Journal of the American Chemical Society. vol. 98. No. 18, 1976, pp. 5512-5517. |
| Kakehi N. et al.: “A Novel Wireless Glucose Sensor Employing Direct Electron Transfer Principle Based Enzyme Fuel Cell”, Biosensors and Bioelectronics, vol. 22, Jan. 1, 2007, pp. 2250-2255. |
| Extended European Search Report issued in EP Patent Application No. 17183254.6 dated Oct. 19, 2017. |
| The Potential Use of Radio Frequency Identification Devices for Active Monitoring of Blood Glucose Levels. Moore, Bert. Journal of Diabetes Science and Technology. vol. 3, Issue 1, Jan. 2009. |
| Mastrototaro, J. J., et al., “An Electroenzymatic Glucose Sensor Fabricated on a Flexible Substrate”, Sensors and Actuators B. vol. 5, 1991, pp. 139-144. |
| Matthews, DR., et al., “An Amperometric Needle-Type Glucose Sensor Tested in Rats and Man”, Original Articles, pp. 248-252 (1988). |
| McKean, B. D., et al., “A Telemetry-Instrumentation System for Chronically Implanted Glucose and Oxygen Sensors”, IEEE Transactions on Biomedical Engineering. vol. 35. No. 7, 1988, pp. 526-532. |
| McNeil, C. J., et al., “Thermostable Reduced Nicotinamide Adenine Dinucleotide Oxidase: Application to Amperometric Enzyme Assay”, Analytical Chemistry. vol. 61. No. 1, 1989, pp. 25-29. |
| Miyawaki, 0., et al., “Electrochemical and Glucose Oxidase Coenzyme Activity of Flavin Adenine Dinucleotide Covalently Attached to Glassy Carbon at the Adenine Amino Group”, Biochimica et Biophysica Acta. vol. 838, 1985, pp. 60-68. |
| Moatti-Sirat, D., et al., “Evaluating In Vitro and In Vivo the Interference of Ascorbate and Acetaminophen on Glucose Detection by a Needle-Type Glucose Sensor”, Biosensors & Bioelectronics. vol. 7, 1992, pp. 345-352. |
| Moatti-Sirat, D., et al., “Reduction of Acetaminophen Interference in Glucose Sensors by a Composite Nation Membrane: Demonstration in Rats and Man”, Diabetologia. vol. 37, 1994, pp. 610-616. |
| Moatti-Sirat, D., et al., “Towards Continuous Glucose Monitoring: In Vivo Evaluation of a Miniaturized Glucose Sensor Implanted for Several Days in Rat Subcutaneous Tissue”, Diabetologia. vol. 35, 1992, pp. 224-330. |
| Moore, “The Potential Use of Radio Frequency Identification Devices for Active Monitoring of Blood Glucose Levels.” Journal of Diabetes Science and Technology, vol. 3, Issue 1, pp. 180-183 (Jan. 2009). |
| Moser, I. et al., “Advanced Immobilization and Protein Techniques on thin Film Biosensors”, Sensors and Actuators, 87:356-362 (1992). |
| Moussy, F. et al., “Performance of Subcutaneously Implanted Needle-Type Glucose Sensors Employing a Novel Trilayer Coating”, Anal. Chem., 65:2072-2077 (1993). |
| Nagy, G., et al., “A New Type of Enzyme Electrode: The Ascorbic Acid Eliminator Electrode”, Life Sciences. vol. 31. No. 23, 1982, pp. 2611-2616. |
| Nakamura, S., et al., “Effect of Periodate Oxidation on the Structure and Properties of Glucose Oxidase”, Biochimica et Biophysica Acta. vol. 445, 1976, pp. 294-308. |
| Narasimhan, K., et al., “p-Benzoquinone Activation of Metal Oxide Electrodes for Attachment of Enzymes”, Enzyme and Microbial Technology. vol. 7, 1985, pp. 283-286. |
| Notification of Transmittal of International Preliminary Examination Report, dated Jul. 19, 2002, In International Application No. PCT/US02/15706 of TheraSense, Inc. |
| Notification of Transmittal of the International Search Report or the Declaration, dated May 11, 2004, in International Application No. PCT/US03/13806 of TheraSense, Inc. |
| Notification of Transmittal of International Preliminary Examination Report, dated Oct. 7, 2004, in International Application No. PCT/US03/13806 of TheraSense, Inc. |
| Ohara et al., “Wired' Enzyme Electrodes for Amperometric Determination of Glucose or Lactate in the Presence of Interfering Substance,” Analytical Chemistry, Aug. 1, 1994, vol. 66, No. 15, pp. 2451-2457. |
| Ohara, T. J., “Osmium Bipyridyl Redox Polymers Used in Enzyme Electrodes”, Platinum Metals Review, vol. 39. No. 2, 1995, pp. 54-62. |
| Ohara, T. J., et al., “Glucose Electrodes Based on Cross-Linked [0s(bpy)2C1]+2+ Complexed Poly(1-Vinylimidazole) Films”, Analytical Chemistry. vol. 65. No. 23, 1993, pp. 3512-3517. |
| Olievier, C. N., et al., “In Vivo Measurement of Carbon Dioxide Tension with a Miniature Electrode”, Pflugers Archiv: European Journal of Physiology. vol. 373, 1978, pp. 269-272. |
| Paddock, R. M., et al., “Electrocatalytic Reduction of Hydrogen Peroxide via Direct Electron Transfer From Pyrolytic Graphite Electrodes to Irreversibly Adsorbed Cyctochrome C Peroxidase”, Journal of Electroanalytical Chemistry. vol. 260, 1989, pp. 487-494. |
| Palleschi, G. et al., “Ideal Hydrogen Peroxide-Based Glucose Sensor”, Applied Biochemistry and Biotechnology, 31:21-35 (1991). |
| Palleschi, G., et al., “A Study of Interferences in Glucose Measurements in Blood by Hydrogen Peroxide Based Glucose Probes”, Analytical Biochemistry. vol. 159, 1986, pp. 114-121. |
| Palmore et al., “A Methanol/Dioxygen Biofuel Cell that uses NAD+-Dependent Dehydrogenases as Catalysts: Application of an Electro-Enzymatic Method to Regenerate Nicotinamide Adenine Dinucleotide at Low Overpotentials,” Journal of Electroanalytical Chemistry, 1998, vol. 443, pp. 155-161. |
| Palmore et al., “Electro-Enzymatic Reduction of Dioxygen to Water in the Cathode Compartment of a Biofuel Cell,” Journal of Electroanalytical Chemistry, 1999, vol. 464, pp. 110-117. |
| Palmore et al., “Microbial and Enzymatic Biofuel Cells,” Enzymatic Conversion of Biomass for Fuels Production, 1994, Chapter 14, pp. 271-290. |
| Pankratov, I., et al., “Sol-Gel Derived Renewable-Surface Biosensors”, Joumal of Electroanalytical Chemistry. vol. 393, 1995, pp. 35-41. |
| Pathak, C., et al., “Rapid Photopolymerization of Immunoprotective Gels in Contact with Cells and Tissue”, Journal of the American Chemical Society. vol. 114. No. 21, 1992, pp. 8311-8312. |
| Pickup, J., et al., “Potentially-Implantable, Amperometric Glucose Sensors with Mediated Electron Transfer: Improving the Operating Stability”, Biosensors. vol. 4, 1989, pp. 109-119. |
| Pishko, M. V., et al., “Amperometric Glucose Microelectrodes Prepared Through Immobilization of Glucose Oxidase in Redox Hydrogels”, Analytical Chemistry. vol. 63. No. 20, 1991, pp. 2268-2272. |
| Poitout, V., et al., “A Glucose Monitoring System for On Line Estimation in Man of Blood Glucose Concentration Using a Miniaturized Glucose Sensor Implanted in the Subcutaneous Tissue and a Wearable Control Unit”, Diabetolgia. vol. 36, 1993, pp. 658-663. |
| Poitout, V., et al., “Calibration in Dogs of a Subcutaneous Miniaturized Glucose Sensor Using a Glucose Meter for Blood Glucose Determination”, Biosensors & Bioelectronics. vol. 7, 1992, pp. 587-592. |
| Poitout, V., et al., “In Vitro and In Vivo Evaluation in Dogs of a Miniaturized Glucose Sensor”, ASAIO Transactions. vol. 37. No. 3, 1991, pp. M298-M300. |
| Pollak, A., et al., “Enzyme Immobilization by Condensation Copolymerization into Cross-Linked Polyacrylamide Gels”, Journal of the American Chemical Society. vol. 102. No. 20, 1980, pp. 6324-6336. |
| Pons, B. S. et al., Application of Deposited Thin Metal Films as Optically Transparent Electrodes for Internal Reflection Spectometric Observation of Electrode Solution Interfaces, Analytical Chemistry, 39(6):685-688, (May 1967). |
| Quinn et al., “Biocompatible, Glucose-Permeable Hydrogel for in situ Coating of Implantable Biosensors,” Biomaterials, 1997, vol. 18, No. 23, pp. 1665-1670. |
| Quinn et al., “Photo-Crosslinked Copolymers of 2-Hydroxyethyl Methacrylate, Poly(ethylene Glycol) Tetra-Acrylate and Ethylene Dimethacrylate for Improving Biocompatibility of Biosensors,” Biomaterials. 1995. vol. 15. No. 5. pp. 389-396. |
| Rajagopalan et al., “Effect of Quaternization of the Glucose Oxidase ‘Wiring’ Redox Polymer on the Maximum Current Densities of Glucose Electrodes,” The Journal of Physical Chemistry, 1996, vol. 100, No. 9, pp. 3719-3727. |
| Rajagopalan et al., “Electrical ‘Wiring’ of Glucose Oxidase in Electron Conducting Hyrogels,” Molecular Electronics, Chapter 7, pp. 241-254. |
| Rao et al., “Metal-Oxygen and Glucose-Oxygen Cells for Implantable Devices,” Biomedical Engineering, 1974, vol. 9, No. 3, pp. 98-102. |
| Reach, G. et al., “A Method for Evaluating in vivo the Functional Characteristics of Glucose Sensors”, Biosensors 2:211-220 (1986). |
| Reach, G., et al., “Can Continuous Glucose Monitoring Be Used for the Treatment of Diabetes?”, Analytical Chemistry. vol. 64. No. 6, 1992, pp. 381-386. |
| Rebrin, K., et al., “Automated Feedback Control of Subcutaneous Glucose Concentration in Diabetic Dogs”, Diabetologia. vol. 32, 1989, pp. 573-576. |
| Sasso, S. V., et al., “Electropolymerized 1,2-Diaminobenzene as a Means to Prevent Interferences and Fouling and to Stabilize Immobilized Enzyme in Electrochemical Biosensors”, Analytical Chemistry. vol. 62. No. 11, 1990, pp. 1111-1117. |
| Sayka et al., “The Effect of Plasma Treatment on the Wettabiity of Substrate Materials,” Solid State Technology, 1989, vol. 32, No. 5, pp. 69-70. |
| Schalkhammer, T. et al, “Electrochemical Glucose Sensors on Permselective Non-conducting Substituted Pyrrole Polymers”, Sensors and Actuators, B4:273-281 (1991). |
| Scheller, F., et al., “Enzyme Electrodes and Their Application”, Philosophical Transactions of the Royal Society of London B. vol. 316, 1987, pp. 85-94. |
| Service, “Can Chip Devices Keep Shrinking?” Science, Dec. 13, 1996, vol. 274, pp. 1834-1836. |
| Shichiri, M., et al., “Glycaemic Control in Pancreatectomized Dogs with a Wearable Artificial Endocrine Pancreas”, Diabetologia. vol. 24, 1983, pp. 179-184. |
| Foulds, N. C., et al., “Enzyme Entrapment in Electrically Conducting Polymers. Immobilisation of Glucose Oxidase in Polypyrrole and its Application in Amperometric Glucose Sensors”, Journal of the Chemical Society. Faraday Transactions 1. vol. 82, 1986, pp. 1259-1264. |
| Foulds, N. C., et al , “Immobilization of Glucose Oxidase in Ferrocene-Modified Pyrrole Polymers”, Analytical Chemistry. vol. 60. No. 22, 1988, pp. 2473-2478. |
| Fresht “Stucture and Mechanism in Protein Science: A Guide to Enzyme Catalysis and Protein Folding” WH Freeman, 1999, pp. 473-474. |
| Frew, J. E., et al., “Electron-Transfer Biosensors”, Philosophical Transactions of the Royal Society of London. vol. 316, 1987, pp. 95-106. |
| Frew, J.E., et al. “Electrochemical Biosensors” Analytical Chemistry, Aug. 1, 1987, vol. 59, pp. 933A-944A. |
| Gernet, S. et al., “Fabrication and Characterization of a Planar Electrochemical Cell and its Application as a Glucose Sensor”, Biosensors & Actuators, 18:59-70 (1989). |
| Gorton, L., et al., “Selective Detection in Flow Analysis Based on the Combination of Immobilized Enzymes and Chemically Modified Electrodes”, Analytica Chimica Acta. vol. 250, 1991, pp. 203-248. |
| Gorton, L. “Carbon Paste Electrodes Modified with Enzymes, Tissues, and Cells” Electroanalysis, Jan. 1995, vol. 7, pp. 23-45. |
| Greenfield et al., “Inactivation of Immobilized Glucose Oxidase by Hydrogen Peroxide,” Analytical Biochemistry. 1975, vol. 65, pp. 109-124. |
| Gregg, B. A., et al., “Cross-Linked Redox Gels Containing Glucose Oxidase for Amperometric Bionsensor Applications”, Analytical Chemistry. vol. 62. No. 3, 1990, pp. 258-263. |
| Gregg, B. A., et al., “Redox Polymer Films Containing Enzymes. 1. A Redox Conducting Epoxy Cement: Synthesis, Characterization, and Electrocatalytic Oxidation of Hydroquinone”, Journal of Physical Chemistry. vol. 95. No. 15, 1991, 5970-5975. |
| Hale, P. D., et al., “A New Class of Amperometric Biosensor Incorporating a Polymeric Electron-Transfer Mediator”, Journal of the American Chemical Society. vol. 111. No. 9, 1989, pp. 3482-3484. |
| Harrison, D. J., et al., “Characterization of Perfluorosulfonic Acid Polymer Coated Enzyme Electrodes and a Miniatureized Integrated Potentiostat for Glucose Analysis in Whole Blood”, Analytical Chemistry. vol. 60. No. 19, 1988, pp. 2002-2007. |
| Hawkridge, E M., et al., “Indirect Coulometric Titration of Biological Electron Transport Components”, Analytical Chemistry. vol. 45. No. 7, 1973, pp. 1021-1027. |
| Heineman, W.R. “Spectro-electro-chemistry”, Analytical Chemistry, 50(3):390-392, 394, 396, 398, 400, 402 (Mar. 1978). |
| Heineman, W.R. et al., “Measurement of Enzyme E°' Values by Optically Transparent Thin Layer Electrochemical Cells”, Analytical Chemistry, 47(1):79, 82-84 (Jan. 1975). |
| Heller, “Electrical Connection of Enzyme Redox Centers to Electrodes,” The Journal of Physical Chemistry, 1992, vol. 96, No. 9, pp. 3579-3587. |
| Heller, A., “Electrical Wiring of Redox Enzymes”, Accounts of Chemical Research vol. 23. No. 5, 1990, 128-134. |
| Heller, A., et al., “Amperometric Biosensors Based on Three-Dimensional Hydrogel Forming Epoxy Networks”, Sensors and Actuators B. vol. 13-14, 1993, pp. 180-183. |
| Ianniello, R. M., et al., “Differential Pulse Voltammetric Study of Direct Electron Transfer in Glucose Oxidase Chemically Modified Graphite Electrodes”, Analytical Chemistry. vol. 54. No. 7, 1982, pp. 1098-1101. |
| Ianniello, R. M., et al , “Immobilized Enzyme Chemically Modified Electrode as an Amperometric Sensor”, Analytical Chemistry. vol. 53. No. 13, 1981, pp. 2090-2095. |
| Ikeda, T., et al., “Glucose Oxidase-Immobilized Benzoquinone-Carbon Paste Electrode as a Glucose Sensor”, Agricultural and Biological Chemistry. vol. 49. No. 2, 1985, pp. 541-543. |
| Jaremko et al., “Advances Toward the Implantable Artificial Pancreas for Treatment of Diabetes,” Diabetes Care, vol. 21, No. 3, Mar. 1998, pp. 444-450. |
| Jin et al., “Electron Transfer Between Cytochrome C and Copper Enzymes,” Biochemistry and Bioenergetics, 1996, vol. 39, pp. 221-225. |
| Johnson, J. M., et al., “Potential-Dependent Enzymatic Activity in an Enzyme Thin Layer Cell”, Analytical Chemistry. vol. 54. No. 8, 1982, pp. 1377-1383. |
| Johnson, K W., “Reproducible Electrodeposition of Biomolecules for the Fabrication of Miniature Electroenzymatic Biosensors”, Sensors and Actuators B. vol. 5, 1991, pp. 85-89. |
| Johnson, K W., et al., “In vivo Evaluation of an Electroenzymatic Glucose Sensor Implanted in Subcutaneous Tissue”, Biosensors & Bioelectronics. vol. 7, 1992, pp. 709-714. |
| Jonsson, G., et al., “An Amperometric Glucose Sensor Made by Modification of a Graphite Electrode Surface With Immobilized Glucose Oxidase and Adsorbed Mediator”, Biosensors. vol. 1, 1985, pp. 355-368. |
| Josowicz, M., et al., “Electrochemical Pretreatment of Thin Film Platinum Electrodes”, Journal of the Electrochemical Society. vol. 135 No. 1, 1988, pp. 112-115. |
| Katakis et al., “Electrostatic Control of the Electron Transfer Enabling Binding of Recombinant Glucose Oxidase and Redox Polyelectrolytes,” Journal of the American Chemical Society, 1994, vol. 116, No. 8, on. 3617-3618. |
| Katakis, I., et al., “L-a-Glycerophosphate and L-Lactate Electrodes Based on the Electrochemical ‘Wiring’ of Oxidases”, Analytical Chemistry. vol. 64. No. 9, 1992, pp. 1008-1013. |
| Katz et al., “A Biofuel Cell Based on Two Immiscible Solvents and Glucose Oxidase and Microperoxidase-11 Monolayer-Functionalized Electrodes,” New J. Chem., 1999, pp. 481-487. |
| Katz, et al. “A Non-Compartmentalized Glucose I 02 Biofuel Cell by Bioengineered Electrode Surfaces,” Journal of Electronanalytical Chemistry, 1999, vol. 479, pp. 64-68. |
| Kenausis et al., “‘Wiring’ of Glucose Oxidase and Lactate Oxidase Within a Hydrogel Made with Poly(vinyl pyridine) Complexed with [0s (4,4 ′-dimethoxy-2,2 ′bipyridine) 2C1]+ i2 +” The Journal of the Chemical Society. Faraday Transactions. 1996. vol. 92. No. 20, pp. 4131-4136. |
| Kondo, T. et al., “A Miniature Glucose Sensor, Implantable in the Blood Stream” , Diabetes Care, 5(3):218-221 (May-Jun. 1982). |
| Koroljova-Skorobogat'ko et al., “Purification and Characterization of the Constitutive Form of Laccase from the Basidiomycete Coriolus 1-lirsutus and Effect of Inducers on Laccase Synthesis,” Biotechnol Appi Biochem. 1998, vol. 28, pp. 47-54. |
| Kulys, J. et al., “Mediatorless Peroxidase Electrode and Preparation of Bienzyme Sensors”, Bioelectrochemistry and Bioenergetics. vol. 24, 1990, pp. 305-311. |
| Lager, W., et al., “Implantable Electrocatalytic Glucose Sensor”, Hormone Metabolic Research, vol. 26, 1994, pp. 526-530. |
| Lee et al., “Catalysis of the Reduction of Dioxygen at Graphite Electrodes Coated with Fungal Laccase A,” J. Electroanal. Chem., 1984, vol. 172, pp. 289-300. |
| Lee, J. et al., “A New Glucose Sensor using Microporous Enzyme Membrane”, Sensors and Actuators, B3:215-219 (1991). |
| Lewandowski, 3.3. et al., “Evaluation of a Miniature Blood Glucose Sensor”, Trans Am Soc Artif Intern Organs, XXXIV: 255-258 (1988). |
| Lindner, E., et al., “Flexible (Kapton-Based) Microsensor Arrays of High Stability for Cardiovascular Applications”, Journal of the Chemical Society. Faraday Transactions. vol. 89. No. 2, 1993, pp. 361-367. |
| Lisdat, et al., “Laccase-Modified Thick Film Electrodes”, (Germany) Technical Digest of The 7th International Meeting on Chemical Sensors, 1998, pp. 786-788. |
| Lortz, J., et al., “What is Bluetooth? We Explain the Newest Short-Range Connectivity Technology”, Smart Computing Learning Series. Wireless Computing. vol. 8. Issue 5, 2002, pp. 72-74. |
| Maidan, R., et al., “Elimination of Electrooxidizable Interferant-Produced Currents in Amperometric Biosensors”, Analytical Chemistry. vol. 64. No. 23, 1992, pp. 2889-2896. |
| Mann-Buxbaum, E. et al, “New Microminiaturized Glucose Sensors Using Covalent Immobilization Techniques”, Sensors and Actuators, B1:518-522 (1990). |
| Mano et al. A Miniature Biofuel Cell Operating at 0.78 V, Chem Commun. 2003, pp. 518-519. |
| Mano et al., “An Oxygen Cathode Operating in a Physiological Solution,” J. Am. Chem. Soc. 2002,124, 6480-6486. |
| Mano et al., “A Miniature Biofuel Cell Operating in a Physiological Buffer,” J. Am. Chem Soc. 2002, vol. 124, pp. 12962-12963. |
| Mano et al., “On the Relationship Between the Characteristics of Bilirubin Oxidases and 02 Cathodes Based on Their Wiring',” J Phys Chem B. 2002, vol. 106, pp. 8842-8848. |
| Shigeru, T. et al, “Simultaneous Determination of Glucose and 1,5-= Anydroglucitol”, ChemicalAbstracts, 111:394 (1989). |
| Sittampalam, G., et al., “Surface-Modified Electrochemical Detector for Liquid Chromatography”, Analytical Chemistry. vol. 55. No. 9, 1983, pp. 1608-1610. |
| Skoog, D. A., et al., “Evaluation of Analytical Data,” Fundamentals of Analytical Chemistry, 1966, pp. 55. |
| Soegijoko, S., et al., “External Artificial Pancreas: A New Control Unit Using Microprocessor”, Hormone and Metabolic Research Supplement Series. vol. 12, 1982, pp. 165-169. |
| Sprules, S.D., et al., “Evaluation of a New Disposable Screen-Printed Sensor Ship for the Measurement of NADH and its Modification to Produce a Lactate Biosensor Employing Microliter Volumes”, Electroanalysis. vol. 8. No. 6, 1996, pp. 539-543. |
| Sternberg, F., et al., “Calibration Problems of Subcutaneous Glucosensors when Applied ‘In Situ’ in Man”, Hormone and Metabolic Research. vol. 26, 1994, pp. 523-526. |
| Sternberg, R., et al., “Covalent Enzyme Coupling on Cellulose Acetate Membranes for Glucose Sensor Development”, Analytical Chemistry. vol. 60. No. 24, 1988, pp. 2781-2786. |
| Stocklein et al., “Laccase: A Marker Enzyme for Solvent-Modified Immunoassays”, Ann. N.Y. Acad. Sci., 2006, vol. 799, pp. 525-528. |
| Stocklein et al., “Effects of Organic Solvents on Semicontinuous Immunochemical Detection of Coumarin Derivatives”, Sens. Actuators, vol. B 24(1-3), 1995, pp. 80-84. |
| Suekane, M., “Immobilization of Glucose Isomerase”, Zettschrift fiir Allemeine Mikrobioloie. vol. 22. No. 8, 1982, pp. 565-576. |
| Tarasevich et al., “Electrocatalysis of Cathodic Molecular Oxygen Reduction with Biopolymers-Enzymes and Their Models” J. Electroanal. Chem., 1986, vol. 206, pp. 217-227. |
| Tarasevich et al., “Electrocatalysis of a Cathodic Oxygen Reduction by Laccase” Bioelectrochemistry and Bioenergetics, 1979, vol. 6, pp. 393-403. |
| Tarasevich, M. R., “Bioelectrocatalysis”, Comprehensive Treatise of Electrochemistry. vol. 10, 1985, pp. 231-295. |
| Taylor, “Wiring' of Glucose Oxidase Within a Hydrogel Moade with Polyvinyl Imidazole Complexed with 10s-4,4′-dimethoxy-2,2′-bipyridine) C1J2,” Journal of Electroanalytical Chemistry, 1995, vol. 396, pp. 511-515. |
| Thuesen et al., “Cyclic Voltammetry and Electrocatalysis of the Blue Copper Oxidase Polyporus versicolor Laccase,” Ada Chemica Scandinavica, 1998, vol. 52, pp. 555-562. |
| Trojanowicz, M., et al., “Enzyme Entrapped Polypyrrole Modified Electrode for Flow Injection Determination of Glucose”, Biosensors & Bioelectronics. vol. 5, 1990, pp. 149-156. |
| Trudeau et al., “Reagentless Mediated Laccase Electrode for the Detection of Enzyme Modulators,” Analytical Chemistry, Mar. 1, 1997, vol. 69, No. 5, pp. 882-886. |
| Tsujimura et al., “Bioelectrocatalytic Reduction of Dioxygen to Water at Neutral pH Using Bilirubin Oxidase as an Enzyme and 2,2′-Azinobis (3-ethylbenzothiazolin-6 sulfonate) as an Electron Transfer Mediator.” Journal of Electroanalytical Chemistry. 2001, vol. 496, pp. 69-75. |
| Tsujimura et al., “Glucose/02 Biofuel Cell Operating at Physiological Conditions,” Electrochemistry. 2002, vol. 70, No. 12, pp. 940-942. |
| Turner, A., et al., “Diabetes Mellitus: Biosensors for Research and Management”, Biosensors. vol. 1, 1985, pp. 85-115. |
| Turner, R. F., et al., “A Biocompatible Enzyme Electrode for Continuous in vivo Glucose Monitoring in Whole Blood”, Sensors and Actuators B. vol. 1, 1990, pp. 561-564. |
| Umana, M., “Protein-Modified Electrochemically Active Biomaterial Surface”, U.S. Army Research Office. Analytical and Chemical Sciences Research Triangle Institute, 1988, pp. 1-9. |
| Urban, G., et al., “Miniaturized Thin-Film Biosensors Using Covalently Immobilized Glucose Oxidase”, Biosensors & Bioelectronics. vol. 6, 1991, pp. 555-562. |
| Velho, G., et al., “Strategies for Calibrating a Subcutaneous Glucose Sensor”, Biomedica Biochimica Acta. vol. 48, 1989, pp. 957-964. |
| Vidal, J.C. et al., “A chronoamperometric sensor for hydrogen peroxide based on electron transfer between immobilized horseradish peroxidase on a glassy carbon electrode and a diffusing ferrocene mediator”, Sensors and Actuators B 21, pp. 135-141 (1994). |
| Von Woedtke, T., et al., “In Situ Calibration of Implanted Electrochemical Glucose Sensors”, Biomedica Biochimica Acta. vol. 48, 1989, pp. 943-952. |
| Vreeke, M., et al., “Hydrogen Peroxide and13-Nicotinamide Adenine Dinucleotide Sensing Amperometric Electrodes Based on Electrical Connection of Horseradish Peroxidase Redox Centers to Electrodes through a Three-Dimensional Electron Relaying Polymer Network”, Analytical Chemistry. vol. 64. No. 24, 1992, pp. 3084-3090. |
| Vreeke, M. S., et al., “Hydrogen Peroxide Electrodes Based on Electrical Connection of Redox Centers of Various Peroxidases to Electrodes through a Three-Dimensional Electron Relaying Polymer Network”, Diagnostic Biosensors Polymers. Chapter 15, 1993, pp. 180-193. |
| Wagner et al., “Continuous Amperometric Monitoring of Glucose in a Brittle Diabetic Chimpanzee with a Miniature Subcutaneous Electrode,” Proc. Nati. Acad., May 1998, vol. 95 No. 11, pp. 6379-6382. |
| Wang, J., et al., “Activation of Glassy Carbon Electrodes by Alternating Current Electrochemical Treatment”, Analytica Chimica Acta. vol. 167, 1985, pp. 325-334. |
| Wang, J., et al., “Amperometric Biosensing of Organic Peroxides with Peroxidase Modified Electrodes”, Analytica Chimica Acta. vol. 254, 1991, pp. 81-88. |
| Wang, J., et al., “Screen-Printable Sol-Gel Enzyme-Containing Carbon Inks”, Analytical Chemistry. vol. 68. No. 15, 1996, pp. 2705-2708. |
| Wang, J., et al., “Sol-Gel-Derived Metal-Dispersed Carbon Composite Amperometric Biosensors”, Electroanalysis. vol. 9. No. 1, 1997, pp. 52-55. |
| Williams, D. L., et al., “Electrochemical-Enzymatic Analysis of Blood Glucose and Lactate”, Analytical Chemistry. vol. 42. No. 1, 1970, pp. 118-121. |
| Willner et al., “A Biofuel Cell Based on PyRroloquinoline Quinone and Microperoxidase-1 1 Monolayer-Functionalized Electrodes,” Bioelectrochemistry and Bioenergetics. 1998, vol. 44, pp. 209-214. |
| Willner et al., “Biofuel Cell Based on Glucose Oxidase and Microperoxidase-1 1 MonolayerFunctionalized Electrodes,” Journal of the Chemical Society—Perkin Transactions, 1998, vol. 2, No. 8, pp. 1817-1822. |
| Willner, et al. “Integration of Layered Redox Proteins and Condictive Supports for Bioelectronic Applications” Agnew. Chem. Int. Ed., 2000 vol. 39 pp. 1180-1218. |
| Yabuki, S., et al., “Electro-Conductive Enzyme Membrane”, Journal of the Chemical Society. Chemical Communications, 1989, pp. 945-946. |
| Yahiro et al., “Bioelectrochemistry I. Enzyme Utilizing Bio-Fuel Cell Studies,” Biochimica et Biophysica Acta, 1964, vol. 88, pp. 375-383. |
| Yang, L., et al., “Determination of Oxidase Enzyme Substrates Using Cross-Flow Thin Layer Amperometry”, Electroanalysis. vol. 8. No. 8-9, 1996, pp. 716-721. |
| Yao, S. J., et al., “The Interference of Ascorbate and Urea in Low-Potential Electrochemical Glucose Sensing”, Proceedings of the Twelfth Annual International Conference of the IEEE Engineering in Medicine and Biology Society. vol. 12. Part 2, 1990, pp. 487-489. |
| Yao, T., “A Chemically-Modified Enzyme Membrane Electrode as an Amperometric Glucose Sensor”, Analytica Chimica Acta. vol. 148, 1983, pp. 27-33. |
| Yaropolov et al., “Electrochemical Properties of Some Copper-Containing Oxidases,” Bioelectrochemistry and Bioenergetics, 1996, vol. 40, pp. 49-57. |
| Ye et al., “High Current Density ‘Wired’ Quinoprotein Glucose Dehydrogenase Electrode,” Analytical Chemistry. Feb. 1, 1993, vol. 65, No. 3, pp. 238-241. |
| Yildiz, A., et al., “Evaluation of an Improved Thin-Layer Electrode”, Analytical Chemistry. vol. 40. No. 7, 1968, pp. 1018-1024. |
| Zakeeruddin et al., “Towards Mediator Design: Characterization of Tris-(4,4 ′ substituted-2,2 ′-bipyridine) Complexes of Iron(II), Ruthenium(II) and Osmium(II) as Mediators for Glucose Oxidase of Aspergillus nizer and other Redox Proteins.” J. Electroanal. Chem. 1992. vol. 337. pp. 253-283. |
| Zamzow, K., et al., “New Wearable Continuous Blood Glucose Monitor (BGM) and Artificial Pancreas (AP)”, Diabetes. vol. 39, 1990, pp. 5A-20. |
| Katz et al. (1999) “A Non-Compartmentalized glucose | O2 Biofuel Cell by Bioengineere Electrode Surfaces” Journal of Electroanalytical Chemistry 479:64-68. |
| Song et al. “Electrocatalytic Oxygen Reduction Reaction, IPEM Full Cell Electrocatalysts and Catalyst Layers Fundamentals and Applications,” Zhang, Jiujum (Ed.) 1st Edition, XII, 1137 pages (2008). |
| Number | Date | Country | |
|---|---|---|---|
| 20180296139 A1 | Oct 2018 | US |
| Number | Date | Country | |
|---|---|---|---|
| Parent | 15351162 | Nov 2016 | US |
| Child | 16015882 | US | |
| Parent | 12717634 | Mar 2010 | US |
| Child | 15351162 | US | |
| Parent | 12393921 | Feb 2009 | US |
| Child | 12717634 | US |
