Personalized modeling of blood glucose concentration influenced by individualized sensor characteristics and individualized physiological characteristics

Abstract

A method for providing clinical data indicative of a concentration of a patient's blood analyte includes receiving a signal from a continuous analyte sensor located in the patient's interstitial fluid and independently modeling two or more factors that affect the signal, the factors resulting from individualized characteristics of the sensor and / or individualized physiological characteristics of the patient.

Description

[Technical field]

[0001] CROSS-REFERENCE TO RELATED APPLICATIONS This application claims the benefit of U.S. Provisional Patent Application No. 63 / 170,272, filed April 2, 2021, entitled "PERSONALIZED MODELING OF BLOOD GLUCOSE CONCENTRATION IMPACTED BY INDIVIDUALIZED SENSOR CHARACTERISTICS AND INDIVIDUALIZED PHYSIOLOGICAL CHARACTERISTICS," the contents of which are incorporated herein by reference. [Background technology]

[0002] Diabetes mellitus is a disease in which the pancreas cannot make enough insulin (Type I or insulin-dependent) and / or insulin is ineffective (Type II or non-insulin-dependent). In the diabetic state, the patient or user suffers from hyperglycemia, which can cause many physiological disorders associated with the deterioration of small blood vessels, such as kidney failure, skin ulcers, or bleeding into the vitreous of the eye. A hypoglycemic reaction (hypoglycemia) can be induced by inadvertent overdosing of insulin or after normal administration of insulin or glucose-lowering agents accompanied by abnormal exercise or inadequate food intake.

[0003] Traditionally, diabetics have carried self-monitoring blood glucose (SMBG) monitors, but these typically required an uncomfortable method such as pricking a finger. Due to lack of comfort and convenience, diabetics typically measure their glucose levels only two to four times per day. Unfortunately, such time intervals are so far apart that diabetics may not know about their hyperglycemic or hypoglycemic condition until it is too late, sometimes resulting in dangerous side effects. Not only may diabetics be slow to realize that they are in danger, but traditional methods may also increase the likelihood that they will not know that their blood glucose concentration values ​​are rising (high) or falling (low). Thus, diabetics may not be able to make educated insulin treatment decisions.

[0004] Another device that some diabetics use to monitor blood glucose is a continuous analyte sensor, e.g., a continuous glucose monitor (CGM) system. A CGM system typically includes a sensor that is placed invasively, minimally invasively, or non-invasively. The sensor measures the concentration of a given analyte, e.g., glucose, in the body and generates a raw signal using electronics associated with the sensor. The raw signal is converted to an output value that is rendered on a display. The output value resulting from the conversion of the raw signal is typically represented in a form that provides meaningful information to the user and is expressed in a format that the user is accustomed to analyzing, such as blood glucose expressed in mg / dL.

[0005] Some CGM systems rely on blood glucose (BG) fingerstick values ​​to correlate the sensor signal to clinical blood glucose, while others do not require real-time BG fingerstick values ​​(e.g., rely instead on factory information) to correlate (calibrate / convert) the raw signal from the sensor to a clinical blood glucose equivalent value that represents the glucose concentration in the patient. Both types of systems can suffer from inaccuracies, especially near the beginning or end of the sensor's life, due to simplistic interpretation of the BG values ​​or calibration codes. Summary of the Invention [Means for solving the problem]

[0006] In a first aspect, a method is presented for providing data representative of a concentration of an analyte in a patient, the method comprising: receiving a signal from an analyte sensor located within the patient's body; independently modeling at least one factor affecting the signal, where the at least one factor results from an individualized characteristic of the sensor and / or an individualized physiological characteristic of the patient; receiving individualized characteristic data associated with the individualized characteristic of the sensor and / or the individualized physiological characteristic of the patient; modifying a model of one or more of the at least one factor independently modeled based on receiving the individualized characteristic data; and outputting data representative of the concentration of the analyte in the patient based at least in part on the modified one or more models.

[0007] In one embodiment of the first aspect, the factors that are modeled independently include diffusion time lag, diffusion enzyme activity, and / or IG-BG kinetics.

[0008] In one embodiment of the first aspect, the factors that are modeled independently include the sensitivity of the sensor and / or the baseline response of the sensor.

[0009] In one embodiment of the first aspect, the analyte sensor is an enzyme-based electrochemical sensor and the individualized sensor characteristic is a sensor characteristic associated with the enzyme-based electrochemical sensor.

[0010] In one embodiment of the first aspect, the enzyme-based electrochemical sensor uses a glucose oxidase enzyme.

[0011] In one embodiment of the first aspect, the enzyme-based electrochemical sensor measures H2O2 produced by an enzyme-catalyzed reaction of glucose.

[0012] In an embodiment of the first aspect, the individualized sensor characteristic comprises a physiochemical characteristic of the analyte sensor.

[0013] In an embodiment of the first aspect, modeling the electrochemical break-in factors includes modeling factors associated with an interference layer of the analyte sensor independent of factors associated with a catalytic surface of the analyte sensor.

[0014] In one embodiment of the first aspect, the individualized physiological patient characteristic that modifies one or more of the models is a change in signal during cellular consumption of the analyte surrounding the insertion site of the sensor.

[0015] In one embodiment of the first aspect, modeling the IG-BG kinetics comprises modeling compartment bias.

[0016] In an embodiment of the first aspect, modeling the compartment biases includes modeling a steady-state compartment bias component separately from a time-lag compartment bias component.

[0017] In one embodiment of the first aspect, the independently modeled factors include gradual sensor degradation, which causes the signal to decrease as the analyte sensor approaches the end of its life.

[0018] In one embodiment of the first aspect, the individualized characteristic data includes data received prior to sensor insertion.

[0019] In one embodiment of the first aspect, at least a portion of the individualized characteristic data is received after an in-vivo sensor session begins.

[0020] In one embodiment of the first aspect, the personalized characteristic data associated with the personalized sensor characteristic includes factory-derived information.

[0021] In an embodiment of the first aspect, the individualized characteristic data associated with the individualized physiological patient characteristic comprises a measure of compartment bias.

[0022] In an embodiment of the first aspect, the individualized characteristic data comprises data comprising a measure of in vivo impedance.

[0023] In one embodiment of the first aspect, the factors to be modelled comprise enzyme activity, and modelling of the enzyme activity is performed using the Michaelis-Menten equation.

[0024] In one embodiment of the first aspect, the personalized characteristic data associated with the personalized sensor characteristic includes levels of glucose or oxygen exposure over the lifetime of the analyte sensor.

[0025] In one embodiment of the first aspect, one or more factors that affect the analyte component of the signal are modeled independently of one or more factors that affect the non-analyte component of the signal.

[0026] In one embodiment of the first aspect, one of the factors modeled is an enzymatic reaction with the analyte.

[0027] In one embodiment of the first aspect, the enzyme is glucose oxidase and the analyte is glucose.

[0028] In one embodiment of the first aspect, one of the factors modeled is diffusion of the analyte through one or more membrane layers of the analyte sensor.

[0029] In one embodiment of the first aspect, the diffusion time lag factor is modeled by the physical and thermodynamic sensor properties.

[0030] In an embodiment of the first aspect, the diffusion time lag factor is modeled based on a time shift model, a transfer function of the diffusion process, or deconvolution.

[0031] In one embodiment of the first aspect, the factors modeled include electrochemical break-in, and the electrochemical break-in is modeled as a function of hydration.

[0032] In an embodiment of the first aspect, receiving the individualized characteristic data includes receiving a physiological characteristic of the patient during a previous sensor session.

[0033] In one embodiment of the first aspect, one of the factors that is modeled independently is the non-stationary noise component of the signal.

[0034] In one embodiment of the first aspect, the enzyme-based electrochemical sensor employs a membrane system disposed over at least a portion of the electroactive surface of the analyte sensor, and one or more of the factors modeled is diffusion through the membrane system of electroactive compounds that interfere with the signal.

[0035] In one embodiment of the first aspect, revising the one or more models includes revising a model of a sensitivity curve of the analyte sensor over a sensor session using factory-derived sensor characteristic information.

[0036] In an embodiment of the first aspect, modifying the one or more models includes modifying a model of diffusion or loss of hydrogen peroxide in and around the sensor.

[0037] In an embodiment of the first aspect, the model of diffusion or loss of hydrogen peroxide in and around the sensor is based on pre-set model parameters.

[0038] In one embodiment of the first aspect, modifying the model of diffusion or loss of hydrogen peroxide in and around the sensor includes adaptively modifying pre-set model parameters based on individualized sensor characteristics.

[0039] In one embodiment of the first aspect, the individualized sensor characteristic comprises a measure of the cumulative exposure of the analyte sensor to in vivo temperature, glucose and / or oxygen concentration over time.

[0040] In one embodiment of the first aspect, modifying the one or more models includes modifying a model of a glucose oxidase enzyme with glucose.

[0041] In one embodiment of the first aspect, the model of the glucose oxidase enzyme with glucose uses the Michaelis-Menten equation.

[0042] In an embodiment of the first aspect, modifying the model of the glucose oxidase enzyme with glucose includes adaptively modifying pre-set model parameters based on the individualized sensor characteristics.

[0043] In one embodiment of the first aspect, the individualized sensor characteristics include data obtained from a factory calibration check.

[0044] In one embodiment of the first aspect, modifying the one or more models includes modifying a model of diffusion of glucose through one or more membrane layers of the analyte sensor based on pre-set model parameters.

[0045] In an embodiment of the first aspect, modifying the model of glucose diffusion through one or more membrane layers of the analyte sensor includes adaptively modifying pre-set model parameters based on the individualized sensor characteristics.

[0046] In one embodiment of the first aspect, the individualized sensor characteristic is selected from the group including maximum enzyme reaction rate, sensor resistive layer thickness, sensor enzyme layer thickness, sensor wire interference layer thickness, and sensor wire dimensions.

[0047] In an embodiment of the first aspect, modifying the one or more models includes modifying a model of electrochemical break-in based on pre-set model parameters.

[0048] In an embodiment of the first aspect, modifying the model of electrochemical break-in includes adaptively modifying pre-set model parameters based on the individualized sensor characteristics.

[0049] In one embodiment of the first aspect, the personalized sensor characteristics are selected from a group including factory-derived sensor measurements, field data, time since sensor insertion, and information obtained during sensor insertion verification.

[0050] In one embodiment of the first aspect, modifying the one or more models includes modifying a model of systemic and / or local responses to the analyte sensor generated by physiological species based on pre-set model parameters.

[0051] In one embodiment of the first aspect, modifying the model of the systemic and / or local responses generated by the physiological species includes adaptively modifying pre-set model parameters based on individualized physiological patient characteristics.

[0052] In one embodiment of the first aspect, the individualized physiological patient characteristic comprises in vivo oxygen concentration over time.

[0053] In one embodiment of the first aspect, the individualized physiological patient characteristic comprises a metabolic or wound healing response of the patient at the insertion site of the analyte sensor.

[0054] In one embodiment of the first aspect, modifying the one or more models includes modifying the sensor end of life model based on individualized physiological patient characteristics.

[0055] In one embodiment of the first aspect, modifying the one or more models includes modifying a model of sensor signal degradation over an initial period after sensor insertion based on individualized physiological patient characteristics.

[0056] In one embodiment of the first aspect, the model and model parameters based on averages across population data are modified based on individualized physiological patient characteristics.

[0057] In one embodiment of the first aspect, the individualized physiological patient characteristics include the patient's age and body mass index (BMI).

[0058] In one embodiment of the first aspect, modifying the one or more models includes modifying a descent recovery compensation model that has been pre-optimized for a predetermined period of time.

[0059] In one embodiment of the first aspect, the drop recovery compensation model is modified using individualized physiological patient characteristics and multiple shorter drop recovery compensation models, each of which lasts for a shorter period of time than the drop recovery compensation model.

[0060] In a second aspect, a system for providing data representative of a concentration of an analyte in a patient is presented, the system comprising: continuous analyte sensor electronics coupled to a continuous analyte sensor that generates data indicative of an analyte concentration in the patient; and a computing device in communication with the continuous analyte sensor, the computing device comprising a continuous analyte monitoring application installed on the computing device, the continuous analyte monitoring application configured to: receive a signal from the continuous analyte sensor located in interstitial fluid of the patient; independently model at least one factor affecting the signal, the at least one factor resulting from an individualized characteristic of the sensor and / or an individualized physiological characteristic of the patient; receive individualized characteristic data associated with the individualized characteristic of the sensor and / or the individualized physiological characteristic of the patient; modify a model of one or more of the at least one factor independently modeled based on receiving the individualized characteristic data; and output data representative of the concentration of the analyte in the patient based at least in part on the modified one or more models. [Brief description of the drawings]

[0061] [Figure 1] FIG. 1 is a block diagram of an example of an integrated system including a continuous glucose sensor and a drug delivery device. [Diagram 2] FIG. 1 is a front view of an electronic device configured for use with the present systems and methods. [Diagram 3] FIG. 2 is a functional block diagram of the electronic device of FIG. 1. [Figure 4A] FIG. 2 is a perspective view of an in vivo portion of an analyte sensor according to one embodiment. [Figure 4B] FIG. 2 is a perspective view of an in vivo portion of an analyte sensor according to one embodiment. [Figure 4C] FIG. 2 is a perspective view of an in vivo portion of an analyte sensor according to one embodiment. [Figure 4D] FIG. 2 is a perspective view of an in vivo portion of an analyte sensor according to one embodiment. [Figure 4E] 2E is a cross-sectional view of the analyte sensor of FIG. 2A taken along line 2E-2E. [Figure 4F] FIG. 1 illustrates a cross-sectional view of a membrane system in one embodiment. [Figure 5A] FIG. 1 is a cross-sectional view of a membrane system in one embodiment, showing the diffusion distance D1 between H2O2 generated within the enzyme domain and the electroactive surface. [Figure 5B] FIG. 2 is a cross-sectional view of a membrane system in another embodiment, showing the diffusion distance D2 between H2O2 generated within the enzyme domain and the electroactive surface. [Figure 5C] FIG. 13 is a cross-sectional view of a membrane system in another embodiment, showing the diffusion distance D3 between H2O2 generated within the enzyme domain and the electroactive surface. [Figure 6] 1 is a flow chart illustrating an example of a method for providing clinical data indicative of a blood analyte concentration in a patient. [Figure 7] 13 shows the sensor sensitivity change from temperature excursions resulting in an increase and decrease in sensor sensitivity before sensor insertion. [Figure 8]13 shows how a model of sensor sensitivity over time changes based on temperature over time. DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

[0062] The following description and examples show preferred embodiments of the present invention in detail. Those skilled in the art will recognize that there are many variations and modifications of the present invention that are encompassed by the scope of the present invention. Therefore, the description of the preferred embodiment should not be considered as limiting the scope of the present invention.

[0063] A. Overview Exemplary embodiments disclosed herein relate to the use of glucose sensors to measure the concentration of a substance indicative of the concentration or presence of glucose or another analyte. In some embodiments, the glucose sensor is a continuous device, e.g., subcutaneous, transdermal, transcutaneous, non-invasive, ocular, and / or intravascular (e.g., intravenous) device. In some embodiments, the device can analyze multiple intermittent blood samples. The glucose sensor can use any method of glucose measurement, including enzymatic, chemical, physical, electrochemical, optical, photochemical, fluorescence-based, spectrophotometric, spectroscopic (e.g., optical absorption spectroscopy, Raman spectroscopy, etc.), polarimetric, calorimetric, iontophoretic, and radiometric.

[0064] The glucose sensor can use any known detection method, including invasive, minimally invasive, and non-invasive sensing techniques, to provide a data stream indicative of the concentration of an analyte in a host. The data stream is typically a raw data signal that is used to provide a useful value of the analyte to a user, such as a patient or a medical professional (e.g., a physician), who may be using the sensor.

[0065] While many of the descriptions and examples are directed to an implantable glucose sensor that can measure the concentration of glucose in a host, the systems and methods of the embodiments may be applied to any measurable analyte. Some exemplary embodiments described below utilize an implantable glucose sensor. However, it should be understood that the devices and methods described herein may be applied to any device that can detect the concentration of an analyte and provide an output signal representative of the concentration of the analyte.

[0066] The following description and examples explain the present embodiments with reference to the drawings, in which reference numbers indicate elements of the present embodiments, and which are reproduced below with reference to the description of the corresponding drawing features.

[0067] b. Exemplary Continuous Glucose Monitoring System Exemplary embodiments disclosed herein relate to the use of glucose sensors to measure the concentration of a substance indicative of the concentration or presence of glucose or another analyte. In some embodiments, the glucose sensor is a continuous device, e.g., subcutaneous, transdermal, transcutaneous, non-invasive, ocular, and / or intravascular (e.g., intravenous) device. In some embodiments, the device can analyze multiple intermittent blood samples. The glucose sensor can use any method of glucose measurement, including enzymatic, chemical, physical, electrochemical, optical, photochemical, fluorescence-based, spectrophotometric, spectroscopic (e.g., optical absorption spectroscopy, Raman spectroscopy, etc.), polarimetric, calorimetric, iontophoretic, and radiometric.

[0068] The glucose sensor can use any known detection method, including invasive, minimally invasive, and non-invasive sensing techniques, to provide a data stream indicative of the concentration of an analyte in a host. The data stream is typically a raw data signal that is used to provide a useful value of the analyte to a user, such as a patient or a medical professional (e.g., a physician), who may be using the sensor.

[0069] While many of the descriptions and examples are directed to glucose sensors that can measure the concentration of glucose in a host, the systems and methods of the embodiments may be applied to any measurable analyte. Some exemplary embodiments described below utilize an implantable glucose sensor. However, it should be understood that the devices and methods described herein may be applied to any device that can detect the concentration of an analyte and provide an output signal representative of the concentration of the analyte.

[0070] As mentioned above, in some embodiments, the analyte sensor is an implantable glucose sensor as described with reference to U.S. Patent No. 6,001,067 and U.S. Patent Publication No. 2011-0027127 (A1). In some embodiments, the analyte sensor is a transcutaneous glucose sensor as described with reference to U.S. Patent Publication No. 2006-0020187 (A1). In yet other embodiments, the analyte sensor is a dual electrode analyte sensor as described with reference to U.S. Patent Publication No. 2009-0137887 (A1). In yet other embodiments, the sensor is configured to be implanted within a host's blood vessel or outside the body as described in U.S. Patent Publication No. 2007-0027385 (A1). These patents and publications are incorporated herein by reference in their entirety.

[0071] The following description and examples explain the present embodiments with reference to the drawings, in which reference numbers indicate elements of the present embodiments, and which are reproduced below with reference to the description of the corresponding drawing features.

[0072] FIG. 1 is a block diagram of a preferred embodiment integrated system including a continuous glucose sensor and a drug delivery device. This is an exemplary environment in which some embodiments described herein may be implemented. Here, an analyte monitoring system 100 includes a continuous analyte sensor system 8. The continuous analyte sensor system 8 includes a sensor electronics module 12 and a continuous analyte sensor 10. The system 100 may also include other devices and / or sensors, such as a drug delivery pump 2 and a reference analyte meter 4. The continuous analyte sensor 10 may be physically connected to the sensor electronics module 12, may be integral with the continuous analyte sensor 10 (e.g., non-releasably attached thereto), or may be releasably attachable thereto. Alternatively, the continuous analyte sensor 10 may be physically separate from the sensor electronics module 12, but electronically coupled thereto, such as via inductive coupling. Additionally, the sensor electronics module 12, the drug delivery pump 2, and / or the analyte reference meter 4 may communicate with one or more additional devices, such as any or all of the display devices 14, 16, 18, and / or 20. Display devices 14, 16, 18, and 20 generally include sufficient processors, memory, storage, and other components to run applications, including decision support modules.

[0073] In some implementations, the system 100 of FIG. 1 may also include a cloud-based processor 22 configured to analyze analyte data, drug delivery data, and / or other user-related data provided directly or indirectly via the network 24 from one or more of the sensor system 8, drug delivery pump 2, reference analyte meter 4, and display devices 14, 16, 18, 20. Based on the received data, the processor 22 may further process the data, generate reports providing statistics based on the processed data, trigger notifications to electronic devices associated with the host or a caregiver of the host, or provide the processed information to any of the other devices of FIG. 1. In some exemplary implementations, the cloud-based processor 22 comprises one or more servers. When the cloud-based processor 22 comprises multiple servers, the servers may be either geographically local or remote from each other. The network 24 may include any wired and wireless communication medium for transmitting data, including WiFi networks, cellular networks, the Internet, and any combination thereof.

[0074] In some example implementations, sensor electronics module 12 may include electronic circuitry associated with measuring and processing data generated by continuous analyte sensor 10. This generated continuous analyte sensor data may also include algorithms that may be used to process and calibrate the continuous analyte sensor data, although these algorithms may be provided in other manners as well, such as by devices 14, 16, 18, and / or 20. Sensor electronics module 12 may include hardware, firmware, software, or a combination thereof for providing a measurement of an analyte level via a continuous analyte sensor, such as a continuous glucose sensor.

[0075] Sensor electronics module 12 may be coupled (e.g., wirelessly, etc.) with one or more devices, such as any or all of display devices 14, 16, 18, and 20, as described above. Display devices 14, 16, 18, and / or 20 may be configured to process and present information, such sensor information, transmitted by sensor electronics module 12 for display on the display devices. Display devices 14, 16, 18, and 20 may also trigger alarms and / or provide decision support recommendations based on the analyte sensor data.

[0076] 1, display device 14 is a key fob-like display device, display device 16 is a handheld application-specific computing device 16 (e.g., a DexCom G4® Platinum receiver available from DexCom, Inc.), display device 18 is a general-purpose smartphone or tablet computing device 20 (e.g., a phone running the Android® OS, an Apple® iPhone®, iPad®, or iPod Touch® available from Apple, Inc.), and display device 20 is a computer workstation 20. In some exemplary implementations, the relatively small key fob-like display device 14 may be a computing device embodied in a watch, belt, necklace, pendant, jewelry, adhesive patch, pager, key fob, plastic card (e.g., credit card), identification (ID) card, or the like. This small display device 14 may include a relatively small (e.g., smaller than display device 18) display and may be configured to display a limited set of displayable sensor information, such as numbers 26 and arrows 28. Some systems may also include a wearable device 21, such as that described in U.S. Provisional Patent Application No. 61 / 904,341, filed November 14, 2013, entitled "Devices and Methods for Continuous Analyte Monitoring," the entire disclosure of which is expressly incorporated herein by reference. Wearable device 21 may include any device that is attached to or integrated into the user's optics, clothing, and / or body.Examples of devices include wearable devices, anklets, eyeglasses, rings, necklaces, armbands, pendants, belt clips, hair clips / ties, pins, cufflinks, tattoos, stickers, socks, sleeves, gloves, clothing (e.g., shirts, pants, underwear, bras, etc.), "clothing jewelry" such as zipper pulls, buttons, watches, shoes, contact lenses, subdermal implants, eyeglasses, cochlear implants, shoe inserts, braces (mouth), braces (body), medical wrappings, sports bands (wristbands, headbands), hats, bandages, hair weaves, nail polish, artificial joints / body parts, orthopedic pins / devices, implantable cardiac or neurological devices, etc. Miniature wearable device 14 and / or display device 21 may include a relatively small (e.g., smaller than display device 18) display and may be configured to display graphical and / or numerical representations of sensor information, such as numbers 26 and / or arrows 28. In contrast, display devices 16, 18, and 20 may be larger display devices that may be capable of displaying a larger set of displayable information, such as trend graph 30 depicted on handheld receiver 16, in addition to other information such as numbers and arrows.

[0077] It should be understood that any other user equipment (e.g., computing device) configured to present at least information (e.g., drug delivery information, discrete self-monitoring analyte readings, heart rate monitor, calorie intake monitor, etc.) may be used in addition to or instead of those described with reference to FIG. 1.

[0078] 1, the continuous analyte sensor 10 comprises a sensor for detecting and / or measuring an analyte, and the continuous analyte sensor 10 may be configured to continuously detect and / or measure an analyte as a non-invasive, subcutaneous, transcutaneous, and / or intravascular device. In some exemplary implementations, the continuous analyte sensor 10 may analyze multiple intermittent blood samples, although other analytes may be used as well.

[0079] In some example implementations of FIG. 1, the continuous analyte sensor 10 may include a glucose sensor configured to measure glucose in blood using one or more measurement techniques, such as enzymatic, chemical, physical, electrochemical, fluorescent, spectrophotometric, polarimetric, calorimetric, iontophoretic, radiometric, immunochemical, etc. In implementations in which the continuous analyte sensor 10 includes a glucose sensor, the glucose sensor may include any device capable of measuring the concentration of glucose and may provide data, such as a data stream indicative of the concentration of glucose in a host, using a variety of techniques for measuring glucose, including invasive, minimally invasive, and non-invasive sensing techniques (e.g., fluorescence monitoring). This data stream may be a raw data signal that is converted into a calibrated and / or filtered data stream that is used to provide glucose values ​​to a host, such as a user, patient, or caregiver (e.g., a parent, relative, guardian, teacher, doctor, nurse, or any other individual interested in the health of the host). Additionally, the continuous analyte sensor 10 may be implanted as at least one of the following types of sensors: an implantable glucose sensor, a transcutaneous glucose sensor implanted within a host's blood vessel or externally, a subcutaneous sensor, a replaceable subcutaneous sensor, a transocular sensor, or an intravascular sensor.

[0080] 2 illustrates one embodiment of an electronic device 200 configured for use with the present systems and methods. The electronic device 200 includes a display 202 and one or more input / output (I / O) devices, such as one or more buttons 204 and / or switches 206 that perform one or more functions when activated or clicked. In some embodiments, the electronic device 200 may be a mobile communication device. For example, in the illustrated embodiment, the electronic device 200 is a smartphone and the display 202 includes a touch screen that also functions as an I / O device. In other embodiments, the electronic device 200 may include one or more devices other than a smartphone, such as a receiver for a CGM system, a smart watch, a tablet computer, a mini tablet computer, a handheld personal digital assistant (PDA), a game console, a multimedia player, a wearable device as described above, a screen in an automobile or other vehicle, etc. Although electronic device 200 is shown in the figures as a smartphone, electronic device 200 may be any of the other electronic devices mentioned herein and / or may incorporate any or all of the functionality of the other electronic devices, including some or all of the functionality being embodied on a remote server.

[0081] Figure 3 is a block diagram of the electronic device 200 shown in Figure 2 and illustrates its functional components according to some embodiments. The electronic device 200 includes a display 202 and one or more input / output ("input / output, I / O") devices 204, 206, as described above with respect to Figure 2. The display 202 can be any device capable of displaying output, such as an LCD or LED screen. The input / output (I / O) devices 202, 204, 206 may include, for example, a keyboard (not shown), one or more buttons 204, one or more switches 206, etc. In embodiments including a touch screen, the display 202 also functions as an I / O device.

[0082] The electronic device 200 further includes a processor 208 (also referred to as a central processing unit (CPU)), a memory 210, a storage device 212, a transceiver 214, and may include other components or devices (not shown). The memory 210 is coupled to the processor 208 via a system bus or a local memory bus 216. The processor 208 may be or include one or more programmable general-purpose or special-purpose microprocessors, digital signal processors (DSPs), programmable controllers, application specific integrated circuits (ASICs), programmable logic devices (PLDs), or the like, or a combination of such hardware-based devices.

[0083] The memory 210 provides the processor 208 with access to data and program information stored therein during execution. Typically, the memory 210 includes random access memory (RAM) circuitry, read-only memory (ROM), flash memory, etc., or a combination of such devices.

[0084] Storage device 212 may include one or more internal and / or external mass storage devices, which may be or include any conventional medium for non-volatile storage of large amounts of data. For example, storage device 212 may include conventional magnetic disks, optical disks, magneto-optical (MO) storage, flash-based storage devices, or any other type of non-volatile storage device suitable for storing structured or unstructured data. Storage device 212 may also include storage in the "cloud" using so-called cloud computing. Cloud computing relates to computing capabilities that provide abstraction between computing resources and their underlying technical architecture (e.g., servers, storage, networks), enabling convenient, on-demand network access to a shared pool of configurable computing resources that can be rapidly provisioned and released with minimal management effort or service provider interaction.

[0085] The electronic device 200 may perform various processes, such as, for example, correlating data, pattern analysis, and other processes. In some embodiments, the electronic device 200 may perform such processes itself. Alternatively, such processes may be performed by one or more other devices, such as one or more cloud-based processors 22 described above. In still further embodiments, these processes may be performed partly by the electronic device 200 and partly by other devices. Various exemplary processes are described herein with reference to the electronic device 200. It should be understood that these exemplary processes are not limited to being performed solely by the electronic device 200. Furthermore, the term "electronic device" as used herein should be construed to include other devices having interaction with the electronic device 200, such as one or more cloud-based processors, servers, etc.

[0086] The electronic device 200 may also include other devices / interfaces for performing various functions. For example, the electronic device 200 may include a camera (not shown).

[0087] The transceiver 214 allows the electronic device 200 to communicate with other computing systems, storage devices, and other devices over a network. Although the illustrated embodiment includes the transceiver 214, in alternative embodiments, a separate transmitter and a separate receiver can be used in place of the transceiver 214.

[0088] In some embodiments, processor 208 can execute various applications (e.g., a CGM application) that may be downloaded to electronic device 200, such as via the Internet and / or a cellular network. Data for the various applications may be shared between electronic device 200 and one or more other devices / systems and stored by storage 212 and / or on one or more other devices / systems. The CGM application may include a decision support module and / or may include sufficient processing to operate decision support evaluation functions and methods as described below.

[0089] In some of the present embodiments, the sensor 10 of the continuous analyte sensor system 8 of FIG. 1 is inserted into the skin of a host. A new sensor session is then initiated with the sensor 10, the sensor electronics 12, and the electronic device 200. A number of techniques can be used to initialize the sensor 10. For example, initialization may be triggered when the sensor electronics 12 engages the sensor 10. In another example, initialization may be triggered by a mechanical switch, such as a switch (not shown) on a snap-in base that receives the sensor electronics 12. The switch is automatically activated when the sensor electronics 12 snaps into the base. In another example, initialization may be menu-driven, and the user may be prompted by a user interface on the display 202 of the electronic device 200 to initiate initialization by making a selection on the user interface, such as by pressing a button or touching a designated area on the display 202 (which may comprise a touch screen). In another example involving a non-invasive sensor applied to the wearer's skin, the sensor 10 can sense when it is in contact with the skin and initiate automatically. Additionally, the analyte sensor system 8 can detect the use of a new sensor 10 using any of the techniques described above, automatically prompt the user to confirm the new sensor session via a prompt on the user interface of the system 8, and initiate an initialization response to the user confirmation in response to the prompt. Additional examples of initializing the sensor 10 are described in U.S. patent application Ser. No. 13 / 796,185, filed Mar. 12, 2013, the entire disclosure of which is incorporated herein by reference.

[0090] c. Analyte Sensor Configurations / Components Preferred embodiments provide a continuous analyte sensor that measures the concentration of an analyte of interest or a substance indicative of the concentration or presence of the analyte. In some embodiments, the analyte sensor is an invasive, minimally invasive, or non-invasive device, such as a subcutaneous, transcutaneous, intravascular, or extracorporeal device. In some embodiments, the analyte sensor is capable of analyzing multiple intermittent biological samples. The analyte sensor can use any method of analyte measurement, including enzymatic, chemical, physical, electrochemical, spectrophotometric, polarimetric, calorimetric, radiometric, etc.

[0091] In some embodiments, the analyte sensor can be broadly characterized as a diffusion-based sensor. Some particular embodiments of the diffusion-based sensor can be more specifically an electrochemical or electrode-based sensor. In some embodiments, the electrochemical or electrode-based sensor can be an enzyme sensor, such as a GOX-based sensor or a GOX-based H2O2 sensor.

[0092] Generally, an analyte sensor provides at least one working electrode and at least one reference electrode that are configured to measure a signal associated with the concentration of an analyte in a host, as described in more detail below and understood by one of skill in the art. The output signal is typically a raw data stream that is used to provide, for example, a patient or physician with a useful value of the measured analyte concentration in the host. However, the analyte sensor of some embodiments includes at least one additional working electrode configured to measure at least one additional signal, as described elsewhere herein. For example, in some embodiments, the additional signal is associated with the baseline and / or sensitivity of the analyte sensor, thereby enabling monitoring of baseline and / or sensitivity changes that may occur in a continuous analyte sensor over time.

[0093] In general, continuous analyte sensors define a relationship between a sensor-generated measurement (e.g., current in pA, nA, or digital counts after A / D conversion) that is meaningful to a user (e.g., patient or physician) and a reference measurement (e.g., glucose concentration in mg / dL or mmol / L). In the case of implantable diffusion-based glucose oxidase electrochemical glucose sensors, the sensing mechanism generally relies on phenomena that are linear with glucose concentration, such as (1) the diffusion of glucose through a membrane system (e.g., a biointerface membrane and membrane system) located between the implantation site and / or electrode surface, (2) an enzymatic reaction within the membrane system, and (3) the diffusion of H2O2 to the sensor. Because of this linearity, the calibration of the sensor can be understood by solving the following equation: y=mx+b where y represents the sensor signal (e.g., counts), x represents the estimated glucose concentration (e.g., mg / dL), m represents the sensor sensitivity to glucose (e.g., counts / mg / dL), and b represents the baseline signal (e.g., counts). If both the sensitivity m and the baseline (background) b change over time in vivo, calibration generally requires at least two independent matched data pairs (x1, y1; x2, y2) to solve for m and b and thus enable glucose estimation when only the sensor signal y is available. Matched data pairs can be created by matching reference data (e.g., one or more reference glucose data points from a blood glucose meter, etc.) with substantially time-corresponding sensor data (e.g., one or more glucose sensor data points) to provide one or more matched data pairs, as described in U.S. Patent Application Publication No. 2005-0027463(A1). In some implantable glucose sensors, such as those described in more detail in U.S. Patent No. 6,329,161 to Heller et al., which is incorporated herein by reference in its entirety, the sensing layer utilizes an immobilized mediator (e.g., a redox compound) to electrically connect the enzyme to the working electrode rather than using a diffusive mediator. In some implantable glucose sensors, such as those described in more detail in U.S. Patent No. 4,703,756, the system has two oxygen sensors located in an oxygen permeable housing, one sensor unchanged and the other in contact with glucose oxidase, allowing for differential measurement of oxygen content in bodily fluids or tissues indicative of glucose levels. Various systems and methods for measuring glucose in a host are known, all of which can benefit from some embodiments that provide a sensor with a signal-to-noise ratio that is substantially unaffected by non-stationary noise.

[0094] Further description of analyte sensor configurations can be found in U.S. Patent Application Serial No. 11 / 692,154, entitled "DUAL ELECTRODE SYSTEM FOR A CONTINUOUS ANALYTE SENSOR," filed March 27, 2007, U.S. Patent Application Publication No. 2007-0027385(A1), and U.S. Patent Application Publication No. 2005-0143635(A1).

[0095] d. Overview of sensor components In some embodiments, the analyte sensor includes a sensing mechanism 34 with a small structure (e.g., a small structure, a micro or small diameter sensor), e.g., a needle-type sensor, at least in part (see FIG. 4A). As used herein, the term "small structure" refers to an architecture having at least one dimension preferably less than about 1 mm. The small structure sensing mechanism can be wire-based, substrate-based, or any other architecture. In some alternative embodiments, the term "small structure" can also refer to a slightly larger structure, such as one having a minimum dimension greater than about 1 mm, but the architecture (e.g., mass or size) is designed to minimize foreign body response (FBR) due to size and / or mass. In some embodiments, a biointerface membrane (e.g., a membrane system or sensing membrane) is formed on the sensing mechanism 34, as described in more detail below. In some alternative embodiments, the sensor is configured to be fully implanted within the host, such as the abdomen of the host, as described in U.S. Patent Application Publication No. 2006-0020187(A1). In yet other embodiments, the sensor is configured to be implanted within a host's blood vessel or outside the body, as described in U.S. Patent Application Publication No. 2007-0027385(A1), U.S. patent application Ser. No. 11 / 543,396, filed October 4, 2006, U.S. patent application Ser. No. 11 / 691,426, filed March 26, 2007, and U.S. patent application Ser. No. 11 / 675,063, filed February 14, 2007.

[0096] In various embodiments described below, as described in detail herein and understood by those of skill in the art, the sensor is an enzyme-based electrochemical sensor, where the working electrode 38 measures hydrogen peroxide (HO) produced by an enzyme-catalyzed reaction of the glucose being detected, generating a measurable electronic current (e.g., detection of glucose utilizing glucose oxidase produces hydrogen peroxide as a by-product, and HO reacts with the surface of the working electrode to release two protons (2H + ), 2 electrons (2e - ), and one molecule of oxygen (O2), which generates an electronic current that is detected. Preferably, one or more potentiostats are used to monitor the electrochemical reaction at the electroactive surface of the working electrode. The potentiostat applies a constant potential to the working electrode and its associated reference electrode to determine the current produced at the working electrode. The current produced at the working electrode (and flows through the circuit to the counter electrode) is substantially proportional to the amount of H2O2 that diffuses to the working electrode. The output signal is typically a raw data stream that is used to provide a host or physician, for example, with a useful value of the measured analyte concentration in the host. In some alternative embodiments, the sensing mechanism includes electrodes deposited on a planar substrate, and the implantable portion has a thickness of less than about 1 mm, see, for example, U.S. Pat. No. 6,175,752 and U.S. Pat. No. 5,779,665.

[0097] Some alternative analyte sensors that may benefit from the systems and methods of some embodiments include, for example, U.S. Patent No. 8,364,229, U.S. Patent No. 7,828,728, U.S. Patent No. 5,711,861, U.S. Patent No. 6,642,015, U.S. Patent No. 6,654,625, U.S. Patent No. 6,565,509, U.S. Patent No. 6,514,718, U.S. Patent No. 6,465,066, U.S. Patent No. 6,214,185, U.S. Patent No. 5,310,469, and U.S. Patent No. 5,683,562, U.S. Patent No. 6,579,690, U.S. Patent No. 6,484,046, U.S. Patent No. 6,512,939, and U.S. Patent No. 6,424,847. These patents are not inclusive of all applicable analyte sensors. It should be understood that the disclosed embodiments are generally applicable to a variety of analyte sensor configurations.

[0098] 4A is a close-up view of an exemplary embodiment of a continuous analyte sensor 34 (also referred to as a transcutaneous analyte sensor or needle-type sensor), particularly showing the sensing mechanism. Preferably, the sensing mechanism comprises a miniature structure as defined herein and is adapted for insertion under the skin of a host, while the remaining body of the sensor (e.g., electronics, etc.) can reside ex vivo. In the illustrated embodiment, the analyte sensor 34 includes two electrodes, a working electrode 38 and at least one additional electrode 30 (hereinafter referred to as the reference electrode 30) that can function as a counter electrode and / or a reference electrode.

[0099] In some exemplary embodiments, each electrode is formed from a thin wire having a diameter of, for example, about 0.001 inch or less to about 0.010 inch or more, and may be formed from, for example, plated insulation, plated wire, or bulk conductive material. The illustrated electrode configuration and associated text describes one preferred method of forming a transdermal sensor, however, various known transdermal sensor configurations may be used with the transdermal analyte sensor system of some embodiments, such as those described in U.S. Pat. No. 6,695,860 to Ward et al., U.S. Pat. No. 6,565,509 to Say et al., U.S. Pat. No. 6,248,067 to Causey III et al., and U.S. Pat. No. 6,514,718 to Heller et al.

[0100] In a preferred embodiment, the working electrode comprises a wire formed from a conductive material (e.g., platinum, platinum-iridium, palladium, graphite, gold, carbon, conductive polymers, alloys, etc.). Although the electrodes can be formed by a variety of manufacturing techniques (bulk metal processing, deposition of metal on a substrate, etc.), it may be advantageous to form the electrodes from plated wire (e.g., platinum on steel wire) or bulk metal (e.g., platinum wire). Electrodes formed from bulk metal wires are believed to offer superior performance (e.g., as opposed to deposited electrodes), including increased assay stability, simplified manufacturability, resistance to contamination (e.g., that may be introduced in the deposition process), and improved surface reactions without peeling or delamination (e.g., due to the purity of the material).

[0101] The working electrode 38 is configured to measure the concentration of an analyte, such as, but not limited to, glucose, uric acid, cholesterol, lactate, etc. For example, in an enzymatic electrochemical sensor for detecting glucose, the working electrode measures hydrogen peroxide produced by an enzyme-catalyzed reaction of the analyte being detected, generating a measurable electronic current. For example, in the detection of glucose, where glucose oxidase (GOX) produces hydrogen peroxide as a by-product, H2O2 reacts with the surface of the working electrode to release two protons (2H + ), 2 electrons (2e -) and one oxygen molecule (O2), which generates an electronic current that is detected.

[0102] The working electrode 38 is covered with an insulating material, such as a non-conductive polymer. Dip coating, spray coating, vapor deposition, or other coating or deposition techniques can be used to deposit the insulating material on the working electrode. In one embodiment, the insulating material includes parylene, which can be an advantageous polymer coating for its strength, lubricity, and electrical insulation properties. Generally, parylene is produced by vapor deposition and polymerization of paraxylylene (or its substituted derivatives). However, any suitable insulating material can be used, such as fluorinated polymers, polyethylene terephthalate, polyurethane, polyimide, other non-conductive polymers, and the like. Glass or ceramic materials can also be used. Other materials suitable for use include surface energy modification coating systems, such as those commercially available under the trade names AMC18, AMC148, AMC141, and AMC321 by Advanced Materials Components Express, Bellafonte, Pennsylvania. However, in some alternative embodiments, the working electrode may not require a coating of an insulator.

[0103] Preferably, the reference electrode 30, which may function as a reference electrode alone or as a dual reference and counter electrode, is formed from silver, silver / silver chloride, or the like. Preferably, the electrodes are juxtaposed and / or twisted on or around each other, although other configurations are possible. In one example, the reference electrode 30 is spirally wound around the working electrode 38, as shown in FIG. 4A. The assembly of wires may then optionally be coated with an insulating material similar to that described above to provide an insulating attachment (e.g., to secure the working and reference electrodes together).

[0104] As discussed above, conventional transdermal devices are believed to suffer from motion artifacts associated with host movement when the host is using the device. For example, it is known that when a transdermal analyte sensor is inserted into a host, various movements on the sensor (e.g., relative movements within and between the subcutaneous space, dermis, skin, and outer portions of the sensor) create stresses on the device that create artifacts (e.g., non-stationary noise) on the sensor signal.

[0105] Thus, there are different design considerations (e.g., stress considerations) for various portions of the sensor. For example, the in vivo portion of the sensor (e.g., the portion inserted through the host's skin and into the underlying tissue) may generally benefit from greater flexibility since it is subject to greater mechanical stresses caused by tissue movement within the patient and relative movement between the in vivo and ex vivo portions of the sensor. On the other hand, the ex vivo portion of the sensor (the portion of the sensor that remains outside the host's body) may generally benefit from a stiffer, more robust design to ensure structural integrity and / or reliable electrical connection. Additionally, in some embodiments where the needle is retracted beyond the ex vivo portion of the device, a stiffer design may minimize the ease of crimping and / or retraction of the needle from the sensor. Thus, by designing greater flexibility into the in vivo portion, it is believed that this flexibility compensates for patient movement and associated noise. By designing a higher stiffness into the ex vivo portion, column strength (against needle retraction on the sensor), electrical connection, and integrity may be improved. In some alternative embodiments, a stiffer distal end and / or a more flexible proximal end may be advantageous, as described in US Patent Application Publication Nos. 2006-0015024(A1) and 2006-0020187(A1).

[0106] Some preferred embodiments provide an in vivo portion of the sensor that is designed to be more flexible than the ex vivo portion of the sensor. The variable stiffness of the sensor of the preferred embodiments may be provided by a variable pitch of any one or more helically wound wires of the device, a variable cross-section of any one or more wires of the device, and / or a variable hardening and / or softening of any one or more wires of the device, as described in more detail with reference to U.S. Patent Application Publication Nos. 2006-0015024(A1) and 2006-0020187(A1).

[0107] In embodiments where an outer insulator is disposed, a portion of the coated assembly structure can be peeled or otherwise removed, for example, by hand, excimer laser oscillation, chemical etching, laser ablation, grit blasting (e.g., with sodium bicarbonate or other suitable grit), etc., to expose the electroactive surface. Alternatively, a portion of the electrode can be masked prior to depositing the insulator to maintain exposed electroactive surface area. In one exemplary embodiment, grit blasting is implemented to expose the electroactive surface, preferably utilizing a grit material that is hard enough to ablate the polymeric material, but soft enough to minimize or avoid damage to the underlying metal electrode (e.g., platinum electrode). While a variety of "grit" materials (e.g., sand, talc, walnut shells, crushed plastic, sea salt, etc.) can be used, in some preferred embodiments, sodium bicarbonate is an advantageous grit material because it is hard enough to ablate, for example, the parylene coating, without damaging the underlying platinum conductor. One additional benefit of sodium bicarbonate blasting includes its abrasive action on metal as it strips away the polymer layer, thereby eliminating a cleaning step that might otherwise be necessary.

[0108] In some embodiments, a radial window is formed through the insulating material to expose the circumferential electroactive surface of the working electrode. In addition, a portion of the electroactive surface of the reference electrode is exposed. For example, the portion of the electroactive surface can be masked during deposition of the outer insulating layer or etched after deposition of the outer insulating layer. In some applications, cellular attack or migration of cells to the sensor can cause a decrease in sensitivity and / or function of the device, especially after the first day of implantation. However, when the exposed electroactive surface is distributed circumferentially around the sensor (e.g., as in a radial window), the surface area available for reaction can be sufficiently distributed to minimize the effect of local cellular intrusion of the sensor on the sensor signal. Alternatively, a tangentially exposed electroactive window can be formed, for example, by peeling off only one side of the coated assembly structure. In other alternative embodiments, a window can be provided at the tip of the coated assembly structure such that the electroactive surface is exposed at the tip of the sensor. Other methods and configurations for exposing the electroactive surface can also be used.

[0109] In general, the sensor electrodes can be configured to provide a sensor having a signal-to-noise ratio that is substantially unaffected by non-stationary noise, such as by systems and methods configured to increase the analyte signal component and / or reduce the non-stationary noise component.

[0110] In some circumstances, noise can be caused (e.g., during the use of an amperometric GOX sensor with a platinum working electrode) by the accumulation of molecular oxygen (O2) on a platinum electrode, which is produced during the electro-oxidation of H2O2 to water and O2. Platinum black applied to the working electrode can prevent O2 accumulation on the platinum working electrode, thereby preventing noise on the sensor. Platinum black is a fine black powder of metallic platinum that can be formed into a paste, ink, or paint-like material that can be applied to a surface (e.g., wire, plastic support) to produce a rough, large surface area coating that is relatively non-polarizable. Platinum black can be applied to the electrode surface using a platinization process (e.g., on platinum wire or other platinum surfaces) or known thin film techniques (e.g., dipping, painting, or screen printing). One further advantage of using platinum black is that platinum electrodes coated with platinum black provide substantial signals from hydrogen peroxide oxidation at working potentials as low as 150 mV, whereas platinum electrodes not coated with platinum black must be operated at potentials of at least 600 mV. Because fewer interferents can be electrooxidized / reduced at the lower working potential, sensors operated at potentials below 600 mV have reduced non-stationary noise on the signal. Thus, in one embodiment, at least the working electrode is coated with platinum black, thereby reducing the non-stationary noise component of the signal and thus adjusting (e.g., increasing) the signal-to-noise ratio.

[0111] In some embodiments, the signal-to-noise ratio can be made substantially unaffected by non-stationary noise by distributing the electroactive surface area along a substantial length of the in vivo portion of the sensor. In some embodiments, the electroactive surface area is distributed along 50%, 55%, 60%, 65%, 70%, 75%, 80%, 85%, 90%, 95%, or more of the in vivo portion of the sensor. It is believed that certain interfering electroactive species scatter (e.g., along the length of the in vivo portion of the sensor) and impinge on the sensor. In other words, some transient interferents do not necessarily contact the sensor uniformly along the in vivo portion of the sensor. For example, the tissue surrounding an implanted sensor is highly variable. In some circumstances, the electroactive surface may be positioned adjacent to a lymphatic vessel (when the sensor is implanted), which can reduce the local concentration of electroactive species that may interfere with the analyte signal and result in a minimal non-stationary noise component of the signal. In other situations, the electroactive surface may be located in fat with poor circulation (when the sensor is implanted), which can result in the accumulation of electroactive species that can interfere with the sensor signal adjacent the electroactive surface and can result in more non-stationary noise in the sensor signal (e.g., than a surface adjacent to a lymphatic vessel).

[0112] 4B is a perspective view of an in vivo portion of an analyte sensor in one embodiment, where the electroactive surface area is divided into four equal portions (38a, 38b, 38c, 38d) and distributed (e.g., spaced apart from one another by distances of 1 mm, 2 mm, etc.) along a substantial length of the in vivo portion of the sensor. Thus, in some embodiments, the sensor is configured such that the electroactive surface area is distributed (e.g., extended, spaced apart, divided, and / or distributed) along a substantial length of the in vivo portion of the sensor, such that the signal is measured (and can be integrated or averaged) over a more dispersed or distributed portion of the in vivo portion of the sensor (e.g., for the sensor of FIG. 4A ) such that the signal contribution from non-stationary non-analyte components is less than about 20% of the total signal (e.g., after sensor break-in is complete). In some embodiments, the area of ​​the electroactive surface may be distributed in various ways, such as, but not limited to, two or more areas (e.g., 38a, 38b, 38c, 38d, etc.) that are cumulatively substantially equal to the desired total area of ​​the electroactive surface. The exposed area of ​​the electroactive surface may have any geometric shape, such as a circle, a dot, a rectangle, an oval, a star, etc. In some embodiments, the exposed surface area essentially functions as a microelectrode along the in vivo portion of the sensor. The microelectrode may increase the sensor sensitivity by increasing the utilization of the electroactive compound being measured, for example, due to the beneficial edge effect associated with multiple small surface areas in close proximity to each other. In one exemplary embodiment, when multiple electroactive surfaces are spaced apart (from each other) within the diffusion distance of H2O2 (e.g., compared to one or more electroactive surface areas spaced apart greater than the diffusion distance of H2O2), the multiple small spaced apart electroactive surface areas may detect more H2O2 produced by the glucose oxidase enzyme layer. In the illustrated embodiment, a wire-type miniature structure sensor is illustrated, however, the surface area of ​​a variety of other analyte sensor configurations (e.g., planar type sensor configurations including fully implantable sensors, intravascular sensors, implantable continuous sensors, and in vitro test strips) can be distributed as described herein.Methods for exposing the electroactive surface of the sensor are detailed in US Patent Application Publication No. 2006-0020187(A1).

[0113] FIG. 4C is a perspective view of an in vivo portion of an analyte sensor in one embodiment, where the area of ​​the electroactive surface is distributed along a substantial length of the in vivo portion of the analyte sensor. In this embodiment, the area of ​​the electroactive surface 38 is distributed by selecting a working electrode having a preferred width such that the exposed electroactive surface covers a substantial length of the in vivo portion of the working electrode while still requiring a relatively low current draw. For example, the electroactive surface can be distributed by using longer and thinner areas such that the electroactive surface covers a substantial length of the in vivo portion of the working electrode, while the total exposed surface area remains unchanged as compared to another sensor having a wider and shorter exposed electroactive surface area. For example, in some embodiments, the electrodes are formed from bulk metal wires having a diameter of about 0.001 to about 0.010 inches. For example, comparing two sensors, a first sensor having a working electrode formed of a 0.001 inch diameter wire with an electroactive surface length of 1 mm, and a second sensor having a working electrode formed of a 0.010 inch diameter wire with an electroactive surface length of 0.1 mm, the two sensors may have the same sensitivity, but the signal to noise ratio will be substantially less affected by non-stationary noise compared to the signal to noise ratio of the second sensor. In some embodiments, the preferred surface area of ​​the working electrode is approximately 0.0000839 cm, assuming a diameter of about 0.001 inch to about 0.010 inch and a length of about 0.004 inch to about 0.078 inch. 2 Less than ~ approx. 0.016129cm 2 That's all.

[0114] 4A-4D, some examples of discontinuous surfaces on a sensor are shown, including relatively sharp abrupt edges, substantially raised surface features, or substantial and / or abrupt changes in sensor diameter, such as at the sensor tip 31a and the in vivo end of the reference electrode 31b. In some circumstances, pockets of localized enhanced inflammatory response (e.g., accumulation of inflammatory cells) may form around portions of the sensor having discontinuous surfaces, such as the sensor tip 31a and the in vivo end of the reference electrode 31b. This phenomenon has been observed in histological sections of rat tissue explants (one day after miniature sensor implantation) as an accumulation of inflammatory cells around both the sensor tip 31a and the in vivo end of the reference electrode 31b. In some embodiments, slight diameter changes, such as at the edges of the electroactive surface area of ​​the working electrode, are designed to avoid or minimize enhanced inflammatory response as described above. Such enhanced inflammatory responses may be caused by the presence of increased populations of macrophages, lymphocytes, neutrophils, and / or foreign body giant cells, which stimulate the production of highly diffusive electroactive noise-causing species in the body, particularly free radicals (e.g., reactive oxygen and nitrogen species). Some of the noise-causing compounds produced by enhanced inflammatory responses may react at the working electrode of the sensor, resulting in noise (e.g., non-stationary noise). Thus, by separating the electroactive surface of the working electrode from the discontinuous surface by a distance sufficient to minimize or avoid the effects of inflammation at the discontinuous surface (e.g., a distance substantially greater than the diffusion distance of at least one non-stationary noise-generating electroactive species produced by inflammation), it is believed that the non-stationary noise on the signal can be reduced to less than about 20% of the total signal. In some preferred embodiments, the electroactive surface of the working electrode is spaced from about 0.020, 0.03, 0.04, or 0.05 inches or less to about 0.06, 0.07, 0.08, 0.09, or 0.100 inches or more from a discontinuous surface such that non-stationary noise is less than about 20% of the total signal and / or the analyte components are at least 80% of the total signal.For example, in one exemplary embodiment, the sensor is configured such that the reference electrode 30 is spaced from the radial window of the working electrode 38 such that the distance between the radial window (e.g., the edge closest to the reference electrode) and the in vivo termination 31b of the reference electrode (referred to herein as the "first distance") is at least about 0.020 inches, such that the non-stationary noise component is less than 20% of the total signal and / or the analyte component is at least 80% of the total signal. In some exemplary embodiments, the first distance is at least 0.030, 0.04, 0.050, 0.060, 0.070, 0.080, 0.090, or 0.100 inches or more. In some exemplary embodiments, the sensor is configured such that the distance between the tip 31a of the sensor and the nearest edge of the radial window (referred to herein as the "second distance") is at least 0.020 inches, such that the non-stationary noise component is less than 20% of the total signal and / or the analyte component is at least 80% of the total signal. In some exemplary embodiments, the second distance is at least 0.030, 0.04, 0.050, 0.060, 0.070, 0.080, 0.090, or 0.100 inches or more. In some circumstances, the preferred distance between the electroactive surface and the discontinuous surface of the working electrode to minimize or avoid enhanced inflammatory responses may vary depending on factors such as, for example, the type / size / characteristics of the material (e.g., electrode material, membrane system components, etc.) used to form the sensor, the tissue differences in which the sensor is implanted (e.g., fat type, lean or fat, etc.), the physical condition of the host (e.g., disease or injury), the condition of the wound produced during sensor implantation, etc.

[0115] FIG. 4D is a perspective view of the in vivo portion of an analyte sensor in one embodiment. In a preferred embodiment, the sensor is configured to minimize discontinuities, thereby reducing non-stationary noise on the signal caused by electroactive species produced from the body's inflammatory response to such discontinuous portions of the sensor. For example, in the embodiment shown in FIG. 4D, the sensor is configured to be substantially non-discontinuous (or substantially continuous) to minimize inflammatory responses, thus minimizing or preventing noise on the signal such that non-stationary noise is less than 20% of the total signal. In one exemplary embodiment as shown in FIG. 4D, the tip 31a of the sensor is tapered 31c. In another exemplary embodiment, discontinuities at the in vivo end of the reference electrode 31b are reduced (e.g., by minimizing steps at the in vivo end of the reference electrode) such that inflammatory responses are substantially avoided, non-stationary noise is less than about 20% of the total signal, and / or analyte components are at least about 80% of the total signal.

[0116] In some embodiments, the sensor is configured to substantially reduce the effect of noise-causing electroactive species caused by inflammation and / or FBR in response to discontinuous portions of the sensor. In some embodiments, the in vivo portion of the sensor is configured such that the electroactive surface 38 is farther from the end of the sensor tip 31a and / or reference electrode 31b than the diffusion distance of at least one noise-causing electroactive species resulting from a metabolic process of the host (e.g., H2O2 produced outside the sensor). In some embodiments, some or all of the discontinuous portions of the sensor are sufficiently smoothed and / or tapered such that inflammation and / or FBR are substantially minimized, so that noise-causing compounds produced by inflammation and / or FBR associated with the discontinuous portions do not substantially contribute to the signal. Thus, in a preferred embodiment, the in vivo portion of the sensor is configured to enable a signal with a non-stationary noise component of less than about 20% of the total signal.

[0117] In the exemplary sensors described above, an overall diameter of about 0.030 inches or less is preferred, more preferably about 0.020 inches or less, and even more preferably about 0.016 inches or less. In some embodiments, the exposed electroactive surface area has a width of about 0.001 inches or less to about 0.010 inches or more, preferably about 0.002 inches to about 0.008 inches, and more preferably about 0.004 inches to about 0.005 inches. The length of the window can be about 0.1 mm (about 0.004 inches) or less to about 2 mm (about 0.078 inches) or more, preferably about 0.5 mm (about 0.02 inches) to about 0.75 mm (0.03 inches). In such embodiments, the exposed surface area of ​​the working electrode is preferably about 0.000013 inches or less. 2 (0.0000839cm 2 ) or less ~ approx. 0.0025in 2 (0.016129cm 2 ) or greater (assuming a diameter of about 0.001 inches to about 0.010 inches and a length of about 0.004 inches to about 0.078 inches). In some embodiments, the exposed surface area of ​​the working electrode is selected to generate an analyte signal having a current in the picoamp range, as described in more detail elsewhere herein.

[0118] However, currents in the picoamp range may depend on various factors, such as the electronic circuit design (e.g., sample rate, current draw, A / D converter bit resolution, etc.), the membrane system (e.g., permeability of analytes through the membrane system), and the exposed surface area of ​​the working electrode. Thus, the exposed electroactive working electrode surface area may be selected to have a value greater or less than the ranges mentioned above, taking into account modifications to the membrane system and / or electronic circuitry. In preferred embodiments of glucose sensors, it may be advantageous to minimize the surface area of ​​the working electrode while maximizing the diffusion rate of glucose so that the signal-to-noise ratio is not substantially affected by non-stationary noise while maintaining sensor performance in both high and low glucose concentration ranges.

[0119] In some alternative embodiments, the exposed surface area of ​​the working electrode (and / or other electrodes) can be increased by modifying the cross-section of the electrode itself. For example, in some embodiments, the cross-section of the working electrode can be defined by a cross, star, cloverleaf, rib, dimple, ridge, irregular, or other non-circular configuration. Thus, for any given length of electrode, a certain increased surface area can be achieved (compared to the area achieved by a circular cross-section). Increasing the surface area of ​​the working electrode can be advantageous in providing an increased signal in response to analyte concentration, which in turn can help improve, for example, the signal-to-noise ratio.

[0120] In some alternative embodiments, additional electrodes may be included in the assembly, such as a three-electrode system (working, reference, and counter electrodes) and / or additional working electrodes (e.g., an electrode that can be used to generate oxygen, an electrode configured as a baseline subtraction electrode, or an electrode configured to measure an additional analyte). U.S. Pat. No. 7,081,195, U.S. Patent Application Publication No. 2005-0143635(A1), and U.S. Patent Application Publication No. 2007-0027385(A1), each of which is incorporated herein by reference, describe some systems and methods for implementing and using additional working, counter, and / or reference electrodes. In one implementation in which the sensor includes two working electrodes, the two working electrodes are juxtaposed (e.g., extend parallel to one another) around which the reference electrode is disposed (e.g., spirally wound). In some embodiments where two or more working electrodes are provided, the working electrodes can be formed in a double, triple, quadruple, etc., helical configuration along the length of the sensor (e.g., surrounding a reference electrode, an insulating rod, or other support structure). The resulting electrode system can be configured with a suitable membrane system, with a first working electrode configured to measure a first signal including glucose and a baseline, and an additional working electrode configured to measure a baseline signal consisting of only the baseline (e.g., configured substantially similarly to the first working electrode without an enzyme disposed thereon). In this manner, the baseline signal can be subtracted from the first signal to generate a glucose-only signal that is substantially free of fluctuations in baseline and / or interfering species on the signal. Thus, the dimensions described above can be modified as desired.

[0121] In some embodiments, the sensing area may include a reference electrode and / or other electrodes associated with the glucose-measuring working electrode, and / or a separate reference electrode and / or counter electrode associated with any auxiliary working electrode. In yet another embodiment, the sensing area may include a glucose-measuring working electrode, an auxiliary working electrode, two counter electrodes (one for each working electrode), and one shared reference electrode. In yet another embodiment, the sensing area may include a glucose-measuring working electrode, an auxiliary working electrode, two reference electrodes, and one shared counter electrode. However, a variety of electrode materials and configurations may be used with the implantable analyte sensors of the preferred embodiments.

[0122] U.S. Patent Application No. 11 / 543,396, filed October 4, 2006, and U.S. Patent Application Publication No. 2005-0245799(A1) describe additional configurations for use at different body locations. In one exemplary embodiment, the sensor is configured for percutaneous implantation into a host. In another exemplary embodiment, the sensor is configured for insertion into a circulatory system, such as a peripheral vein or artery. However, in other embodiments, the sensor is configured for insertion into a central circulatory system, such as, but not limited to, a vena cava. In still other embodiments, the sensor may be configured for insertion into an extracorporeal circulatory system, such as, but not limited to, a shunt (e.g., artery to vein), an extracorporeal blood chemistry analysis device, a dialysis machine, or a heart-lung machine (e.g., to pump blood during cardiac surgery). In yet another embodiment, the sensor may be configured to be fully implantable, as described in U.S. Patent No. 6,001,067.

[0123] Although some embodiments show one electrode configuration including one bulk metal wire helically wound around another bulk metal wire, other electrode configurations are contemplated. In an alternative embodiment, the working electrode includes a tube with a reference electrode disposed or coiled inside with an insulator therebetween. Alternatively, the reference electrode includes a tube having a working electrode disposed or coiled inside with an insulator therebetween. In another alternative embodiment, a polymer (e.g., insulating) rod is provided and the electrodes are deposited (e.g., electroplated) thereon. In yet another alternative embodiment, a metal (e.g., steel) rod coated with an insulating material is provided and the working and reference electrodes are deposited thereon. In yet another alternative embodiment, one or more working electrodes are helically wound around the reference electrode.

[0124] FIG. 4E is a cross-sectional view of the sensor of FIG. 4A taken along line 2E-2E, showing the membrane system 32 in one embodiment. In this embodiment, the membrane system includes an electrode domain 43 wrapped around a platinum wire working electrode 38, an interference domain 44, an enzyme domain 46, and a diffusion resistance domain 48. In some embodiments, the membrane system also includes a cell impermeable domain as described elsewhere herein. In some embodiments, a single resistance domain and a cell impermeable domain are included in the membrane system (shown in this figure as resistance domain 48). In some embodiments, the transcutaneous wire sensor is configured for short-term implantation (e.g., about 1-30 days).

[0125] 4F is a cross-sectional view of the membrane system 32 in another embodiment. The membrane system 32 can be used with a glucose sensor as described herein. In this embodiment, the membrane system 32 includes an electrode domain 43 that is closest to the electrochemically reactive surface of the working electrode, an (optional) interference domain 44 that is no closer to the electrochemically reactive surface of the working electrode than the electrode domain, an enzyme domain 46 that is no closer to the electrochemically reactive surface of the working electrode than the interference domain, a diffusion resistance domain 48 that is no closer to the electrochemically reactive surface of the working electrode than the enzyme domain, a cell impermeable domain 42 (also referred to as a bioprotective layer) that is no closer to the electrochemically reactive surface of the working electrode than the diffusion resistance domain, and an optional cell disruption domain 40 that is the most distal of all domains from the electrochemically reactive surface of the working electrode. However, it is understood that the membrane system 32 can be modified for use in other devices by including only two or more layers, or additional layers not listed above.

[0126] Generally, the sensing membrane 32 of some embodiments includes multiple domains or layers, such as an interference domain 44, an enzyme domain 46, and a resistance domain 48, and may include additional domains, such as an electrode domain 43, a cell impermeable domain 42 (also referred to as a bioprotective layer), and / or an oxygen domain (not shown), as described in more detail in the above-cited U.S. Patent Application Publications. However, it should be understood that sensing membranes modified for other sensors, for example by including fewer or additional domains, are within the scope of some embodiments. In some embodiments, one or more domains of the sensing membrane are formed from materials such as silicone, polytetrafluoroethylene, ethylene tetrafluoroethylene copolymers, polyolefins, polyesters, polycarbonates, biostable polytetrafluoroethylene, homopolymers, copolymers, terpolymers of polyurethane, polypropylene (PP), polyvinylchloride (PVC), polyvinylidene fluoride (PVDF), polybutylene terephthalate (PBT), polymethylmethacrylate (PMMA), polyether ether ketone (PEEK), polyurethanes, cellulosic polymers, poly(ethylene oxide), poly(propylene oxide) and copolymers and blends thereof, polysulfones and block copolymers thereof (e.g., diblock, triblock, alternating, random, and graft copolymers, etc.). US Patent Application Publication No. 2005-024579912(A1) describes biological interface and sensing membrane configurations and materials that may be applied in some embodiments.

[0127] In some embodiments, the sensing membrane can be deposited on the electroactive surface of the electrode material using known thin or thick film techniques (e.g., spraying, electrodeposition, dipping, etc.). Note that the sensing membrane surrounding the working electrode does not have to be of the same structure as the sensing membrane surrounding the reference electrode, etc. For example, an enzyme domain deposited on the working electrode does not necessarily have to be deposited on the reference and / or counter electrodes.

[0128] e. Sensor membrane system Generally, the analyte sensor of the preferred embodiment comprises a membrane system such as that illustrated in Figures 4E and 4F. Preferably, the membrane system is deposited over at least a portion of the electroactive surfaces of the sensor (working electrode and optionally the reference electrode) to provide protection of the exposed electrode surfaces from the biological environment, diffusion resistance (limitation) of the analyte if necessary, catalysis to enable enzymatic reactions, limitation or blocking of interferents, and / or hydrophilicity at the electrochemically reactive surfaces of the sensor interface. Some examples of suitable membrane systems are described in US Patent Application Publication No. 2005-0245799(A1).

[0129] Generally, the membrane system 32 includes multiple domains, such as one or more of an electrode domain 43, an interference domain 44, an enzyme domain 46 (e.g., including glucose oxidase), and a resistance domain 48, as shown in Figures 4B and 4C, and may include a high oxygen solubility domain, a bioprotective domain, and / or a cytocidal domain, as described in more detail in US Patent Application Publication No. 2005-0245799 (A1) and as described in more detail below. Although the embodiment illustrated in Figures 4E and 4F shows an interference domain between the electrode domain and the enzyme domain, the interference domain can be located more proximal or more distal to the electroactive surface. For example, in some embodiments, the interference domain 44 is more distal to the electroactive surface than the enzyme domain. In some embodiments, the interference domain is the most distal layer / domain of the membrane system to the electroactive surface. In some embodiments, the interference domain can be the most proximal domain / layer to the electroactive surface. In still other embodiments, the interference can be combined with one or more other membrane domains / layers. For example, in some embodiments, the interference domain and the resistance domain are combined into a single domain that provides both interference blocking and control of analyte flux. One of skill in the art will appreciate that a wide variety of configurations and combinations are encompassed by preferred embodiments.

[0130] The film system can be deposited on the exposed electroactive surfaces using known thin film techniques (e.g., evaporation, spraying, electrodeposition, dipping, etc.). However, in alternative embodiments, other deposition processes (e.g., physical and / or chemical vapor deposition processes) may be useful for providing one or more of the insulating and / or membrane layers, including, for example, ultrasonic deposition, electrostatic deposition, evaporation, deposition by sputtering, pulsed laser deposition, high rate oxygen fuel deposition, thermal evaporation, electron beam evaporation, deposition by reactive sputtering molecular beam epitaxy, atmospheric pressure chemical vapor deposition (CVD), atomic layer CVD, hot wire CVD, low pressure CVD, microwave plasma assisted CVD, plasma enhanced CVD, rapid thermal CVD, remote plasma enhanced CVD, ultra-high vacuum CVD, and ion implantation. However, the film system can be disposed over (or deposited on) the electroactive surfaces using any known method, as will be appreciated by those skilled in the art.

[0131] In some embodiments, one or more domains of the membrane system are formed from materials such as silicone, polytetrafluoroethylene, ethylene tetrafluoroethylene copolymers, polyolefins, polyesters, polycarbonates, biostable polytetrafluoroethylene, homopolymers, copolymers, terpolymers of polyurethane, polypropylene (PP), polyvinyl chloride (PVC), polyvinylidene fluoride (PVDF), polybutylene terephthalate (PBT), polymethyl methacrylate (PMMA), polyether ether ketone (PEEK), polyurethane, cellulosic polymers, polysulfones, and block copolymers thereof (e.g., diblock, triblock, alternating, random, and graft copolymers, etc.). US Patent Application Publication No. 2005-0245799 (A1) describes the configurations and materials of biointerfaces and membrane systems that can be applied to preferred embodiments.

[0132] The functionality of the domains of the membrane system 32 depends on a combination of factors, including, but not limited to, the thickness of the domain, the composition of the domain, the number of layers in the domain (or in the membrane as a whole), and the way in which the layers are applied (e.g., multiple thin layers can give better / complete coverage where one thick layer does not). In some embodiments, these factors are configured to provide a membrane system that causes the signal-to-noise ratio of the sensor to be substantially unaffected by non-stationary noise. The signal-to-noise ratio of the sensor can be substantially unaffected by non-stationary noise either by substantially increasing the analyte component (e.g., without a corresponding increase in the noise component) or by substantially reducing the noise component (e.g., without a substantial reduction in the analyte component).

[0133] Thus, in some preferred embodiments, the membrane system 32 is configured such that the signal-to-noise ratio of the sensor is substantially unaffected by non-stationary noise by substantially increasing the diffusion of glucose therein, while the diffusion of at least one interfering substance (e.g., H2O2 formed outside the membrane system) into the membrane system is substantially unaffected, and the analyte component is at least 80% of the total signal. Alternatively, or in addition, in some preferred embodiments, the membrane system is configured such that the signal-to-noise ratio of the sensor is substantially unaffected by non-stationary noise by reducing the noise component (e.g., non-stationary noise) of the total signal without a corresponding reduction in the analyte component. In some preferred embodiments, the membrane system is configured to increase the analyte signal component and reduce the non-stationary noise signal component.

[0134] In one preferred embodiment, the analyte sensor comprises a membrane system disposed on an electrode (e.g., working electrode), the membrane system configured for inactivation of electroactive species that may interfere with the analyte signal. In another preferred embodiment, the membrane system is configured to substantially consume at least one electroactive compound (that interferes with the analyte signal) that diffuses therein, such that the compound is substantially prevented from reaching the electroactive surface. The term "consumed" as used herein is a broad term and is used in its ordinary sense, including, but not limited to, rendering the interferent substantially unreactive with the electroactive surface and / or the voltage potential of the sensor, such as by oxidation or reduction of the interferent. For example, the membrane system may interact with the interferent such that, at the voltage potential at which the sensor operates, when the interferent comes into contact with the electroactive surface, the redox potential of the interferent is altered, such that the interferent cannot be substantially oxidized and / or reduced. In various embodiments, the membrane system is configured to substantially consume at least one interfering species by at least one of a tortuous diffusion path, a thickness of about 2 μm to about 100 μm or more, peroxidase, oxidase, catalase, and / or heme compound, etc., as described in more detail elsewhere herein. In some preferred embodiments, the membrane thickness is about 5 μm to about 50 μm.

[0135] f. Electrode domain In selected embodiments, the membrane system includes an electrode domain. The electrode domain 43 is provided to ensure that an electrochemical reaction occurs between the electroactive surface of the working electrode and the electroactive surface of the reference electrode, and thus the electrode domain 43 is preferably located closer to the electroactive surface than the interference domain and / or the enzyme domain. Preferably, the electrode domain includes a coating that maintains a layer of water on the electrochemically reactive surface of the sensor. In other words, the electrode domain exists to provide an environment between the surface of the working electrode and the surface of the reference electrode, facilitating the electrochemical reaction between the electrodes. For example, a humectant in a binder material can be used as the electrode domain. This allows for the complete transport of ions in an aqueous environment. The electrode domain can also help stabilize the operation of the sensor by accelerating the start-up of the electrodes and accelerating drift problems caused by insufficient electrolyte. The material forming the electrode domain can also provide an environment that protects against pH-mediated damage that may result from the formation of a large pH gradient due to the electrochemical activity of the electrodes.

[0136] In one embodiment, the electrode domain 43 comprises a flexible water-swellable hydrogel film having a "dry film" thickness of about 0.05 microns or less to about 20 microns or more, more preferably about 0.05, 0.1, 0.15, 0.2, 0.25, 0.3, 0.35, 0.4, 0.45, 0.5, 1, 1.5, 2, 2.5, 3, or 3.5 microns to about 4, 5, 6, 7, 8, 9, 10, 11, 12, 13, 14, 15, 16, 17, 18, 19, or 19.5 microns, and even more preferably about 3, 2.5, 2, or 1 microns or less to about 3.5, 4, 4.5, or 5 microns or more. "Dry film" thickness refers to the thickness of a cured film cast from a coating formulation by standard coating techniques.

[0137] In certain embodiments, the electrode domains 43 are formed from a curable mixture of a urethane polymer and a hydrophilic polymer. A particularly preferred coating is formed from a polyurethane polymer having carboxylate or hydroxyl functional groups and non-ionic hydrophilic polyether segments, where the polyurethane polymer is crosslinked with a water-soluble carbodiimide (e.g., 1-ethyl-3-(3-dimethylaminopropyl)carbodiimide (EDC)) in the presence of polyvinylpyrrolidone and cured at a moderate temperature of about 50° C.

[0138] In some preferred embodiments, the electrode domain 43 is formed from a hydrophilic polymer (e.g., a polyamide, a polylactone, a polyimide, a polylactam, a functionalized polyamide, a functionalized polylactone, a functionalized polyimide, a functionalized polylactam, or a combination thereof), which renders the electrode domain substantially more hydrophilic than the overlay domain (e.g., an interference domain, an enzyme domain). In some embodiments, the electrode domain is formed substantially entirely and / or primarily from a hydrophilic polymer. In some embodiments, the electrode domain is formed substantially entirely from PVP. In some embodiments, the electrode domain is formed entirely from a hydrophilic polymer. Useful hydrophilic polymers include, but are not limited to, poly-N-vinylpyrrolidone (PVP), poly-N-vinyl-2-piperidone, poly-N-vinyl-2-caprolactam, poly-N-vinyl-3-methyl-2-caprolactam, poly-N-vinyl-3-methyl-2-piperidone, poly-N-vinyl-4-methyl-2-piperidone, poly-N-vinyl-4-methyl-2-caprolactam, poly-N-vinyl-3-ethyl-2-pyrrolidone, poly-N-vinyl-4,5-dimethyl-2-pyrrolidone, polyvinyl imidazole, poly-N,N-dimethylacrylamide, polyvinyl alcohol (PVA), polyHEME, polymethyl methacrylate (PMMA), ethylene vinyl acetate (EVA), and the like. acetate, EVA), PGA-PEG, polyanhydrides, polyacrylic acid, polyethylene oxide, poly-2-ethyl-oxazoline, copolymers thereof and mixtures thereof. In some embodiments, blends of two or more hydrophilic polymers are preferred. In some preferred embodiments, the hydrophilic polymer is not crosslinked. In alternative embodiments, crosslinking is preferred, such as by adding a crosslinking agent (e.g., but not limited to, EDC) or by irradiation at a wavelength sufficient to promote crosslinking between hydrophilic polymer molecules, which is believed to create a more tortuous diffusion path through the domains.

[0139] Electrode domains formed from hydrophilic and / or conductive polymers (e.g., PVP and buffers) have been shown to substantially reduce the break-in time of analyte sensors, such as glucose sensors utilizing cellulosic interference domains as described in more detail elsewhere herein. In some embodiments, single-component electrode domains formed from a single hydrophilic polymer (e.g., PVP) have been shown to substantially reduce the break-in time of glucose sensors to less than about 2 hours, less than about 1 hour, less than about 20 minutes, and / or substantially immediately. Generally, sensor break-in is the amount of time required (after implantation) for the sensor signal to become substantially representative of the analyte concentration. Sensor break-in includes both membrane break-in and electrochemical break-in. In some embodiments, the break-in time is less than about 2 hours. In other embodiments, the break-in time is less than about 1 hour. In still other embodiments, the break-in time is less than about 30 minutes, less than about 20 minutes, less than about 15 minutes, less than about 10 minutes, or less. In preferred embodiments, sensor break-in occurs substantially immediately. Advantageously, in embodiments where the break-in time is about 0 minutes (substantially immediately), for example, the sensor can be inserted and begin providing substantially accurate analyte (e.g., glucose) concentrations almost immediately after insertion, and membrane break-in does not limit start-up time.

[0140] Without being bound by theory, it is believed that providing an electrode domain that is then substantially more hydrophilic than the more distal membrane layer or domain (e.g., the overlay domain). A layer (e.g., interference domain or enzyme domain) more distal to the electroactive surface than the electrode domain decreases the break-in time of the implanted sensor by increasing the rate at which the membrane system is hydrated by the surrounding host tissue. Without being bound by theory, it is generally believed that increasing the amount of hydrophilicity of the electrode domain relative to the overlay layer (e.g., the distal layer in contact with the electrode domain, such as the interference domain, enzyme domain, etc.) increases the rate of water absorption, thereby decreasing the break-in time of the sensor. The hydrophilicity of the electrode domains can be substantially increased based on their hydrophilicity relative to each other and to the overlay layer (e.g., the cellulosic interference domain) by appropriate selection of hydrophilic polymers, with preferred polymers being substantially more hydrophilic than the overlay layer. In one exemplary embodiment, PVP forms the electrode domain and the interference domain is formed from a blend of cellulose derivatives (e.g., but not limited to, cellulose acetate butyrate and cellulose acetate). Because PVP is substantially more hydrophilic than the cellulosic interference domain, it is believed that PVP quickly draws water into the membrane relative to the electrode domain, allowing the sensor to function with the desired sensitivity and accuracy and begin within a substantially reduced period of time after implantation. The reduction in sensor break-in time reduces the amount of time a host must wait to obtain a sensor reading, which is particularly advantageous in ambulatory applications, as well as in hospital environments where time is of the essence. In some alternative embodiments, a hydrophilic (e.g., PVP) electrode domain can be formed beneath the silicone-Pluronic polymer blend interference domain, such that the break-in time of the sensor is substantially reduced.

[0141] Without wishing to be bound by theory, it is believed that when the water uptake of an overlay domain (e.g., an overlay domain of an electrode domain) is less than that of the electrode domain (e.g., during membrane equilibration), the difference in water uptake between the two domains promotes membrane equilibration and therefore drives membrane break-in. That is, it is believed that increasing the difference in hydrophilicity (e.g., between the two domains) results in an increase in the rate of water uptake, which in turn results in a decrease in membrane break-in time and / or sensor break-in time.

[0142] Generally, the molecular weight of an analyte is substantially higher than the molecular weight of many electroactive species (e.g., reactive oxygen and nitrogen species) that may interfere with the analyte signal. As an example, the molecular weight of glucose is 180 g / mol, and the molecular weight of H2O2 is 34.02 g / mol. The diffusion of molecules through a membrane system is substantially regulated by the porosity (e.g., pore size) and hydrophilicity of the membrane. Small molecules such as H2O2, and reactive oxygen and nitrogen species, etc., generally diffuse through many membrane systems at substantially their maximum rate (e.g., due to their small size), so that changes in many membrane system configurations have substantially little effect on their diffusion rate. In contrast, the diffusion of an analyte (e.g., glucose) through a membrane system is substantially slower due to its large size. It is believed that increasing the diffusion rate (through a membrane system) of an analyte does not substantially affect the diffusion rate of the small molecules mentioned above. However, increasing the difference in hydrophilicity (e.g., between the two domains) substantially increases the diffusion rate of the analyte without a corresponding increase in the diffusion of small molecules, so that the analyte component (e.g., of the total signal) increases while the noise component (e.g., of the total signal) remains substantially unchanged, thereby adjusting (e.g., increasing) the signal-to-noise ratio. The adjusted signal-to-noise ratio resulting from the inclusion of a hydrophilic electrode domain (e.g., 2×PVP in some embodiments) increases the sensor sensitivity and reduces the sensor error. This advantageously adjusts the signal-to-noise ratio of the sensor in which the domain is located.

[0143] As described elsewhere herein, the relative hydrophilicity of the electrode domain 43 compared to the overlay domain can be adjusted by selecting a more hydrophilic material for the formation of the electrode domain (and / or a more hydrophobic material for the overlay domain). For example, an electrode domain having a hydrophilic polymer capable of absorbing a greater amount of water can be selected instead of a second hydrophilic polymer capable of absorbing less water than a first hydrophilic polymer. In some embodiments, the difference in water content between the electrode domain and the overlay domain (e.g., during or after membrane equilibration) is about 1% or less to about 90% or more. In other embodiments, the difference in water content between the electrode domain and the overlay domain is about 10% or less to about 80% or more. In still other embodiments, the difference in water content between the electrode domain and the overlay domain is about 30% or less to about 60% or more. In preferred embodiments, the electrode domain absorbs from 5% to 95% or more water by weight than an adjacent (overlay) domain (e.g., a domain that is more distal to the electroactive surface than the electrode domain), preferably from 5, 10, 15, 20, 25, 30, 35, 40, 45, or 50% to about 55, 60, 65, 70, 75, 80, 85, 90, or 95% water by weight.

[0144] In another example, the rate of water absorption by a polymer can be influenced by other factors, such as, but not limited to, the molecular weight of the polymer. For example, the rate of water absorption by PVP depends on its molecular weight, typically about 40 kDa or less to about 360 kDa or more. Lower molecular weight PVP (e.g., 40 kDa) absorbs water faster than higher molecular weight PVP. Thus, modulating factors such as molecular weight that affect the rate of water absorption by a polymer can facilitate the appropriate selection of materials for electrode domain fabrication. In one embodiment, a lower molecular weight PVP is selected to reduce break-in time.

[0145] Preferably, the electrode domains are deposited by known thin film deposition techniques (e.g., spray coating or dip coating of the electroactive surface of the sensor). In some embodiments, the electrode domains are formed by dip coating the electroactive surface in an electrode domain solution (e.g., 5, 10, 15, 20, 25, or 30% or more PVP in deionized water) and curing the domain for about 15 minutes to about 30 minutes at a temperature of about 40° C. to about 55° C. (and can be accomplished under vacuum (e.g., 20-30 mmHg)). In embodiments where dip coating is used to deposit the electrode domains, a preferred insertion rate of about 1 inch to about 3 inches / minute into the electrode domain solution, a preferred residence time of about 0.5 minutes to about 2 minutes in the electrode domain solution, and a preferred withdrawal rate of about 0.25 inches to about 2 inches / minute from the electrode domain solution provide a functional coating. However, as will be appreciated by those skilled in the art, values ​​other than those listed above may be acceptable or even desirable in certain embodiments, depending, for example, on solution viscosity and solution surface tension. In one embodiment, the electroactive surface of the electrode system is dip coated once (one layer) and cured at 50° C. under vacuum for 20 minutes. In another embodiment, the electroactive surface of the electrode system is dip coated and cured a first time at 50° C. under vacuum for 20 minutes, followed by dip coating and curing a second time at 50° C. under vacuum for 20 minutes (two layers). In yet other embodiments, the electroactive surface can be dip coated three or more times (three or more layers). In other embodiments, one, two, three or more layers of PVP are applied to the electroactive surface by spray coating or vapor deposition. In some embodiments, a crosslinking agent (e.g., EDC) can be added to the electrode domain casting solution to promote crosslinking within the domains (e.g., between electrode domain polymeric components, latex, etc.). However, in some alternative embodiments, no crosslinking agent is used and the electrode domains are not substantially crosslinked.

[0146] In some embodiments, the deposited PVP electrode domain 43 has a "dry film" thickness of about 0.05 microns or less to about 20 microns or more, more preferably about 0.05, 0.1, 0.15, 0.2, 0.25, 0.3, 0.35, 0.4, 0.45, 0.5, 1, 1.5, 2, 2.5, 3, or 3.5 microns to about 4, 5, 6, 7, 8, 9, 10, 11, 12, 13, 14, 15, 16, 17, 18, 19, or 19.5 microns, and even more preferably about 2, 2.5 or 3 microns to about 3.5, 4, 4.5, or 5 microns.

[0147] Although separate electrode domains 43 are described herein, in some embodiments, sufficient hydrophilicity can be provided in the interference domain and / or enzyme domain (domains adjacent to the electroactive surface) to provide complete transport of ions in an aqueous environment (e.g., without a separate electrode domain). In these embodiments, an electrode domain is not required.

[0148] g. Interference domain In some embodiments, the membrane system 34 comprises an interference domain 44 configured to substantially reduce and / or block the diffusion of one or more noise-causing interferents into the membrane system, thereby increasing the signal-to-noise ratio of the sensor. In some embodiments, the interference domain 44 is a component of the membrane system as shown in Figures 4E and 4F. However, the interference domain can be located at any level of the membrane system (e.g., layer or domain) (e.g., closer or more distal to the electroactive surface than shown in Figures 4E and 4F). In some other embodiments, the interference domain is combined with an additional membrane domain (e.g., a resistance domain or an enzyme domain).

[0149] As explained elsewhere herein, noise may occur during periods of inactivity, such as the first few hours or days after sensor implantation (e.g., intermittent sedimentary noise), and is believed to be caused by a local increase in interferents (e.g., electroactive metabolites) that interfere with sensor function, resulting in an apparent glucose signal that is generally independent of the host's glucose concentration. Without being bound by theory, it is believed that the noise intensity and / or the number of occurrences of intermittent sedimentary noise can be reduced or eliminated by reducing the local concentration of interferents, such as by incorporating an interference domain 44 in the membrane system 34. In general, the term "interference domain" includes any noise reduction mechanism that substantially blocks, reduces, eliminates, reacts with, or prevents interferents from reacting at the working electrode. As used herein, "noise reduction mechanism" is a broad term and should be given its ordinary and customary meaning to one of skill in the art (and should not be limited to any special or customized meaning), and refers to, without limitation, any sensor system component or configuration that reduces and / or eliminates noise on the sensor signal. Some noise reduction mechanisms include, but are not limited to, electrode configurations (e.g., two or more working electrodes), membrane configurations (e.g., interference domains), algorithm configurations (e.g., signal processing to remove identified noise components of the signal), and the like. In addition, the noise reduction mechanisms described herein, including structures, membrane materials, bioactive agents, and the like that can reduce the effect of interference species (noise) on the sensor signal, can be considered at least part of an "interference domain." Some examples of interference domain structures are described herein in this section entitled "Interference Domain." However, other known interference domain structures can be implemented with the sensors described herein. Although the embodiment shown in FIG. 4E and FIG. 4F shows an interference domain 44 located between the electrode domain and the enzyme domain, the interference domain can be located at any level of the membrane system (e.g., more proximal or more distal to the electroactive surface). For example, the interference domain can be located between the enzyme domain and the resistor domain, between the electroactive surface and the electrode domain, as the outermost membrane domain, and the like.In some embodiments, any domain of the membrane system may be configured to function as an interference domain or may be combined with an interference domain, for example, an enzyme domain and an interference domain may be combined into an enzyme-interference domain that performs the functions of an enzyme domain and an interference domain.

[0150] In one preferred embodiment, the membrane system includes an interference domain configured to substantially reduce noise (e.g., non-stationary noise) caused by one or more endogenous or exogenous interferents. In a preferred embodiment, the signal-to-noise ratio can be adjusted (e.g., increased) by incorporating an interference domain of a preferred embodiment onto the sensor. In some preferred embodiments, the interference domain is configured such that the analyte component is at least about 80% of the total signal over a period of at least about one day. In some preferred embodiments, the interference domain is configured such that the non-stationary noise component is less than about 20% of the total signal over a period of at least about one day.

[0151] As shown in Figures 4E and 4F, the membrane system 32 of the preferred embodiment includes an interference domain 44. In some preferred embodiments, an interference domain is provided that substantially restricts or blocks the flow of one or more interfering species therethrough. In some embodiments, the interference domain may be configured to reduce noise (and adjust the signal-to-noise ratio) using one, two, or more noise-reduction mechanisms. For example, in some embodiments, the interference domain is configured to substantially block the passage of at least one interfering species into the membrane system. In some embodiments, the interference domain is configured to substantially reduce the concentration of at least one interferent. For example, the interferent may be diluted, such as by increasing the bulk of the fluid and / or promoting the formation of a fluid pocket around the sensor. Alternatively, or in addition, the concentration of the interferent may be substantially reduced by configuring the interference domain to increase bulk fluid flow (e.g., carrying the interferent away via the lymphatic system). In other embodiments, the interference domain is configured to oxidize and / or reduce the interferent such that it no longer substantially affects the sensor. In some embodiments, the interference domain is configured to reduce non-stationary noise (and adjust the signal-to-noise ratio) by combining two or more noise reduction mechanisms, as described below. As described in more detail herein, some known interference species for glucose sensors include acetaminophen, ascorbic acid, bilirubin, cholesterol, creatinine, dopamine, ephedrine, ibuprofen, L-dopa, methyldopa, salicylate, tetracycline, tolazamide, tolbutamide, triglycerides, and uric acid. In some embodiments, the interference domain of the preferred embodiment is less permeable to one or more interference species than the species being measured (e.g., the product of the enzymatic reaction being measured on the electroactive surface), such as, but not limited to, H2O2.

[0152] In one embodiment, the interference domain 44 is formed from one or more cellulose derivatives. Cellulose derivatives can include, but are not limited to, cellulose esters and cellulose ethers. Generally, cellulose derivatives include polymers such as cellulose acetate, cellulose acetate butyrate, 2-hydroxyethyl cellulose, cellulose acetate phthalate, cellulose acetate propionate, cellulose acetate trimellitate, and copolymers and terpolymers thereof with other cellulose or non-cellulosic monomers. Cellulose is a polysaccharide polymer of beta-D-glucose. Although cellulose derivatives are generally preferred, other polymeric polysaccharides having similar properties to cellulose derivatives can also be used in preferred embodiments.

[0153] In one preferred embodiment, the interference domain 44 is formed from cellulose acetate butyrate. A cellulose acetate butyrate having a molecular weight of about 10,000 Daltons to about 75,000 Daltons, preferably about 15,000, 20,000, or 25,000 Daltons to about 50,000, 55,000, 60,000, 65,000, or 70,000 Daltons, more preferably about 20,000 Daltons, is used. However, in certain embodiments, higher or lower molecular weights may be preferred. In some embodiments, blends of two or more cellulose acetate butyrates having different molecular weights are preferred. Although "blends" as defined herein (compositions of two or more substances that are not substantially chemically bound to each other and can be separated) are generally preferred, in certain embodiments, a single polymer incorporating different components (e.g., separate components as monomer units and / or substituents on a single polymer chain) may be used instead. Further, a casting solution or dispersion of cellulose acetate butyrate with a weight percent of about 5% to about 25%, preferably about 5%, 6%, 7%, 8%, 9%, 10%, 11%, 12%, 13%, 14% or 15% to about 16%, 17%, 18%, 19%, 20%, 21%, 22%, 23%, 24% or 25%, more preferably about 5% to about 15% is preferred. Preferably, the casting solution includes a solvent or solvent system, such as an acetone:ethanol solvent system. In certain embodiments, higher or lower concentrations may be preferred. In alternative embodiments, a single solvent (e.g., acetone) is used to form symmetric membrane domains. A single solvent is used in the casting solution to form the symmetric membrane layers. In some embodiments, multiple layers of cellulose acetate butyrate can be advantageously combined to form interference domains, for example, three layers can be used. It may be desirable to use a mixture of cellulose acetate butyrate components having different molecular weights in a single solution, or to deposit multiple layers of cellulose acetate butyrate from different solutions containing cellulose acetate butyrates of different molecular weights, different concentrations, and / or different chemical properties (e.g., functional groups). It may also be desirable to include additional materials in the casting solution or dispersion, such as functionalizing agents, crosslinking agents, other polymeric materials, materials that can modify the hydrophilicity / hydrophobicity of the resulting layer, etc.

[0154] In an alternative embodiment, the interference domain 44 is formed from cellulose acetate. Preferred are cellulose acetates having a molecular weight of about 30,000 daltons or less to about 100,000 daltons or more, preferably about 35,000, 40,000, or 45,000 daltons to about 55,000, 60,000, 65,000, 70,000, 75,000, 80,000, 85,000, 90,000, or 95,000 daltons, more preferably about 50,000 daltons. In some embodiments, blends of two or more cellulose acetates having different molecular weights are preferred. Further, a casting solution or dispersion of cellulose acetate with a weight percent of about 3% to about 10%, preferably about 3.5%, 4.0%, 4.5%, 5.0%, 5.5%, 6.0%, or 6.5% to about 7.5%, 8.0%, 8.5%, 9.0%, or 9.5%, more preferably about 8%, is preferred. However, in certain embodiments, higher or lower molecular weights and / or weight percents of cellulose acetate may be preferred. It may be desirable to use a mixture of cellulose acetates with molecular weights in a single solution, or to deposit multiple layers of cellulose acetate from different solutions containing cellulose acetates of different molecular weights, different concentrations, or different chemical properties (e.g., functional groups). It may also be desirable to include additional materials in the casting solution or dispersion, such as those described in more detail above.

[0155] In addition to forming interference domains from only cellulose acetate or only cellulose acetate butyrate, the interference domain 44 may be formed from combinations or blends of cellulose derivatives (such as, but not limited to, cellulose acetate and cellulose acetate butyrate), or combinations of layers of cellulose acetate and layers of cellulose acetate butyrate. In some embodiments, the blend of cellulose derivatives (for forming the interference domain) comprises up to about 10% or more by weight of cellulose acetate. For example, in some embodiments, about 1, 2, 3, 4, 5, 6, 7, 8, 9% or more by weight of cellulose acetate is preferred. In some embodiments, the blend of cellulose derivatives comprises about 90% or less to about 100% by weight of cellulose acetate butyrate. For example, in some embodiments, the blend comprises about 91, 92, 93, 94, 95, 96, 97, 98, or 99% by weight of cellulose acetate butyrate. In some embodiments, the cellulosic blend comprises about 1.5, 2.0, 2.5, 3.0, or 3.5% cellulose acetate by weight to about 98.5, 98.0, 97.5, 97.0, or 96.5% cellulose acetate butyrate by weight. In other embodiments, the blend comprises about 4, 4.5, 5, 5.5, 6, 6.5, 7, 7.5, or 8% cellulose acetate by weight to about 96, 95.5, 95, 94.5, 94, 93.3, 93, 92.5, or 92% cellulose acetate butyrate by weight. In yet other embodiments, the blend comprises about 8.5, 9.0, 9.5, 10.0, 10.5, or 11.0% cellulose acetate by weight to about 91.5, 91.0, 90.5, 90, 89.5, or 89% cellulose acetate butyrate by weight.

[0156] In some embodiments, preferred blends of cellulose acetate and cellulose acetate butyrate contain from about 1.5 parts or less to about 60 parts or more of cellulose acetate butyrate per part of cellulose acetate. In some embodiments, the blend contains from about 2 parts to about 40 parts of cellulose acetate butyrate per part of cellulose acetate. In other embodiments, about 4, 6, 8, 10, 12, 14, 16, 18, or 20 parts of cellulose acetate butyrate per part of cellulose acetate are preferred for forming interference domains 26. In still other embodiments, blends having 22, 24, 26, 28, 30, 32, 34, 36, or 38 parts of cellulose acetate butyrate per part of cellulose acetate are preferred. As explained elsewhere herein, cellulose acetate butyrate is relatively more hydrophobic than cellulose acetate. Thus, cellulose acetate / cellulose acetate butyrate blends contain substantially more hydrophobic components than hydrophilic components.

[0157] Cellulose acetate butyrate is a cellulose polymer that has both acetyl and butyl groups in addition to hydroxyl groups. The acetyl group is more hydrophilic than the butyl group, and the hydroxyl group is more hydrophilic than both the acetyl and butyl groups. Thus, the relative amounts of acetyl, butyl, and hydroxyl groups can be used to adjust the hydrophilicity / hydrophobicity of the cellulose acetate butyrate of the cellulose acetate / cellulose acetate butyrate blend. The cellulose acetate butyrate can be selected based on the relative amounts of acetate, butyrate, and hydroxyl compounds, and the cellulose acetate can be selected based on the relative amounts of acetate and hydroxyl compounds. For example, in some embodiments, cellulose acetate butyrate having about 35% or less acetyl groups, about 10% to about 25% butyl groups, and the remaining hydroxyl groups is preferred for forming interference domains 44. In other embodiments, cellulose acetate butyrate having about 25% to about 34% acetyl groups and about 15% to about 20% butyl groups is preferred. In yet another embodiment, the preferred cellulose acetate butyrate contains about 28% to about 30% acetyl groups and about 16% to about 18% butyl groups. In yet another embodiment, the cellulose acetate butyrate can have no acetate groups and about 20% to about 60% butyrate groups. In yet another embodiment, the cellulose acetate butyrate can have about 55% butyrate groups and no acetate groups.

[0158] While asymmetric interference domains can be used in some alternative embodiments, in some embodiments, symmetric interference domains 44 (e.g., of a cellulose derivative blend, such as, but not limited to, a blend of cellulose acetate and cellulose acetate butyrate components) are preferred. Symmetric membranes are uniform throughout their structure, e.g., have no gradient in pore density or size, or have a skin on one side but not the other. In various embodiments, symmetric interference domains can be formed by appropriate selection of a solvent to make the casting solution (e.g., no antisolvent is used). Suitable solvents include those belonging to the ketone family that can solvate cellulose acetate and cellulose acetate butyrate. Solvents include, but are not limited to, acetone, methyl ethyl ketone, methyl n-propyl ketone, cyclohexanone, and diacetone alcohol. In some embodiments, other solvents, such as furans (e.g., tetrahydrofuran and 1,4-dioxane), may be preferred. In one exemplary embodiment, about 7% to about 9% by weight solids (e.g., a blend of cellulose derivatives (e.g., cellulose acetate and cellulose acetate butyrate)) are blended with a single solvent (e.g., acetone) to form a casting solution for a symmetric interference domain. In another embodiment, about 10% to about 15% solids are blended with acetone to form the casting solution. In yet another embodiment, about 16% to about 18% solids are blended with acetone to form the casting solution. In some embodiments, relatively low or high weight percent solids are preferred to form the casting solution.

[0159] The casting solution can be applied either directly to the electroactive surface of the sensor or on top of the electrode domain layer (if included in the membrane system). The casting solution can be applied using any known thin film technique, as described elsewhere herein. Additionally, in various embodiments, the symmetric interference domain 44 includes at least one layer, and in some embodiments, two, three, or more layers are formed by successive application and curing of the casting solution.

[0160] The concentration of solids in the casting solution can be adjusted to deposit a sufficient amount of solids on the electrode in one layer (e.g., one dip or spray) to form a membrane layer with sufficient blocking capacity such that the equivalent glucose signal of an interferent measured by the sensor (e.g., a compound with an oxidation or reduction potential (e.g., HO) that overlaps with that of the species being measured) is about 60 mg / dL or less. For example, in some embodiments, the percentage of solids in the casting solution is adjusted such that only a single layer (e.g., one dip) is required to deposit a sufficient amount of cellulose acetate / cellulose acetate butyrate blend to form a functional symmetric interference domain that substantially blocks the passage of at least one interferent, such as, but not limited to, acetaminophen, ascorbic acid, dopamine, ibuprofen, salicylic acid, tolbutamide, tetracycline, creatinine, uric acid, ephedrine, L-dopa, methyldopa, and tolazamide. In some embodiments, the amount of interference domain material deposited by a single dip is sufficient to reduce the equivalent glucose signal of the interferent (e.g., as measured by a sensor) to about 60 mg / dl or less. In a preferred embodiment, the equivalent glucose signal response of the interferent (as measured by a sensor) is 50 mg / dl or less. In a more preferred embodiment, the interferent produces an equivalent glucose signal response of 40 mg / dl or less. In even more preferred embodiments, the interferent produces an equivalent glucose signal response of less than about 30, 20, or 10 mg / dl. In one exemplary embodiment, the interference domain is configured to substantially block the passage of acetaminophen, and the equivalent glucose signal response of acetaminophen is less than about 30 mg / dl.

[0161] In alternative embodiments, the interference domain is configured to substantially block a therapeutic dose of acetaminophen. The term "therapeutic dose" as used herein is a broad term and should be given its ordinary and customary meaning to those skilled in the art (and should not be limited to a special or customized meaning), and refers to the amount of any substance required to effect a cure of a disease, to relieve pain, or to correct the symptoms of a deficiency of a particular factor in the diet, without limitation, the effective dose used with a therapeutically applied compound, such as a drug. For example, a therapeutic dose of acetaminophen can be the amount of acetaminophen required to relieve a headache or reduce a fever. As a further example, 1,000 mg of acetaminophen taken orally, such as by swallowing two 500 mg tablets of acetaminophen, is a therapeutic dose frequently taken for headaches. In some embodiments, the interference membrane is configured to block a therapeutic dose of acetaminophen, and the equivalent glucose signal response of acetaminophen is less than about 60 mg / dl. In a preferred embodiment, the interference membrane is configured to block a therapeutic dose of acetaminophen, the equivalent glucose signal response of which is less than about 40 mg / dl. In a more preferred embodiment, the interference membrane is configured to block a therapeutic dose of acetaminophen, the equivalent glucose signal response of which is less than about 30 mg / dl.

[0162] Without being bound by theory, it is believed that for symmetrical cellulose-based membranes, there is an inverse balance between interferent blocking and analyte sensitivity. That is, in some embodiments, changes to the interference domain configuration that increase interferent blocking may result in a corresponding decrease in sensor sensitivity. Sensor sensitivity is described in more detail elsewhere herein. It is believed that the balance between interferent blocking and sensor sensitivity depends on the relative proportions of hydrophobic and hydrophilic components of the membrane layer (e.g., interference domain), with sensors having more hydrophobic interference domains increasing interferent blocking but decreasing sensitivity, and sensors having more hydrophilic interference domains decreasing interferent blocking but increasing sensitivity. It is believed that the hydrophobic and hydrophilic components of the interference domains can be balanced to promote a desired level of interferent blocking while simultaneously maintaining a desired level of analyte sensitivity. The hydrophobic-hydrophilic balance of the interference domain can be manipulated and / or maintained by appropriate selection and blending of hydrophilic and hydrophobic interference domain components (e.g., cellulose derivatives having acetyl, butyryl, propionyl, methoxy, ethoxy, propoxy, hydroxyl, carboxymethyl, and / or carboxyethyl groups). For example, cellulose acetate is relatively more hydrophilic than cellulose acetate butyrate. In some embodiments, increasing the percentage of cellulose acetate (or decreasing the percentage of cellulose acetate butyrate) can increase the hydrophilicity of the cellulose acetate / cellulose acetate butyrate blend, which promotes increased permeability to hydrophilic species, such as, but not limited to, glucose, H2O2, and some interferents (e.g., acetaminophen). In another embodiment, the percentage of cellulose acetate butyrate is increased to increase the blocking of interferents, but also reduces the lower permeability to some desired molecules, such as H2O2 and glucose.

[0163] One way to manipulate the hydrophobic-hydrophilic balance of the interference domain is to select the appropriate proportions of acetyl groups (which are relatively more hydrophilic than butyl groups), butyl groups (which are relatively more hydrophobic than acetyl groups), and hydroxyl groups on the cellulose acetate butyrate used to form the interference domain 44. For example, increasing the proportion of acetate groups on the cellulose acetate butyrate makes it more hydrophilic. In another example, increasing the proportion of butyl groups on the cellulose acetate butyrate makes it more hydrophobic. In yet another example, increasing the proportion of hydroxyl groups increases the hydrophilicity of the cellulose acetate butyrate. Thus, by selecting a cellulose acetate butyrate that is more or less hydrophilic (or more or less hydrophobic), the overall hydrophilicity of the cellulose acetate / cellulose acetate butyrate blend can be adjusted. In one exemplary embodiment, the interference domain can be configured to be relatively hydrophobic (and therefore more strongly block interferents) by decreasing the proportion of acetyl or hydroxyl groups or increasing the proportion of butyl groups on the cellulose acetate butyrate used in the casting solution (while maintaining the relative ratio of cellulose acetate to cellulose acetate butyrate).

[0164] In some alternative embodiments, the interference domain is formed of a blend of cellulose derivatives, and the hydrophilic and hydrophobic components of the interference domain are balanced such that the glucose sensitivity is between about 1 pA / mg / dL and about 100 pA / mg / dL, and the interference domain is sufficiently blocked from passing at least one interferent such that the equivalent glucose signal response of the at least one interferent is less than about 60 mg / dL. In preferred embodiments, the glucose sensitivity is between about 5 pA / mg / dL and about 25 pA / mg / dL. In more preferred embodiments, the glucose sensitivity is between about 5 pA / mg / dL and about 25 pA / mg / dL, and the equivalent glucose signal response of the at least one interferent is less than about 40 mg / dL. In even more preferred embodiments, the glucose sensitivity is between about 5 pA / mg / dL and about 25 pA / mg / dL, and the equivalent glucose signal response of the at least one interferent is less than about 30 mg / dL. In some embodiments, the balance between the hydrophilic and hydrophobic components of the interference domain can be achieved by adjusting the amounts of the hydrophilic and hydrophobic components relative to each other, as well as adjusting the hydrophilic and hydrophobic groups (e.g., acetyl, butyryl, propionyl, methoxy, ethoxy, propoxy, hydroxyl, carboxymethyl, and / or carboxyethyl groups) of the components themselves (e.g., cellulose derivatives such as, but not limited to, cellulose acetate and cellulose acetate butyrate).

[0165] In some alternative embodiments, additional polymers such as Nafion® can be used in combination with the cellulose derivatives to provide equivalent and / or enhanced functionality of the interference domain 44. As an example, a layer of 5 wt. % Nafion® casting solution was applied onto a previously applied (e.g., cured) layer of 8 wt. % cellulose acetate, for example by dip-coating at least one layer of cellulose acetate, followed by dip-coating at least one layer of Nafion® onto a needle-type sensor as described with reference to the preferred embodiment. Any number of coatings or layers formed in any order may be suitable for forming the interference domain of the preferred embodiment.

[0166] In some alternative embodiments, more than one cellulose derivative may be used to form the interference domain 44 of the preferred embodiment. Generally, the formation of interference domains on a surface utilizes a solvent or solvent system to solvate the cellulose derivative (or other polymer) prior to film formation thereon. In a preferred embodiment, acetone and ethanol are used as solvents for cellulose acetate. However, those skilled in the art will recognize numerous solvents suitable for use with cellulose derivatives (and other polymers). Furthermore, those skilled in the art will recognize that the preferred relative amounts of solvents may depend on the cellulose derivative (or other polymer) used, its molecular weight, its deposition method, its desired thickness, and the like.

[0167] However, it is preferred to use a solute percentage of about 1% to about 25% by weight to form an interference domain solution to obtain an interference domain having the desired properties. The cellulose derivative (or other polymer) used, its molecular weight, the deposition method, and the desired thickness can be adjusted depending on one or more other parameters and can be varied accordingly, as will be appreciated by those skilled in the art.

[0168] In some alternative embodiments, other polymer types that can be utilized for the interference domain 44 include, for example, polyurethanes, polymers with pendant ionic groups, and polymers with controlled pore sizes. In one such alternative embodiment, the interference domain includes a thin hydrophobic membrane that is non-swelling and limits the diffusion of high molecular weight species. The interference domain is permeable to relatively low molecular weight substances such as hydrogen peroxide, but limits the passage of higher molecular weight substances, including glucose and ascorbic acid. Other systems and methods for reducing or eliminating interference species that can be applied to the membrane system of the preferred embodiment are described in U.S. Pat. No. 7,074,307, U.S. Pat. App. Pub. No. 2005-0176136 (A1), U.S. Pat. No. 7,081,195, and U.S. Pat. App. Pub. No. 2005-0143635 (A1). In some alternative embodiments, a separate interference domain is not included.

[0169] In some embodiments, the interference domain 44 is deposited directly on the electroactive surface of the sensor or on the distal surface of the electrode domain with a domain thickness of about 0.05 microns or less to about 20 microns or more, more preferably about 0.05, 0.1, 0.15, 0.2, 0.25, 0.3, 0.35, 0.4, 0.45, 0.5, 1, 1.5, 2, 2.5, 3, or 3.5 microns to about 4, 5, 6, 7, 8, 9, 10, 11, 12, 13, 14, 15, 16, 17, 18, 19, or 19.5 microns, and even more preferably about 1, 1.5 or 2 microns to about 2.5 or 3 microns.

[0170] Thicker membranes may also be desirable in certain embodiments, but thinner membranes are generally preferred because they have less effect on the rate of diffusion of hydrogen peroxide from the enzyme membrane to the electrode. In some embodiments, the interference domain may be deposited either proximally or more distally to the electroactive surface than the electrode domain, depending on the composition of the interference domain and the configuration of the membrane system.

[0171] Generally, the membrane system of the preferred embodiments can be formed and / or deposited on the exposed electroactive surfaces (e.g., one or more of the working and reference electrodes) using known thin film techniques (e.g., casting, spray coating, drawdown, electrodeposition, dip coating, etc.), although casting or other known application techniques can also be utilized. Preferably, the interference domain is deposited by spray coating or dip coating. In one exemplary embodiment of a needle-type (transdermal) sensor as described herein, the interference domain is formed by dip coating the sensor in an interference domain solution using an insertion speed of about 0.5 inches / minute to about 60 inches / minute, preferably 1 inch / minute, a dwell time of about 0 minutes to about 2 minutes, preferably about 1 minute, and a withdrawal speed of about 0.5 inches / minute to about 60 inches / minute, preferably about 1 inch / minute, and curing (drying) the domain for about 1 minute to about 30 minutes, preferably about 3 minutes to about 15 minutes (which can be accomplished at room temperature or under vacuum (e.g., 20-30 mmHg)). In one exemplary embodiment including a cellulose acetate butyrate interference domain, a 3 minute cure (i.e., drying) time is preferred between each layer applied. In another exemplary embodiment using a cellulose acetate interference domain, a 15 minute cure (i.e., drying) time is preferred between each layer applied.

[0172] In some embodiments, the dipping process can be repeated at least once and up to 10 or more times. In other embodiments, only one dipping is preferred. The preferred number of repeated dipping processes depends on the cellulose derivatives used, their concentrations, conditions during deposition (e.g., dipping), and the desired thickness (e.g., a thickness sufficient to provide functional blocking of a particular interferent), etc. In some embodiments, 1-3 microns may be preferred as the thickness of the interference domain. However, as will be appreciated by those of skill in the art, values ​​other than these may be acceptable or even desirable in certain embodiments, depending, for example, on viscosity and surface tension. In one exemplary embodiment, the interference domain is formed from three layers of cellulose acetate butyrate. In another exemplary embodiment, the interference domain is formed from ten layers of cellulose acetate. In another embodiment, the interference domain is formed from one layer of a blend of cellulose acetate and cellulose acetate butyrate. In alternative embodiments, the interference domain may be formed using any known method and combination of cellulose acetate and cellulose acetate butyrate, as will be appreciated by those of skill in the art.

[0173] In some embodiments, the electroactive surface can be washed prior to application of the interference domain 44. In some embodiments, the interference domains of preferred embodiments can be useful as bioprotective or biocompatible domains (i.e., domains that interact with host tissue when implanted in an animal (e.g., human) due to their stability and biocompatibility).

[0174] In some embodiments, the interference domain is formed from a silicone-hydrophilic / hydrophobic polymer blend, such as, but not limited to, a silicone-pluronic polymer blend, as described in the section entitled "Silicone / Hydrophilic Polymer Blend Materials."

[0175] h enzyme domain In a preferred embodiment, the membrane system further comprises an enzyme domain 46 located distally from the electroactive surface than the interference domain. However, other configurations may be desirable (FIGS. 4E-4F). In a preferred embodiment, the enzyme domain provides an enzyme for catalyzing the reaction of the analyte and its co-reactant, as described in more detail below. In a preferred embodiment of a glucose sensor, the enzyme domain comprises glucose oxidase. However, other oxidases, such as galactose oxidase or uricase oxidase, may also be used.

[0176] For an enzyme-based electrochemical glucose sensor to perform well, it is preferred that the sensor response is not limited by either the enzyme activity or the coreactant concentration. Since enzymes, including glucose oxidase (GOX), undergo inactivation as a function of time even in ambient conditions, this behavior is compensated for when forming the enzyme domain. Preferably, the enzyme domain is composed of an aqueous dispersion of colloidal polyurethane polymer containing the enzyme. However, in an alternative embodiment, the enzyme domain is constructed from an oxygen-enhancing material (e.g., silicone or fluorocarbon) to provide an excess supply of oxygen during transient ischemia. Preferably, the enzyme is immobilized within the domain. See, for example, US Patent Application Publication No. 2005-0054909 (A1).

[0177] In preferred embodiments, the enzyme domain 46 is deposited on the interference domain 44 at a domain thickness of about 0.05 microns or less to about 20 microns or more, more preferably about 0.05, 0.1, 0.15, 0.2, 0.25, 0.3, 0.35, 0.4, 0.45, 0.5, 1, 1.5, 2, 2.5, 3, or 3.5 microns to about 4, 5, 6, 7, 8, 9, 10, 11, 12, 13, 14, 15, 16, 17, 18, 19, or 19.5 microns, and even more preferably about 2, 2.5, or 3 microns to about 3.5, 4, 4.5, or 5 microns. However, in some embodiments, the enzyme domain can be deposited directly on the electroactive surface. Preferably, the enzyme domain is deposited by spray coating or dip coating. In one embodiment of a needle-type (transdermal) sensor as described herein, the enzyme domain is formed by dip coating the interference domain-coated sensor in an enzyme domain solution and curing the domain for about 15 minutes to about 30 minutes at a temperature of about 40° C. to about 55° C. (and can be accomplished under vacuum (e.g., 20-30 mmHg)). In an embodiment where dip coating is used to deposit the enzyme domain at room temperature, a preferred insertion speed of about 0.25 in / min to about 3 in / min, a preferred residence time of about 0.5 minutes to about 2 minutes, and a preferred withdrawal speed of about 0.25 in / min to about 2 in / min provides a functional coating. However, as will be appreciated by those skilled in the art, values ​​other than those listed above may be acceptable or even desirable in certain embodiments, depending, for example, on viscosity and surface tension. In one embodiment, the enzyme domain is formed by dip coating twice in the enzyme domain solution (i.e., forming two layers) and curing under vacuum at 50° C. for 20 minutes. However, in some embodiments, the enzyme domain may be formed by dip coating and / or spray coating one or more layers at a predetermined concentration, insertion rate, residence time, withdrawal rate, and / or desired thickness of the coating solution.

[0178] 5A is a cross-sectional view of a membrane system in one embodiment, showing the diffusion distance D1 between H2O2 generated in the enzyme domain and the electroactive surface of the electrode 38. In general, when H2O2 310 is generated (in the enzyme domain 46) by the metabolism of glucose by GOX, the generated H2O2 can diffuse in all directions (e.g., from the location in the enzyme domain where H2O2 was generated). A portion of the generated H2O2 diffuses the distance D1 to the electroactive surface and generates a signal associated with the analyte (e.g., FIG. 5A).

[0179] Figure 5B is a cross-sectional view of a membrane system in another embodiment, showing the diffusion distance D2 between the H2O2 generated at the enzyme domain and the electroactive surface. In this embodiment, the distance D2 between the location of H2O2 generation and the electroactive surface is reduced (relative to D.sub.1). Thus, more H2O2 reaches the electroactive surface and is detected in the embodiment of Figure 5B compared to the embodiment of Figure 5A.

[0180] FIG. 5C is a cross-sectional view of a membrane system in yet another embodiment, showing the diffusion distance D3 between the enzyme domain and the H2O2 generated at the electroactive surface. That is, in the embodiment shown in FIG. 5C, the distance D3 between the location of H2O2 generation and the electroactive surface is reduced (relative to D1 and D2). Thus, more H2O2 (relative to the embodiment of FIG. 5A and FIG. 5B) contacts the electroactive surface and is detected in the embodiment of FIG. 5C. Thus, in a preferred embodiment, the system is configured such that the analyte component of the signal is at least 80% of the total signal, at least in part, due to the preferred H2O2 diffusion distance. In some embodiments, the preferred diffusion distance is achieved by including GOX in a layer adjacent to the electrode (i.e., in a H2O2 diffusion-based sensor). In some embodiments, the preferred diffusion distance is less than about 20 μm. In some preferred embodiments, the preferred diffusion distance is less than about 10 μm. In some preferred embodiments, the preferred diffusion distance is less than about 5 μm. In yet other preferred embodiments, the preferred diffusion distance is less than about 1 μm. In a preferred embodiment, the analyte component is at least 80% of the total signal over a period of at least one day.

[0181] In some embodiments, the enzyme domain 46 is located adjacent to the electroactive surface (e.g., by eliminating or combining the functionality of the electrode and / or interference domain). In some embodiments, the enzyme (e.g., GOX) may be contained within the electrode domain, for example, using a coupling agent. Suitable coupling agents include disulfosuccinimidyl tartarate (sulfo-DST), bis(sulfosuccinimidyl) suberate (BS3), ethylene glycol bis(sulfosuccinimidyl succinate) (Sulfo-EGS), 3,3'-dithiobis(sulfosuccinimidyl propionate) (Sulfo-EGS), ... propionate), DTSSP), N,N'1,3-phenylenedimaleimide (mPDM), N,N'-1,2-phenylenedimaleimide (oPDM), N,N'-1,4-phenylenedimaleimide (pPDM), N,N'-(methylene-4-1-phenylene)bismaleimide (BM), naphthalene-1,5-dimaleimide (naphthalene ne-1,5-dimaleimide (NDM), bismaleimidoethane (BMOE), 1,4-bismaleimidobutane (BMB), 1,4-bis-maleimidyl-2,3-dihydroxybutane (BMDB), dithio-bis-maleimdoethane (DTME), 1,6-bismaleimidohexane (BMH), 1,8-bismaleimidotriethylene glycol (1,8-bismaleimidotriethyleneglycol (BM[PEO]3), 1,11-bis-maleimidotetraethyleneglycol (BM[PEO]4), dimethyl adipimidate (DMA), dimethyl pimelimidiate (DMP), dimethyl suberimidate (DMS), dimethyl 3,3'-dithiobis-propionimidate (DTBP), disuccinimidyl tartarate (DST), disuccinimidyl glutarate (DSG), dithiobis(succinimidylpropionate) (DSP), disuccinimidyl suberate suberate (DSS), bis(2-[succinimidooxycarbonyloxy]ethyl)sulfone (BSOCOES), ethylene glycol bis-(succinimidylsuccinate) (EGS), 1,5-difluoro-2,4-dinitrobenzene (DFDNB), 4,4'-difluoro-3,3'-dinitrodiphenylsulfone (4,4'-difluoro-3,Examples of suitable enzymes include, but are not limited to, 3'-dinitrodiphenylsulfone (DFDNPS), dibromobimane (bBBr), and the like. In some embodiments, the enzyme (e.g., GOX) can be absorbed (using techniques known in the art) to the electroactive surface, for example, by dipping the electrode in an enzyme solution, drying the electrode, and then applying at least one membrane domain. In yet other embodiments, the enzyme can be mixed with the electrode domain material (e.g., PVP or other hydrophilic polymer) and applied to the electroactive surface to form the electrode domain. Additionally, the thickness of the enzyme domain itself can be adjusted to increase the analyte component of the signal. In some embodiments, the enzyme domain has a thickness of about 10 μm or less. In some preferred embodiments, the enzyme domain has a thickness of about 5 μm or less. In more preferred embodiments, the enzyme domain has a thickness of about 2 μm or less.

[0182] In some embodiments, the enzyme domain is configured to modulate H2O2 utilization and / or production, for example, by including a coenzyme in the enzyme domain or in a membrane domain more proximal to the electroactive surface than the enzyme domain. In some circumstances, the coenzyme can stabilize the enzyme reaction product (e.g., H2O2 from the metabolism of glucose by GOX) and / or increase the reaction efficiency of the enzyme. For example, NADPH in the presence of other enzyme systems dramatically increases the effectiveness of the enzyme. Suitable coenzymes include, but are not limited to, superoxide dismutase (SOD), hydrogenase, reductase, oxidase, peroxidase, flavoenzyme, and NADPH. For example, the reaction product (e.g., H2O2) can be stabilized by a compound such as SOD, which can scavenge more reactive oxygen radical species and enhance the lifetime of the reaction product (H2O2). Stabilization of the reaction product and / or modulation of the enzyme reaction rate can result in a corresponding increase in the analyte signal. Thus, in some embodiments, the sensor is configured such that (after full sensor break-in) the analyte signal is at least 80% of the full signal over a period of at least one day. In some preferred embodiments, the analyte signal is at least 90% of the full signal over a period of at least two days.

[0183] In some embodiments, the enzymatic electrochemical analyte sensor includes at least one working electrode (including an electroactive surface) and a membrane system (including an enzyme domain) configured such that the enzyme domain is substantially adjacent to the electroactive surface. Additionally, the sensor is configured to detect H2O2 diffusing from its synthesis location (e.g., within the enzyme domain) to the electroactive surface (after sensor break-in is complete), such that the analyte component is at least 80% of the total signal over a period of at least one day. In preferred embodiments, the sensor is configured such that the non-stationary noise component does not substantially contribute to the total signal. For example, in some preferred embodiments, the non-stationary noise component is less than 20% of the total signal over a period of one day or more.

[0184] iResistance Domain In a preferred embodiment, the membrane system 32 includes a resistance domain 48 disposed more distally from the electroactive surface than the enzyme domain 46 (e.g., FIGS. 4E-4F). The following description is directed to a resistance domain for a glucose sensor, but the resistance domain can be similarly modified for other analytes and co-reactants.

[0185] There is a molar excess of glucose relative to the amount of oxygen in the blood, i.e., for each free oxygen molecule in the extracellular fluid, there are typically more than 100 glucose molecules (see Updike et al., Diabetes Care 5:207-21 (1982)). However, immobilized enzyme-based glucose sensors that use oxygen as a coreactant are preferably supplied with a non-rate-limiting excess of oxygen so that the sensor responds linearly to changes in glucose concentration but not to changes in oxygen concentration. Specifically, when the glucose monitoring reaction is oxygen-limited, linearity is not achieved above a minimum concentration of glucose. In the absence of a semi-permeable membrane placed over the enzyme domain to control the flux of glucose and oxygen, a linear response to glucose levels can only be obtained for glucose concentrations up to about 40 mg / dL. However, in a clinical setting, a linear response to glucose levels is desirable up to at least about 400 mg / dL.

[0186] The resistance domain 48 includes a semi-permeable membrane that controls the flow of oxygen and glucose to the underlying enzyme domain 46, preferably making the oxygen non-rate-limiting excess. As a result, the upper linearity limit of glucose measurement is extended to values ​​much higher than that achieved without the resistance domain. In one embodiment, the resistance domain exhibits an oxygen to glucose permeability ratio of about 50:1 or less to about 400:1 or more, preferably about 200:1. As a result, one-dimensional reactant diffusion is sufficient to provide an excess of oxygen at all reasonable glucose and oxygen concentrations found in the subcutaneous matrix (see Rhodes et al., Anal. Chem., 66:1520-1529 (1994)).

[0187] In an alternative embodiment, a lower oxygen to glucose ratio may be sufficient to provide excess oxygen by using a high oxygen solubility domain (e.g., silicone or fluorocarbon based materials or domains) to enhance oxygen supply / transport to the enzyme domain. Also, if more oxygen is supplied to the enzyme, more glucose can be supplied to the enzyme without creating an oxygen-limiting excess. In an alternative embodiment, the resistance domain is formed from a silicone composition such as described in US Patent Application Publication No. 2005-0090607(A1).

[0188] j. Polyurethane polymer materials In a preferred embodiment, the resistance domain comprises a polyurethane membrane having both hydrophilic and hydrophobic regions to control the diffusion of glucose and oxygen to the analyte sensor, which membrane is easily and reproducibly manufactured from commercially available materials. A suitable hydrophobic polymer component is polyurethane, or polyetherurethaneurea. Polyurethane is a polymer produced by the condensation reaction of a diisocyanate with a difunctional hydroxyl-containing material. Polyurethaneurea is a polymer produced by the condensation reaction of a diisocyanate with a difunctional amine-containing material. Preferred diisocyanates include aliphatic diisocyanates containing about 4 to about 8 methylene units. Diisocyanates containing alicyclic moieties may also be useful in preparing the polymer and copolymer components of the membrane of the preferred embodiment. The material forming the base of the hydrophobic matrix of the resistance domain may be any known in the art as suitable for use as a membrane in a sensor device and as having sufficient permeability to allow relevant compounds to pass therethrough, for example, sufficient permeability to allow oxygen molecules to pass through the membrane from the test sample to reach an active enzyme or an electrochemical electrode. Examples of materials that can be used to make non-polyurethane type membranes include inorganic polymers such as vinyl polymers, polyethers, polyesters, polyamides, polysiloxanes and polycarbosiloxanes, natural polymers such as cellulosic and protein-based materials, and mixtures or combinations thereof.

[0189] In a preferred embodiment, the hydrophilic polymer component is polyethylene oxide. For example, one useful hydrophobic-hydrophilic copolymer component is a polyurethane polymer containing about 5% hydrophilic polyethylene oxide. The polyethylene oxide portion of the copolymer is thermodynamically driven to separate from the hydrophobic portion of the copolymer and the hydrophobic polymer component. The 5% polyethylene oxide-based soft segment portion of the copolymer used to form the final blend affects the water absorption and subsequent glucose permeability of the membrane.

[0190] In some embodiments, alternative polyurethanes containing urethane groups and polyurethane ureas (also containing urea groups) can be used to construct the resistance domain and modulate the flux of analytes (e.g., glucose) therethrough, preferably increasing the signal-to-noise ratio. For example, the polyurethanes and polyurethane ureas selected to form the resistance domain can be based on poly(oxyalkylene) glycols, including poly(oxyethylene) glycols. In accordance with conventional usage, both types of polymers are referred to as polyurethanes herein. Poly(oxyalkylene) glycol-based polyurethane membranes do not show a predictable relationship between molecular weight and permeability. The unique separation observed in the membranes of the present invention can be explained on the basis that the substance-membrane or solute-membrane interactions that tend to affect partitioning are not solely due to the hydrophilic poly(oxyalkylene) glycol or "soft" segments, but the hydrophobic or "hard" segments of the block copolymer also contribute to the overall selectivity. Thus, the selectivity of the membrane system can be altered by changing the structure of the hydrophobic segments of the block copolymer and / or by increasing or decreasing the molecular weight of the poly(oxyalkylene) glycol. In some embodiments of the membrane system, the desired selectivity for glucose and hydrogen peroxide can be obtained by using two different membranes, for example, block copolyether urethanes based on poly(oxyalkylene) glycols. Further description of polyurethane resistance domains can be found in PCT International Publication No. WO1992 / 013271.

[0191] In a preferred embodiment, the resistance domain is deposited on the enzyme domain to yield a thickness of about 0.05 microns or less to about 20 microns or more, more preferably about 0.05, 0.1, 0.15, 0.2, 0.25, 0.3, 0.35, 0.4, 0.45, 0.5, 1, 1.5, 2, 2.5, 3, or 3.5 microns to about 4, 5, 6, 7, 8, 9, 10, 11, 12, 13, 14, 15, 16, 17, 18, 19, or 19.5 microns, and even more preferably about 2, 2.5, or 3 microns to about 3.5, 4, 4.5, or 5 microns. Preferably, the resistance domain is deposited on the enzyme domain by vapor deposition, spray coating, or dip coating. In one preferred embodiment, spray coating is the preferred deposition technique. The spray process atomizes and mists the solution so that most or all of the solvent evaporates before the coating material settles onto the underlying domains, thereby minimizing contact of the solvent with the enzyme.

[0192] In a preferred embodiment, the resistance domain is deposited on the enzyme domain by spray coating a solution of about 1% to about 5% by weight of polymer and about 95% to about 99% by weight of solvent. When spraying a solution of resistance domain material containing a solvent onto the enzyme domain, it is desirable to mitigate or substantially reduce any contact with the enzyme of any solvent in the spray solution that may inactivate the enzyme underlying the enzyme domain. Tetrahydrofuran (THF) is one solvent that has minimal or negligible effect on the enzyme of the enzyme domain when sprayed. Other solvents may also be suitable for use, as will be appreciated by those skilled in the art.

[0193] It is believed that the incorporation of silicone-hydrophilic polymer blends into the membrane system can render the signal-to-noise ratio of the sensor substantially unaffected by non-stationary noise by substantially reducing and / or eliminating the noise, for example, by substantially blocking and / or slowing the passage of interfering substances therethrough (e.g., reducing the diffusion rate). In preferred embodiments, a sensor having one or more working electrodes includes a membrane system 34, and a resistive domain 48 includes a blend of silicone polymers and hydrophilic polymers configured to reduce noise-causing species, such as non-stationary noise-causing species. In some embodiments, the membrane domain / layer includes a blend of silicone polymers and hydrophilic polymers configured to reduce noise-causing species. In some preferred embodiments, the sensor includes a silicone-hydrophilic polymer blend membrane domain and / or layer (e.g., interference domain) having a micelle jacket structure (described elsewhere herein). Without wishing to be bound by theory, it is believed that membrane domains comprising silicone-hydrophilic polymer blends may reduce noise by blocking and / or inhibiting the passage of at least one interfering species into the membrane system while simultaneously allowing and / or facilitating the transport of an analyte (e.g., glucose or other such water-soluble molecules such as drugs).

[0194] By "hydrophilic polymer" is meant that the polymer has an affinity for water and is generally primarily soluble in water or has a tendency to absorb water due to the presence of one or more hydrophilic substituents. In one example, the hydrophilic component of the hydrophilic polymer facilitates the movement of water and / or compounds in water (e.g., by diffusion or other means) through a membrane formed from the hydrophilic polymer, such as by lowering the thermodynamic barrier to the movement of compounds in water through the membrane.

[0195] In some embodiments, the hydrophilic polymer comprises a hydrophilic-hydrophobic polymer. In general, the terms "hydrophilic-hydrophobic" and "hydrophobic-hydrophilic" are used interchangeably herein (not meant to imply that either the hydrophilic or hydrophobic substituents are the majority of the polymer) and refer to the property of having both hydrophilic and hydrophobic substituents and / or characteristics in a single molecule (e.g., polymer).

[0196] The hydrophilic and hydrophobic substituents of the polymer can affect the behavior of the polymer in certain situations (e.g., but not limited to, silicone / hydrophilic-hydrophobic blend materials and micelle jackets), which are described elsewhere herein. Using PEO-PPO-PEO as an exemplary polymer, the main component of the polymer (PEO) is hydrophilic and can provide an overall hydrophilic character to the molecule (e.g., the molecule behaves generally hydrophilically). However, the hydrophobic component of the polymer (PPO) can allow the polymer to have some hydrophobic properties (e.g., a portion of the molecule behaves like a hydrophobic molecule) in some situations. In some situations, such as the formation of a micelle jacket in a silicone / hydrophilic-hydrophobic blend material, the polymer self-assembles to the silicone (e.g., silicone globules) such that the hydrophobic PPO is adjacent to the silicone (which is hydrophobic) and the two PEO groups protrude away from the silicone (e.g., due to thermodynamic forces). Depending on the situation (e.g., the polymer selected), variations of the above micelle jacket structure (e.g., opposite orientation) are possible. For example, in a mixture of PPO-PEO-PPO and silicone, the PPO groups will orient themselves towards the silicone and the PEO centers will orient away from the silicone.

[0197] In one embodiment, the hydrophilic polymer has a molecular weight of at least about 1000 g / mol, 5,000 g / mol, 8,000 g / mol, 10,000 g / mol, or 15,000 g / mol or more. In one embodiment, the hydrophilic polymer comprises both hydrophilic and partially hydrophobic domains (e.g., copolymers, also referred to herein as hydrophobic-hydrophilic polymers). The hydrophobic domains facilitate the blending of hydrophilic polymers with hydrophobic silicone polymers, such as, but not limited to, the formation of micellar jackets within and / or around the silicone. In one embodiment, the hydrophobic domains are themselves polymers (i.e., polymeric hydrophobic domains). For example, in one embodiment, the hydrophobic domains are not simple molecular head groups, but rather polymers. In various embodiments, the molecular weight of any covalently connected hydrophobic domain within the hydrophilic polymer is at least about 500 g / mol, 700 g / mol, 1000 g / mol, 2000 g / mol, 5000 g / mol, or 8,000 g / mol or more. In various embodiments, the molecular weight of any covalently connected hydrophilic domain within the hydrophilic polymer is at least about 500 g / mol, 700 g / mol, 1000 g / mol, 2000 g / mol, 5000 g / mol, or 8,000 g / mol or more.

[0198] In some embodiments, within a particular layer, the ratio of silicone polymer to hydrophilic polymer is selected to provide an amount of solubility of oxygen and water-soluble molecules so that the transport of oxygen and water-soluble molecules through the domain is adjusted according to the desired function of that particular layer.Furthermore, in some embodiments, the ratio of silicone polymer to hydrophilic polymer and the polymer composition are selected so that a layer constructed from this material has interference properties that inhibit the transport of one or more interfering species through the layer.Some known interfering species for glucose sensors include, but are not limited to, acetaminophen, ascorbic acid, bilirubin, cholesterol, creatinine, dopamine, ephedrine, ibuprofen, L-dopa, methyldopa, salicylate, tetracycline, tolazamide, tolbutamide, triglycerides, and uric acid.Thus, in some embodiments, the silicone polymer / hydrophilic polymer layer disclosed herein is less permeable to one or more of these interfering species than the analyte, e.g., glucose.

[0199] In some of these embodiments, the ratio of silicone polymer to hydrophilic polymer (in the layer incorporating the blend) varies according to the desired functionality of each layer.The relative amounts of silicone polymer and hydrophilic polymer described below are based on the respective amounts found in the cured polymer blend.When introduced into an aqueous environment, some of the polymer components may leach out, thereby changing the relative amounts of silicone polymer and hydrophilic polymer.For example, a substantial amount of the portion of the hydrophilic polymer that is not crosslinked may leach out, depending, for example, on the molecular weight of the hydrophilic polymer and how tortuous the diffusion path from the membrane is.

[0200] In some embodiments, the silicone and hydrophilic polymer form a substantial blend. That is, the amount of any crosslinking between the silicone polymer and the hydrophilic polymer is substantially limited. In various embodiments, at least about 75%, 85%, 95%, or 99% or more of the silicone polymer can be covalently bonded to the hydrophilic polymer. In some embodiments, the silicone polymer and the hydrophilic polymer are not crosslinked at all unless a crosslinking agent is used (e.g., as described below). Similarly, in some embodiments, the amount of any entanglement (e.g., blending at a molecular level) between the silicone polymer and the hydrophilic polymer is substantially limited. In one embodiment, the silicone polymer and the hydrophilic polymer form a microdomain. For example, in one embodiment, the silicone polymer forms a micellar jacket structure surrounded by a network of the hydrophilic polymer.

[0201] The silicone polymer for use in the silicone / hydrophilic polymer blend can be any suitable silicone polymer. In some embodiments, the silicone polymer is a liquid silicone rubber that can be vulcanized using metal (e.g., platinum), peroxide, heat, ultraviolet light, or other radiation catalyzed processes. In some embodiments, the silicone polymer is a dimethyl- and methylhydrogen-siloxane copolymer. In some embodiments, the copolymer has vinyl substituents. In some embodiments, commercially available silicone polymers can be used. For example, commercially available silicone polymer precursor compositions may be used to prepare the blends as described below. In one embodiment, MED-4840, commercially available from NUSIL® Technology LLC, is used as the precursor to the silicone polymer used in the blend. MED-4840 is composed of a two-component silicone elastomer precursor that includes a vinyl-functionalized dimethyl- and methylhydrogen-siloxane copolymer, amorphous silica, a platinum catalyst, a crosslinker, and an inhibitor. The two components can be mixed together and heated to initiate vulcanization, thereby forming an elastomeric solid material. Other suitable silicone polymer precursor systems include, but are not limited to, MED-2174 peroxide-cured liquid silicone rubber available from NUSIL® Technology LLC, SILASTIC® MDX4-4210 platinum-cured biomedical grade elastomer available from DOW CORNING®, and Implant Grade Liquid Silicone Polymer (durometer 10-50) available from Applied Silicone Corporation.

[0202] The hydrophilic polymer for use in the blend can be any suitable hydrophilic polymer, including, but not limited to, components such as polyvinylpyrrolidone (PVP), polyhydroxyethyl methacrylate, polyvinyl alcohol, polyacrylic acid, polyethers such as polyethylene glycol or polypropylene oxide, and copolymers thereof, including, for example, diblock, triblock, alternating, random, comb, star, dendritic, and graft copolymers (block copolymers are described in U.S. Pat. Nos. 4,803,243 and 4,686,044, which are incorporated herein by reference). In one embodiment, the hydrophilic polymer is a copolymer of poly(ethylene oxide) (PEO) and poly(propylene oxide) (PPO). Suitable such polymers include, but are not limited to, PEO-PPO diblock copolymers, PPO-PEO-PPO triblock copolymers, PEO-PPO-PEO triblock copolymers, alternating block copolymers of PEO-PPO, random copolymers of ethylene oxide and propylene oxide, and blends thereof. In some embodiments, the copolymers may be optionally substituted with hydroxy substituents. Commercially available examples of PEO and PPO copolymers include PLURONIC brand polymers available from BASF. Some PLURONIC polymers are HO--(CH2CH2O) x --(CH2CH2CH2O) y --(CH2CH2O) x --OH It is a triblock copolymer of poly(ethylene oxide)-poly(propylene oxide)-poly(ethylene oxide) having the general molecular structure:

[0203] Here, the repeat units x and y vary among various PLURONIC® products. The poly(ethylene oxide) blocks act as hydrophilic domains that allow dissolution of aqueous drugs in the polymer. The poly(propylene oxide) blocks act as hydrophobic domains that facilitate blending of the PLURONIC® polymer with silicone polymers. In one embodiment, PLURONIC® F-127 is used, which has an x ​​of about 100 and a y of about 65. The molecular weight of PLURONIC® F-127 is about 12,600 g / mol, as reported by the manufacturer. Other PLURONIC® polymers include PPO-PEO-PPO triblock copolymers (e.g., PLURONIC® R products) and PEO-PDMS-PEO triblock copolymers (e.g., PEO-polydimethylsiloxane-PEO, SILSURF®, BASF, USA). Other suitable commercially available polymers include, but are not limited to, SYNPERONICS® products available from UNIQEMA®.

[0204] The polyether structure of the PLURONIC® polymer is relatively inert, and therefore, without being bound to any particular theory, it is believed that the PLURONIC® polymer does not substantially react with the components in MED-4840 or other silicone polymer precursors.

[0205] Those skilled in the art will appreciate that other copolymers having hydrophilic and hydrophobic domains may be used. For example, in another embodiment, a triblock copolymer having a hydrophobic-hydrophilic-hydrophobic structure may be used. In another alternative embodiment, a diblock copolymer having a hydrophilic-hydrophobic structure is used. Additional devices, methods, and compositions are described in U.S. Patent Application Publication No. 2006-0270923(A1) and U.S. Patent Application No. 11 / 404,417, filed April 14, 2006.

[0206] Layers and / or domains comprising silicone polymer-hydrophilic polymer blends can be made using any of the methods known in the art for forming polymer blends. In one embodiment, a silicone polymer precursor (e.g., MED-4840) is mixed with a solution of a hydrophilic polymer (e.g., PLURONIC® F-127 dissolved in a suitable solvent such as acetone, ethyl alcohol, xylene, or 2-butanone). The mixture may then be drawn into a film or applied to a multi-layer thin film structure using any method known in the art (e.g., spraying, painting, dip coating, deposition, molding, 3D printing, lithographic techniques (e.g., photolithography), micro- and nano-pipetting printing techniques, etc.). The mixture can then be cured at elevated temperatures (e.g., 50°C to 150°C). Other suitable curing methods include, for example, ultraviolet radiation or gamma radiation. During curing, the silicone polymer precursor vulcanizes and the solvent evaporates.

[0207] In one embodiment, after the mixture is stretched into a film, another preformed layer of the membrane system is placed on the film. Then, curing of the film provides a bond between the film and the other preformed layer. In one embodiment, the preformed layer is a cell disruption layer. In one embodiment, the cell disruption domain comprises a preformed porous silicone membrane. In other embodiments, the cell disruption domain is also formed from a silicone polymer / hydrophilic polymer blend. In some embodiments, multiple films are applied on top of the preformed layer. Each film may have a finite interface with the adjacent film, or together may form a physically continuous structure with a chemical composition gradient.

[0208] A certain amount of crosslinking agent can also be included in the mixture to induce crosslinking between hydrophilic polymer molecules. For example, when using PLURONIC® polymers, crosslinking can be induced using crosslinking systems that react with pendant or terminal hydroxyl groups or methylene, ethylene, or propylene hydrogen atoms. Non-limiting examples of suitable crosslinking agents include ethylene glycol diglycidyl ether (EGDE), poly(ethylene glycol) diglycidyl ether (PEGDE), or dicumyl peroxide (DCP). Without being bound to any particular theory, it is believed that at low concentrations, these crosslinking agents react primarily with the PLURONIC® polymer and, to a certain extent, can induce crosslinking in or between the PLURONIC® polymer and the silicone polymer. In one embodiment, sufficient crosslinker is added such that the ratio of crosslinker molecules to hydrophilic polymer molecules added when synthesizing the blend is about 10 to about 30 (e.g., about 15 to about 20). In one embodiment, about 0.5% to about 15% w / w (in one example, about 1% to about 10%) of crosslinker is added based on the total dry weight of crosslinker, silicone polymer, and hydrophilic polymer added when blending the components. In one embodiment, about 5% to about 30% of the dry component weight is PLURONIC® polymer.

[0209] In some embodiments, other agents can be added to the mixture to facilitate the formation of the blend. For example, a small amount of butylhydroxy toluene (BHT) (e.g., about 0.01% w / w) or other suitable antioxidant can be mixed with PLURONIC® to stabilize it.

[0210] In some alternative embodiments, precursors of both the silicone polymer and the hydrophilic polymer can be mixed prior to curing, such that polymerization of both the silicone polymer and the hydrophilic polymer occurs during curing, hi another embodiment, an already polymerized silicone polymer is mixed with the hydrophilic polymer, such that no significant polymerization occurs during curing.

[0211] Without wishing to be bound by theory, it is believed that micelle-like structures, referred to herein as micelle jacket structures, can be formed by combining certain hydrophobic polymers (e.g., silicones) with certain amphiphilic polymers (e.g., hydrophilic polymers such as PLURONIC® polymers), which, when substantially blended, create a mechanism by which glucose and other analytes are transported at a limited rate. One example of a limited rate is the diffusion of oxygen and glucose into a membrane at a ratio of 50:1 (50 oxygen molecules for every 1 glucose molecule). In a preferred embodiment, oxygen and glucose diffuse into the membrane at a limited rate of 100:1. In a more preferred embodiment, oxygen and glucose diffuse into the membrane at a limited rate of 200:1.

[0212] In the first mechanism of limited analyte transport, it is believed that the hydrophilic and hydrophobic components of PLURONIC®, together with the silicone, can promote the self-assembly of PLURONIC® molecules into a micellar jacket. The micellar jacket provides a continuous channel (e.g., a tortuous path) through the silicone through which the analyte migrates. For example, on a first side of a membrane / domain, glucose dissolves in the hydrophilic components of the micellar jacket (e.g., within the membrane / domain) and diffuses through the hydrophilic portion of the adjacent micellar jacket to reach the other side of the membrane / domain.

[0213] In the second mechanism of limited analyte transport, it is believed that the micelle jacket may provide a hydrophilic phase within the silicone membrane / domain structure. There is an energy barrier to the diffusion of the analyte (e.g., glucose) into the silicone. However, energetic thermodynamic forces (e.g., analyte concentration gradient) drive the analyte to pass across / through the membrane by "jumping" from one micelle jacket to another. For example, a glucose concentration gradient can provide the energy for a glucose molecule to enter a membrane domain or layer (e.g., a cell-impermeable domain formed from a substantial blend of silicone and PLURONIC®), enter the first micelle jacket, then "jump" to the next micelle jacket and continue until the molecule reaches the other side of the membrane domain / layer.

[0214] In one exemplary embodiment, a silicone-hydrophilic polymer (e.g., the hydrophilic polymer is an amphiphilic polymer, such as, but not limited to, PLURONIC®) blend is believed to promote the polymeric self-assembly of a micellar jacket covering colloidal silicone globules (e.g., silicone granules forming a three-dimensional continuous polymeric structure with silicone-silicone contacts between the hydrophilic polymer-coated silicone granules) within the membrane domain. The hydrophilic groups of the micelle jacket orient toward the silicone, and the hydrophobic portion of the polymer orients away from the silicone core of the structure. For example, in the case of a silicone globule covered with PLURONIC® (PEO-PPO-PEO), it is believed that the PLURONIC® molecule thermodynamically favors orienting itself such that the PPO "lies down" against the silicone and the PEO away from the silicone, e.g., bent in a U-shape. Reverse micelle jackets, e.g., in silicone (e.g., hydrophobic PPO faces outward toward the silicone and hydrophilic PEO faces inward), are also possible. Furthermore, the micelle jackets may not be in direct physical contact with one another, which provides a thermodynamic barrier for molecules to enter the membrane layer and move through / across the layer by energetically "jumping" from one micelle jacket to the next.

[0215] In addition to facilitating the passage of analytes through the membrane domain, the micellar jacket structure has been found to block the diffusion of small reactive oxygen and nitrogen interferents (e.g., H2O2, oxygen radicals, peroxynitrate, etc.) that can cause non-stationary noise. Without being bound by theory, it is believed that the micellar jacket structure sufficiently retards the diffusion of reactive oxygen and nitrogen interferents so that these molecules self-extinguish before reaching the electroactive surface. In contrast, it is believed that high molecular weight interferents (e.g., acetaminophen and ascorbate) do not reach the electroactive surface because they are sterically and / or thermodynamically blocked and / or trapped by the micellar jacket. Thus, in a preferred embodiment, the non-stationary noise produced by both low molecular weight interferents and high molecular weight interferents is substantially attenuated such that the non-stationary noise component of the signal is less than 20% of the total signal over a period of at least one day.

[0216] In one exemplary embodiment, the enzyme-based electrochemical sensor is configured to block non-stationary non-analyte-related noise-generating compounds and includes at least one working electrode and a membrane system (e.g., FIG. 4E and FIG. 4F) including at least one domain formed of a silicone-hydrophilic polymer blend with a micelle jacket structure. In some embodiments, the membrane system includes at least one additional domain, such as, but not limited to, an electrode domain, an interference domain, an enzyme domain, a resistance domain, and a cell destruction domain. In one preferred embodiment, the sensor includes a membrane system with a combined resistance-interference domain formed of a silicone-hydrophilic polymer blend (with a micelle jacket structure) and is configured to reduce non-stationary noise by regulating the flux of an analyte into the membrane system and blocking the passage of at least one interferent (e.g., acetaminophen) into the membrane system.

[0217] In some preferred embodiments, an analyte sensor (e.g., an enzyme-based electrochemical analyte sensor) includes at least one working electrode and a membrane system having a resistive domain configured to substantially consume and / or block at least one intermittent non-stationary noise-causing species (e.g., H.sub.2O.sub.2 from a source external to the sensor) produced by a host metabolic process, such that the signal contribution attributable to non-stationary non-analyte components is less than about 20% over a period of about 1 day or more. In more preferred embodiments, the non-stationary non-analyte components are less than about 20% over a period of about 1, 2, 3, 4, 5, 6, or 7 days or more. In some preferred embodiments, the non-stationary noise is less than about 18%, 16%, 14%, 12%, 10%, 8%, 6%, 5%, 4%, 3%, 2%, or 1% of the total signal for at least about 1 day.

[0218] In a preferred embodiment, the resistance domain is configured to consume and / or block (e.g., by rendering the interferents electrochemically inactive at the applied voltage of the sensor) small molecule interferents (e.g., but not limited to, externally generated H.sub.2O.sub.2, reactive oxygen and nitrogen species) caused by the host's metabolic processes. For example, reactive oxygen and nitrogen species (e.g., oxygen and nitrogen radicals) and externally generated H2O2 (e.g., from local immune cells that have invaded the sensor site as part of the wound healing process) are highly reactive and spontaneously react with anything they come into contact with, such as the material of the membrane system. Thus, as these small molecule interferents diffuse through the membrane system, only a portion of them continues to diffuse to the electroactive surface of the sensor. The remaining small molecule interferents generally impinge on and react with the matrix of the membrane system. When these interferents react with the matrix, they are generally oxidized / reduced, so that they no longer substantially react with the electroactive surface of the sensor at the applied voltage. For example, reactive oxygen species O2 2-can be oxidized to O2. Thus, by reducing the number of interfering substance molecules reaching the electroactive surface, the noise components can be reduced and the signal-to-noise ratio can be increased. Thus, the system of the preferred embodiment can provide a signal in which the substantially non-stationary non-analyte-related components do not substantially contribute to the signal. For example, in some embodiments, the non-stationary noise is less than about 20% of the total signal for at least about one day.

[0219] In some embodiments, the resistance domain is configured to provide a tortuous path for noise-causing compounds such that reactive electroactive species that may interfere with the analyte signal come into contact with the tortuous diffusion path and are thereby consumed (e.g., rendered electrochemically inactive at the electroactive surface at the applied voltage).

[0220] In some embodiments, the thickness of the resistance domain and / or membrane system is configured such that the non-stationary noise component of the signal is less than about 20% of the total signal over a period of at least about 1, 2, or 3 days or more due to self-quenching as some relatively unstable electroactive compounds diffuse therethrough. In some embodiments, the thickness of the resistance domain is about 1 μm to about 25 μm or more. In some embodiments, the thickness of the membrane system is about 5 μm to about 10 μm or more. In alternative embodiments, the resistance domain is configured to consume at least one interferent (e.g., reactive oxygen or nitrogen species, exogenous H2O2) by including a compound that binds and / or metabolizes the interferent such that the interferent is substantially electrochemically unreactive with the electroactive surface (at the applied voltage). In some embodiments, an enzyme (e.g., including but not limited to, a peroxidase (e.g., catalase, horseradish peroxidase, cytochrome c peroxidase, glutathione peroxidase, etc.)) is incorporated into the resistance domain. In one exemplary embodiment, the peroxidase disposed in the resistance domain can metabolize externally generated H2O2 (which diffuses into the membrane system) into water and molecular oxygen, which do not substantially interact with the electroactive surface of the sensor. Thus, substantially only H2O2 produced within the enzyme domain (e.g., from the metabolism of glucose by GOX) diffuses to the sensor electroactive surface and produces a signal. Thus, the desired thickness of the resistance domain and / or membrane system can be achieved by various known techniques, such as applying one or more additional layers of resistance domain material (e.g., two layers instead of one) during membrane system construction.

[0221] In some embodiments, the resistance domain comprises one or more heme compounds, which are known antioxidants that react with reactive interfering species (rendering the interferents unreactive with the electroactive surface) such that the non-stationary noise component is less than about 20% of the total signal over a period of about one day or more. Suitable heme compounds include, but are not limited to, hemin, metmyoglobin, hemoglobin, methemoglobin, and cytochrome c, desferoxamine, or are synthesized by partial modification and crosslinking to a polymer backbone.

[0222] In some embodiments, the signal-to-noise ratio can be increased by including an interferent scavenger in one or more layers of the membrane system. Depending on the nature of the interferent, the interferent scavenger can be incorporated into the membrane domain that is either more distal or more proximal to the electroactive surface than the enzyme domain. In some embodiments, the scavenger can be incorporated into the enzyme domain of the membrane. For example, some interferents are ionic and bind to ionic interferents.

[0223] Thus, by incorporating an interferent-removing ionic component, such as Nafion®, into one or more layers of a membrane system, in some embodiments, interferents having the same charge as the ionic component can be substantially blocked and / or delayed from diffusing through the membrane system, thus allowing fewer interferents to reach the electroactive surface and reducing noise.

[0224] The interferent-removing enzyme can be incorporated into one or more layers of the membrane system. Useful enzymes include, but are not limited to, peroxidases and / or oxidases. Generally, peroxidases catalyze the reduction of compounds using H2O2. Exemplary peroxidases include horseradish peroxidase, glutathione peroxidase, cytochrome C peroxidase, myeloperoxidase, and the like. Horseradish peroxidase is a preferred peroxidase because interferents present in biological fluids, such as ascorbate, urate, acetaminophen, bilirubin, and cysteine, are rapidly oxidized by hydrogen peroxide in the presence of horseradish peroxidase. Generally, oxidases catalyze the oxidation / reduction of compounds using molecular O2. Exemplary oxidases include glucose oxidase, monoamine oxidase, cytochrome P450 oxidase, NADPH oxidase, cytochrome C oxidase, xanthine oxidase, L-gulonolactone oxidase, lactate oxidase, lysyl oxidase, catalase, and the like. In some embodiments, peroxidases can be crosslinked to one or more membrane domains using known protein crosslinking techniques, including, but not limited to, glutaraldehyde crosslinking, NaIO4, oxidation of the enzyme oligosaccharide groups followed by coupling to a matrix. Some useful coupling methods are described in U.S. Pat. Nos. 5,262,305 and 5,356,786.

[0225] In one exemplary embodiment, peroxidase is incorporated into the distal membrane domain (e.g., above the enzyme domain) to scavenge H.sub.2O.sub.2 derived from external sources (e.g., macrophages during wound healing). In one exemplary embodiment, the distal membrane domain comprises horseradish peroxidase. Additional scavenge techniques are described in U.S. Patent No. 5,356,786, U.S. Patent No. 6,284,478, and U.S. Patent No. 7,003,341.

[0226] In some embodiments, non-stationary noise can be reduced by including one or more membrane domains with an interferent-blocking compound. A variety of interferent-blocking compounds can be used, such as, but not limited to, sulfonated polyethersulfone, polyaminophenol, or polypyrrole. In one embodiment, the membrane system contains 3-amino-phenol, which blocks the transport of acetaminophen while allowing the diffusion of H.sub.2O.sub.2. The interferent-blocking compound can be applied to the electrodes using any method known in the art, such as, but not limited to, immersion, spraying, electropolymerization, spin-coating, as described elsewhere herein. In one exemplary embodiment, the sensor is a glucose sensor including two working electrodes, where a solution of 3-amino-phenol is sprayed onto the working electrodes and dried prior to application of the membrane enzyme domains. In further embodiments, the sensor includes additional membrane layers. Further methods and devices can be found in U.S. Patent No. 7,120,483 to Russell, which is incorporated herein by reference in its entirety.

[0227] Other systems and methods for reducing or eliminating interfering species that may be applied to the membrane systems of the preferred embodiments are described in U.S. Patent Application Publication Nos. 2005-0115832 (A1), 2005-0176136 (A1), 2005-0161346 (A1), and 2005-0143635 (A1). In some alternative embodiments, no separate interference domain is included.

[0228] k.Outer hydrophilic surface In some embodiments, the membrane system is comprised of a hydrophilic outer surface (and / or domain) that is at least discontinuously hydrophilic, and in some embodiments may be continuously hydrophilic and configured to contact the host's tissue. The term "discontinuous hydrophilic surface (domain)" as used herein is a broad phrase and is to be given its ordinary and customary meaning to those skilled in the art (and is not to be limited to any special or customized meaning), and refers to a surface that includes, but is not limited to, some hydrophilic domains and some hydrophobic domains located thereon. Without being bound by theory, it is believed that the outer surface of the resistance domain is responsible for a substantial portion of the domain's analyte resistance capability. It is believed that the signal-to-noise ratio can be made substantially unaffected by non-stationary noise by providing at least a discontinuous hydrophilic outer surface, such as by surface treatment with a polymer having sufficient hydrophilic components to provide an outer surface with at least discontinuous hydrophilic properties, and / or by application of a distal discontinuous hydrophilic layer or domain formed from a polymer having substantial hydrophilic components. Such surface treatments and / or the outer surface of the discontinuous hydrophilic domains enable a sensor system having a total signal of at least about 80% analyte content.

[0229] In some embodiments, the discontinuous hydrophilic outer domain (e.g., either the resistive domain itself or applied in addition to the resistive domain in exemplary embodiments) having a surface with discontinuous hydrophilic components is configured such that the analyte component of the sensor signal is at least about 80% of the total signal for at least about 1 day. In some embodiments, the discontinuous hydrophilic outer domain comprises a silicone / hydrophilic polymer blend. For example, in some embodiments, the hydrophilic component of the silicone / hydrophilic polymer blend is at least about 5% by weight. In some preferred embodiments, the hydrophilic component of the silicone / hydrophilic polymer blend is at least about 10% by weight. In some preferred embodiments, the hydrophilic component of the silicone / hydrophilic polymer blend is at least about 15% by weight. In even more preferred embodiments, the hydrophilic component of the silicone / hydrophilic polymer blend is at least about 20% by weight. In yet other more preferred embodiments, the hydrophilic component of the silicone / hydrophilic polymer blend is at least about 25% by weight.

[0230] In some embodiments, the discontinuous hydrophilic outer domain comprises a surface-treated resistance domain, the outer surface of which is configured such that the hydrophilicity of the surface of the resistance domain is increased to such an extent that the analyte component is at least 80% of the total signal over at least about 1 day. In some embodiments, the discontinuous hydrophilic outer domain comprises a polyurethane-based resistance domain that is surface-treated with a polymer containing a hydrophilic moiety (e.g., a PEG compound, a PLURONIC® compound, or a substantially hydrophilic polyurethane compound). In some embodiments, a polymer solution is applied onto the (substantially cured) polyurethane-based resistance domain 48 to form the cell-impermeable domain 42, as described elsewhere herein. In some embodiments, the applied solution is comprised of PEG in acetone, for example, but not limited to, a 1%, 5%, 10%, 20%, 30%, 40% or more solution of PEG in acetone is applied. For example, in experiments using small geometry sensors (e.g., transdermal) that were dipped once in a 30% solution of PEG in acetone, the sensitivity of the treated sensors increased from about 20% to about 75% when compared to untreated sensors.

[0231] In some embodiments, the discontinuous hydrophilic outer domain (e.g., having a discontinuous hydrophilic outer surface) is formed from polyurethane. To adjust the glucose permeability through the polyurethane-based outer hydrophilic domain, the amount of hydrophilic component can be adjusted such that the analyte component of the total signal is at least about 80%. For example, in some embodiments, the analyte sensor includes a discontinuous hydrophilic outer domain formed from a blend of polyurethanes, where the proportion of the more hydrophilic polyurethane is selected such that the analyte component of the total signal is at least 80% or more. In some embodiments of the analyte sensor, the discontinuous hydrophilic outer domain is a resistive domain formed from a polyurethane blend that includes a sufficient proportion of hydrophilic polyurethane (e.g., polyurethane with hydrophilic groups, such as, but not limited to, PEG, PEO, PVP, etc.) to provide a signal in which the analyte component is at least 80% of the total signal, such as for a period of at least about 1 day. In some preferred embodiments, the hydrophilic component is at least about 5% or more of the discontinuous hydrophilic outer domain. In some embodiments, the hydrophilic component of the outer intermittent hydrophilic exterior domain is at least about 1%, 2%, 5%, 7%, 10%, 15%, 20%, or 25% or more of the polyurethane blend.

[0232] In yet other embodiments, the surface of the resistance domain is treated (e.g., coated) with a polymer having a sufficient number of hydrophilic moieties such that the analyte component is at least 80% of the total signal. In one exemplary embodiment, the polyurethane blend resistance domain is treated with a hydrophilic polymer (e.g., but not limited to, PLURONIC® (commercially available from BASF Corp., Florham Park, NJ, USA)) or a sufficiently hydrophilic polyurethane-based polymer. In another exemplary embodiment, the sensor is immersed in PEG or PVP to increase the glucose permeability of the resistance domain. Other known hydrophilic materials, such as those described in the section entitled "Silicone / Hydrophilic Polymer Materials," can be used to increase the glucose permeability of the resistance domain without substantially affecting the permeability of interferents. In some embodiments in which the surface of the resistance domain is treated with a hydrophilic polymer, the analyte component of the signal is at least 90% of the total signal. In preferred embodiments, the analyte component of the signal is at least 99% of the total signal. In preferred embodiments, the analyte component of the signal is at least 80% of the total signal for at least about one day or more.

[0233] These principles and / or resistance domain configurations can be used in a variety of other analyte sensors, such as, but not limited to, the sensors described in U.S. Pat. No. 6,721,587, U.S. Pat. No. 4,484,987, U.S. Pat. No. 4,671,288, U.S. Pat. No. 5,322,063, U.S. Pat. No. 6,654,625, U.S. Pat. No. 6,689,265, and U.S. Patent Application Publication No. 2003-0031699(A1).

[0234] Without being bound by theory, it is believed that non-stationary, non-analyte-related noise can be reduced by diluting and / or removing transient electroactive species that may interfere with the analyte signal, such as by increasing the fluid bulk (e.g., fluid pocket), increasing bulk fluid flow, and / or increasing the diffusivity around at least a portion of the sensor, such as the sensing portion of the sensor. Additionally, a physical spacer can reduce the effects of lymphatic pooling (e.g., accumulation of interfering electroactive species in tissue surrounding an implanted sensor) due to local compression (described elsewhere herein) by mechanically maintaining the fluid pocket. Because the spacer can maintain the fluid bulk around the sensor during local compression, the effects of increased interferent concentration can be inhibited or reduced, thereby reducing noise and facilitating improved sensor function. One preferred embodiment provides a device having an architecture that enables and / or promotes increased bulk fluid flow and / or increased fluid volume in an area surrounding at least a portion of an implanted sensor in vivo, which is believed to enable a sensor signal having a non-stationary noise component that is at least about 20% of the total signal over a period of at least one day.

[0235] Various structures can be incorporated into the sensor configuration to allow and / or facilitate (e.g., stimulate or facilitate) fluid bulk, bulk fluid flow, and / or diffusivity, such as by forming fluid pockets, which may enable a sensor signal with a non-stationary noise component of less than about 20% of the total signal (over about a day or more). These structures can include, but are not limited to, spacers, meshes, sloughing layers, roughened surfaces, machinable materials, nanoporous materials, shape memory materials, porous memory materials, self-assembling materials, crushable materials, biodegradable materials, combinations thereof, and the like. Structures that promote increased fluid bulk and / or increased bulk fluid flow may also include, but are not limited to, structures that promote fluid inflow or outflow (e.g., fluid inflow promoting architectures, fluid outflow promoting architectures), structures that promote vasodilation (e.g., vasodilatory architectures), structures that promote inflammation (e.g., inflammatory architectures), structures that promote wound healing or wound perpetuation (e.g., wound healing architectures and wound architectures, respectively), structures that promote angiogenesis (e.g., angiogenic architectures), structures that inhibit inflammation (e.g., anti-inflammatory architectures), or combinations thereof.

[0236] In some embodiments, the sensor includes a physical spacer disposed between the sensor and the surrounding tissue. The spacer allows a liquid sheath to form around at least a portion of the sensor, such as the area surrounding the electrodes. The fluid sheath can provide a fluid bulk that dilutes or buffers interferents while facilitating glucose and oxygen transport to the sensor. In some embodiments, the spacer is a mesh or optionally a fibrous structure. Suitable mesh materials are known in the art and include open meshes made from biocompatible materials (e.g., but not limited to, PLA, PGA, PP, nylon, etc.). The mesh spacer can be applied directly to the sensing mechanism or on a biointerface membrane, such as the porous biointerface membranes disclosed elsewhere herein. The mesh spacer can act as a structure that facilitates fluid inflow or outflow, providing the advantage of relatively rapid fluid movement, mixing, and / or diffusion within the mesh to reduce local interferent concentrations and increase glucose and oxygen concentrations. The increased fluid volume within the mesh can also facilitate increased fluid movement in and out of the area that provides glucose and oxygen while removing or diluting interferents.

[0237] In one exemplary embodiment, the sensor is encased in a single layer of open polypropylene (PP) biocompatible mesh. When the sensor is inserted, the mesh holds the surrounding tissue away from the sensor surface and allows the influx of extracellular fluid to enter the space within the mesh, thereby creating a fluid pocket around the sensor. Within the fluid pocket, the fluid can mix substantially rapidly as the extracellular fluid enters and leaves the fluid pocket, or due to host movement. Interferants can be carried by the fluid, thereby mixing and / or diluting. Because the host can wear the sensor for multiple days, settlement periods will inevitably occur. During these periods, interferents may accumulate. However, the increased fluid volume provided by the mesh can substantially buffer the accumulated interferents until the settlement period ends. Once the settlement period ends, the accumulated interferents can be diluted or carried away by the influx or outflow of fluid.

[0238] In some embodiments, the mesh may be applied to the sensor either symmetrically or asymmetrically. For example, the mesh may be wrapped tightly around the sensor. In another example, a strip of mesh may be applied to only one side of the sensor. In yet another example, the mesh may form a flat envelope about a few millimeters to about a centimeter wide, with the sensor sandwiched within the envelope. In some embodiments, the mesh may cover only a portion of the sensor, such as the portion that includes the electrochemically reactive surface. In other embodiments, the mesh may cover the entire sensor.

[0239] In another alternative embodiment, noise can be reduced by including a hydrogel on the surface of at least a portion of the sensor (e.g., the sensing region). A hydrogel is a network of superabsorbent (can contain about 20% to about 99% water by weight, preferably more than 80% to 99% water by weight) natural or synthetic polymer chains. A hydrogel may also be viewed as a colloidal gel in which water is the dispersion medium. Because hydrogels are non-porous, fluids and interfering substances within the hydrogel move by diffusion. Thus, the movement of molecules within a hydrogel is relatively slower than that possible within a mesh-based fluid pocket as described above. Optionally, the hydrogel may be biodegradable. A biodegradable hydrogel can provide a fluid pocket that gradually decreases and is eventually eliminated by the surrounding tissue.

[0240] In further embodiments, the hydrogel comprises a flexible water-swellable film (as disclosed elsewhere herein) having a "dry film" thickness of about 0.05 microns or less to about 20 microns or more, more preferably about 0.05, 0.1, 0.15, 0.2, 0.25, 0.3, 0.35, 0.4, 0.45, 0.5, 1, 1.5, 2, 2.5, 3, or 3.5 to about 4, 5, 6, 7, 8, 9, 10, 11, 12, 13, 14, 15, 16, 17, 18, 19, or 19.5 microns, more preferably about 2, 2.5 or 3 microns to about 3.5, 4, 4.5, or 5 microns. "Dry film" thickness refers to the thickness of a cured film cast from a coating formulation by standard coating techniques. The hydrogel material may be applied to the entire sensor or a portion thereof using any method known in the art, such as, but not limited to, dipping, painting, spraying, wrapping, etc.

[0241] In some embodiments, a scavenging agent (e.g., a bioactive agent capable of removing, binding, or substantially inactivating an interferent) can be incorporated into the hydrogel or other aspects of the device (e.g., the membrane system). The scavenging agent can inhibit long-term wound and inflammation by removing signals associated with irritants from the sensor and / or the site of internally generated hydrogen peroxide. One exemplary scavenging agent embodiment incorporates H.sub.2O.sub.2 degrading enzymes (e.g., including but not limited to glutathione peroxidase (GSH peroxidase), catalase, heme-containing peroxidase, eosinophil peroxidase, thyroid peroxidase, or horseradish peroxidase (HRP)) into the hydrogel to degrade available H.sub.2O.sub.2 and generate oxygen. The scavenging agent can act within the hydrogel or can be released into the local environment to act outside of the hydrogel.

[0242] In some embodiments, mesh and hydrogel can be used in combination to provide better mechanical support (hold surrounding tissue away from the sensor) while slowing down the diffusion rate in the mesh-hydrogel layer. For example, a PP mesh can be applied to the sensor, followed by spraying dry hydrogel material onto the PP-wrapped sensor. Alternatively, the hydrogel can be dried within the mesh before application to the sensor. Once the sensor is embedded, the hydrogel can absorb fluid from the surrounding tissue, expand, and fill the mesh pores. In a further example, the hydrogel can be biodegradable. In this example, the hydrogel can initially slow down fluid movement. However, as the hydrogel biodegrades, the mesh pores can open up and fluid movement can accelerate or increase.

[0243] A variety of alternative materials can be used to form the architecture that forms the fluid pocket. For example, a shape memory material can be used as an alternative to mesh to form a fluid pocket around the sensor. Shape memory materials are metals or polymers that "remember" their geometric shape. Shape memory metals (e.g., memory metals or smart wires) include copper-zinc-aluminum, copper-aluminum-nickel, and nickel-titanium (NiTi) alloys. Shape memory polymers include materials such as polynorbornene, segmented poly(epsilon-caprolactone) polyurethane, poly(ethylene glycol)-poly(epsilon-caprolactone) diblock copolymers. Shape memory materials can deform from their "original" form in response to forces such as temperature or pressure, and regain their original geometric shape by themselves.

[0244] In some embodiments, the porous memory material, which has been crushed into a flat, non-porous sheet, can be applied to the exterior of the sensor as a flat film. After insertion into the body, increased temperature or exposure to moisture can stimulate the memory material to transform into a three-dimensional porous architecture that, for example, promotes the formation of fluid pockets.

[0245] In some embodiments, nanoporous materials acting as molecular sieves can be used to eliminate interfering substances surrounding the sensor. In another alternative embodiment, a swellable material (e.g., a material having an initial volume that, upon contact with a fluid, such as water, absorbs the fluid to a second volume greater than the initial volume) or a collapsible material (e.g., a material having an initial volume that collapses to a second volume less than the initial volume) can create or maintain the fluid pockets.

[0246] In some embodiments, materials with different properties can be applied in combination, such as alternating bands or layers, to inhibit uniform capsule formation. For example, alternating bands of crushable and non-crushable swellable materials can be applied around a portion of the sensor. Upon implantation, both materials swell with fluid from the surrounding tissue. However, only segments of the crushable material can deform. The material surrounding the sensor is irregular, which can disrupt the formation of a continuous cell layer, thereby reducing noise and increasing the lifespan of the sensor.

[0247] In addition to providing physical spacers, meshes, porous materials, etc., stimulative sensor configurations can reduce noise by promoting fluid pocket formation and / or increased bulk fluid flow. Thus, one embodiment of an stimulative biointerface includes a structure having a rough surface that can scrape or poke adjacent cells in vivo. The sensor surface can be roughened, for example, by coating the sensor with a machinable material that is or can be etched to form ridges, bristles, spikes, grids, grooves, circles, spirals, dots, bumps, pits, etc. The material can be any convenient biocompatible material such as a machined porous structure overlaid on the sensor, including, but not limited to, machinable metal matrix composites, bone matrices such as hydroxyapatite, coral hydroxyapatite, and beta-tricalcium phosphate (TCP), porous titanium (Ti) mixtures made by sintering elemental powders, bioglass (calcium and silicon based porous glass), ceramics, etc. The material may be "machined" by any convenient means, such as, but not limited to, scraping, etching, lathing, or laser machining.

[0248] The micro-motion of the sensor can increase the stimulating effect of the rough surface. Micro-motion is an inherent characteristic of any implanted device, such as an implanted glucose sensor. The micro-motion of the device (e.g., the fine movement of the device within the host) is caused by the movement of the host, ranging from breathing and small local muscle movements to gross motor actions such as walking, running, or even rising and sitting down. External forces, such as the application of external pressure, can also cause micro-motion. Micro-motion includes back and forth movement, rotation, twisting, and / or rotation of the sensor. Thus, when the sensor moves by micro-motion, the rough surface of the sensor can rub the surrounding tissue violently, causing an increase or enlargement of the wound, resulting in additional stimulation of the wound healing process, increasing the fluid bulk, bulk fluid flow, and / or fluid pocket formation, and concomitantly reducing noise.

[0249] In some embodiments, the stimulable architecture is formed from self-assembling materials. Self-assembling biomaterials contain specific polypeptides that are pre-designed to self-assemble into targeted nano- and microscopic structures. Intramolecular self-assembling molecules are often complex polymers that have the ability to assemble from a random coil conformation into distinct stable structures (secondary and tertiary structures). Various self-assembling materials known in the art can be used in this embodiment. For example, PuraMatrix® (3DM Inc., Cambridge, Massachusetts, USA) can be used to create synthetic self-assembling peptide nanofiber scaffolds and defined 3-D microenvironments.

[0250] In an exemplary embodiment of the stimulating biointerface, a stimulating superstructure is applied to the working electrode or the completed sensor. "Superstructure" as used herein is a broad term and is used in its ordinary sense, including but not limited to referring to any structure built on top of another, such as overlapping portions of structures. The stimulating superstructure may include any substantial structure that prevents cell attachment and stimulates surrounding tissue in vivo. In one example, the stimulating superstructure may include large spaces, such as at least about 50 μm wide and at least about 50 μm deep. Cells surrounding the sensor may be prevented from attaching to the spaces within the superstructure, allowing fluid to fill these spaces. In some exemplary embodiments, the stimulating superstructure utilizes micromotion of the sensor to prevent cell attachment and stimulate fluid pocket formation.

[0251] In one exemplary embodiment, the irritating superstructure is comprised of ridges at least about 0.25 μm to 0.50 μm in diameter and about 50 μm in height, separated by at least about 0.25 μm to 0.50 μm. In another exemplary embodiment, exposed silver wires at least about 0.25 μm to 0.50 μm in diameter are applied to the exterior of the sensor to form grooves about 50 μm wide and about 50 μm deep. Because silver is pro-inflammatory and stimulates fluid influx from surrounding tissues, the combination of the irritating superstructure with a chemical irritant may promote an increased rate of fluid influx or prolong the stimulation and fluid influx.

[0252] In yet another exemplary embodiment, with reference to the embodiment shown in FIG. 4A, the configuration (e.g., diameter) of the reference electrode 30 can be altered (e.g., increased size and / or coil spacing) such that the reference electrode itself becomes the stimulating superstructure.

[0253] l. Porous membrane In addition to the above-mentioned devices, bulk fluid flow and / or bulk fluid flow at and / or adjacent to the sensor can be increased by incorporating a porous membrane into the sensor system such that noise is substantially reduced and sensor accuracy and / or sensitivity is improved. The porous membrane can be referred to as a "bioprotective domain" or "cytolytic domain." In some embodiments, the sensor includes a porous material disposed over a portion thereof that modifies the host tissue response to the sensor, thereby reducing noise (e.g., due to local accumulation of electroactive species that may interfere with the analyte signal). For example, in some embodiments, the porous material surrounding the sensor advantageously enhances and extends the performance and lifespan of the sensor in the short term by slowing or reducing cell migration to the sensor and associated degradation that would otherwise be caused by cell infiltration if the sensor were directly exposed to an in vivo environment. Alternatively, the porous material can provide stabilization of the sensor via tissue ingrowth into the porous material in the long term.Suitable porous materials include silicone, polytetrafluoroethylene, expanded polytetrafluoroethylene, ethylene tetrafluoroethylene copolymers, polyolefins, polyesters, polycarbonates, biostable polytetrafluoroethylene, homopolymers, copolymers, and terpolymers of polyurethane, polypropylene (PP), polyvinyl chloride (PVC), polyvinylidene fluoride (PVDF), polyvinyl alcohol (PVA), polybutylene terephthalate (PBT), polymethyl methacrylate (PMMA), polyether ether ketone (PEEK), polyamides, polyurethanes, cellulosic polymers, poly(ethylene oxide), poly(propylene oxide) and copolymers and blends thereof, polysulfone and block copolymers thereof, such as diblock, triblock, alternating, random, and graft copolymers, as well as metals, ceramics, cellulose, hydrogel polymers, poly(2-hydroxyethyl methacrylate, pHEMA), hydroxyethyl methacrylate, (hydroxyethyl methacrylate, HEMA), polyacrylonitrile-polyvinyl chloride (PAFC), polyvinyl chloride-polyvinyl chloride (PCB ... chloride, PAN-PVC), high density polyethylene, acrylic copolymers, nylon, polyvinyl fluoride, polyanhydrides, poly(L-lysine), poly(L-lactic acid), hydroxyethyl methacrylate, hydroxyapeptite, alumina, zirconia, carbon fiber, aluminum, calcium phosphate, titanium, titanium alloys, nitinol, stainless steel, and CoCr alloys, and the like, as described, for example, in U.S. Patent Application Publication No. 2005-0031689(A1) and U.S. Patent Application Publication No. 2005-0112169(A1).

[0254] In some embodiments, the porous material surrounding the sensor provides unique short-term (e.g., 1-14 days) advantages that can be used to improve and extend the performance and lifespan of the sensor. However, such materials can also provide advantages in the long term (e.g., more than 14 days). In particular, the in vivo portion of the sensor (the portion of the sensor that is implanted in the host's tissue) is encased (partially or completely) in the porous material. The porous material can be wrapped around the sensor (e.g., by wrapping the porous material around the sensor or by inserting the sensor into a section of the porous material sized to receive the sensor). Alternatively, the porous material can be deposited on the sensor (e.g., by electrospinning a polymer directly onto it). In yet other alternative embodiments, the sensor is inserted into a selected portion of the porous biomaterial. Other methods of encasing the in vivo portion of the sensor with a porous material can also be used, as will be understood by those skilled in the art.

[0255] The porous material surrounding the sensor advantageously slows or reduces cell migration to the sensor and associated degradation that would be caused by cell invasion if the sensor were directly exposed to the in vivo environment. That is, the porous material provides a barrier that makes cell migration toward the sensor more tortuous and therefore slower (providing a short-term benefit). This is believed to reduce or delay the loss of sensitivity typically observed in short-term sensors over time.

[0256] In one embodiment, where the porous material is a high oxygen solubility material, such as porous silicone, the high oxygen solubility porous material surrounds some or all of the in vivo portion of the sensor. In some embodiments, a lower oxygen to glucose ratio may be sufficient to provide excess oxygen by using a high oxygen solubility domain (e.g., silicone-based or fluorocarbon-based materials) to enhance oxygen supply / transport to the enzyme membrane and / or electroactive surface. It is believed that some signal noise typically seen with conventional sensors may be due to oxygen deficiency. Silicone has high oxygen permeability, thus facilitating oxygen transport to the enzyme layer. By enhancing oxygen supply through the use of silicone compositions, for example, glucose concentration may not be a limiting factor. In other words, if more oxygen is supplied to the enzyme and / or electroactive surface, more glucose may also be supplied to the enzyme without creating an oxygen-limiting excess. Without being bound to any particular theory, it is believed that silicone materials provide enhanced biostability compared to other polymeric materials, such as polyurethane.

[0257] In certain aspects, sensor performance can be improved by including a biointerface structure, material, matrix, and / or membrane that forms a suitable space on the sensor for fluid to fill in vivo. In some embodiments, the sensor includes a porous biointerface material that allows fluid from the surrounding tissue to form a fluid-filled pocket around at least a portion of the sensor. It is believed that the fluid-filled pocket provides a sufficient source of analyte-containing fluid for accurate sensor measurements in a short period of time. Additionally or alternatively, the inclusion of a bioactive agent can modify the host tissue response, for example, reducing or eliminating tissue ingrowth or other cellular responses into the biointerface.

[0258] In some aspects, sensor performance can be improved by modifying the sensor with structures, materials, and / or membranes / matrices that allow tissue ingrowth without forming barrier cells. For example, a vascularized bed of tissue for long-term analyte sensor measurements. In some embodiments, a porous biointerface membrane that includes multiple interconnected cavities and solid portions and covers at least the sensing portion of the sensor allows vascularized tissue ingrowth therein. The vascularized tissue ingrowth provides a sufficient source of analyte-containing tissue for extended periods of time. Additionally or alternatively, the inclusion of a bioactive agent can modify the host tissue response, for example, reducing or eliminating barrier cell layer formation within the membrane.

[0259] As used herein, the terms "membrane" and "matrix" are meant to be interchangeable. In these embodiments, a first domain is provided that includes an architecture, including cavity size, configuration, and / or overall thickness, that modifies the host tissue response, for example, by forming fluid pockets, promoting ingrowth of vascularized tissue, breaking down underlying tissue contracture, resisting fibrotic tissue growth adjacent to the device, and / or preventing the formation of barrier cells. The biological interface preferably covers at least the sensing mechanism of the sensor and can be of any shape or size, including uniformly, asymmetrically, or axially symmetrically covering or surrounding the sensing mechanism or sensor.

[0260] In some embodiments, a second domain that is impermeable to cells and / or cellular processes is optionally provided. A bioactive agent is optionally provided that is incorporated into at least one of the first domain, the second domain, the sensing membrane, or other portions of the implantable device, where the bioactive agent is configured to modify a host tissue response.

[0261] In one embodiment, the porous material that provides for increased fluid bulk, bulk fluid flow and / or diffusion rates, and the formation of tightly packed vascular structures, is a porous polymer membrane, such as, but not limited to, polytetrafluoroethylene (PTFE), polysulfone, polyvinylidene difluoride, polyacrylonitrile, silicone, polytetrafluoroethylene, expanded polytetrafluoroethylene, polyethylene-co-tetrafluoroethylene, polyolefin, polyester, polycarbonate, biostable polytetrafluoroethylene, polyurethane homopolymers, copolymers, and terpolymers, polypropylene (PP), polyvinyl chloride (PVC), polyvinylidene fluoride (PVDF), polyvinyl alcohol (PVA), polybutylene terephthalate (PBT), polymethyl methacrylate (PMMA), polyether. etherketone (PEEK), polyamide, polyurethane, cellulosic polymer, poly(ethylene oxide), poly(propylene oxide) and copolymers and blends thereof, polysulfone and block copolymers thereof (including, for example, diblock, triblock, alternating, random and graft copolymers), as well as metal, ceramic, cellulose, hydrogel polymer, poly(2-hydroxyethyl methacrylate, pHEMA), hydroxyethyl methacrylate (HEMA), polyacrylonitrile-polyvinyl chloride (PAN-PVC), high density polyethylene, acrylic copolymers, nylon, polyvinyl difluoride, polyanhydrides, poly(l-lysine), poly(L-lactic acid), and hydroxyethyl methacrylate, having an average nominal pore size of at least about 0.6 μm to 20 μm using conventional methods for determining pore size in the art. In one embodiment, at least about 50% of the pores in the membrane have an average size of about 0.6 μm to about 20 μm, as described in U.S. Pat. No. 5,882,354. In this exemplary embodiment, the structural elements that provide the three-dimensional structure may include fibers, strands, globules, cones, or rods of smooth or rough, irregular or uniform geometric shape. These elements (hereinafter referred to as "strands") generally have one dimension larger than the other two, with the smaller dimension not exceeding 5 microns.

[0262] In another further embodiment, the porous polymeric membrane material is comprised of strands that define "apertures" formed by a frame of interconnected strands as described above. The apertures have an average size of about 20 μm or less in any dimension other than the longest dimension. The apertures of the material form a framework of interconnected apertures and define "cavities" that average about 20 μm or less in any dimension other than the longest dimension. In another embodiment, the porous polymeric membrane material has at least some apertures that are of a size sufficient to allow at least some vascular structures to form within the cavity. At least some of these apertures allow vascular structures to form within the cavity, but prevent connective tissue from forming therein due to size restrictions.

[0263] In further embodiments, the porous membrane has a frame of elongated strands of material less than 5 microns in all but the longest dimension, the frame defining openings that interconnect to form a three-dimensional cavity that allows substantially all inflammatory cells that migrate into the cavity to maintain a rounded morphology. Furthermore, the porous material promotes angiogenesis adjacent to, but not substantially within, the porous material upon implantation into a host. Exemplary materials include, but are not limited to, polyethylene, polypropylene, polytetrafluoroethylene (PTFE), cellulose acetate, cellulose nitrate, polycarbonate, polyester, nylon, polysulfone, mixed esters of cellulose, polyvinylidene difluoride, silicone, polyacrylonitrile, and the like.

[0264] In some embodiments, the short-term sensor comprises a spacer adapted to provide a fluid pocket between the sensor and the host tissue, the structure of the spacer, e.g., a biointerface material, matrix, mesh, hydrogel, etc., and the resulting fluid pocket are believed to provide for the transport of oxygen and / or glucose to the sensor.

[0265] In one exemplary embodiment, the sensor includes a biointerface membrane configured to prevent fat cells from contacting the inserted transcutaneous or implanted sensor. Preferably, the porous biointerface membrane surrounds the sensor, covers the sensing mechanism (e.g., at least the working electrode), and is configured to be filled with fluid in vivo, thereby forming a fluid pocket surrounding the sensor. Thus, fat cells surrounding the sensor are held at a distance (e.g., the thickness of the porous biointerface membrane) from the sensor surface. Thus, when the porous biointerface membrane is filled with fluid (e.g., forming a fluid pocket), sufficient amounts of oxygen and glucose are transported to the sensing mechanism to maintain accurate sensor function. In addition, as described elsewhere herein, interferents are diluted, suppressing or reducing interference with sensor function.

[0266] In another exemplary embodiment, a short-term sensor (or short-term function of a long-term sensor) including a biointerface, including but not limited to, for example, a porous biointerface material, a mesh cage, etc., can be used to improve short-term (e.g., first hours to days) sensor function, such as by reducing noise on the sensor signal. A porous biointerface membrane does not necessarily include interconnected cavities to form fluid pockets in the short term.

[0267] m. Bioactive agents Various bioactive agents are known to promote fluid influx or outflux. Thus, incorporation of a bioactive agent into the membrane can increase fluid bulk, bulk fluid flow, and / or diffusion rate (and promote glucose and oxygen influx), thereby reducing non-stationary noise. In some embodiments, fluid bulk and / or bulk fluid flow is increased at the sensor (e.g., adjacent to the sensor exterior surface) by incorporation of one or more bioactive agents. In some embodiments, the sensor is configured to include a bioactive agent that stimulates the release of soluble mediators known to stimulate wounds and cause local fluid influx at the wound site. In some embodiments, the sensor is configured to include a vasodilatory bioactive agent that can cause local influx of fluid from the vasculature.

[0268] A variety of bioactive agents may be found to be useful in preferred embodiments. Exemplary bioactive agents include, but are not limited to, blood-brain barrier disrupting agents and vasodilators, vasodilators, angiogenic factors, and the like.Useful bioactive agents include mannitol, sodium thiosulfate, VEGF / VPF, NO, NO donors, leptin, bradykinin, histamine, blood components, platelet rich plasma (PRP), matrix metalloproteinases (MMPs), Basic Fibroblast Growth Factor (bFGF), (also known as heparin-binding growth factor II and fibroblast growth factor II), Acidic Fibroblast Growth Factor (aFGF), (also known as heparin-binding growth factor-I and fibroblast growth factor-I), Vascular Endothelial Growth Factor (VEGF), Platelet Derived Endothelial Cell Growth Factor BB (PDEGF-BB), Angiopoietin-1, Transforming Growth Factor beta (TGF-1), and the like. Beta, TGF-Beta, Transforming Growth Factor Alpha (TGF-Alpha), Hepatocyte Growth Factor, Tumor Necrosis Factor-Alpha (TNF-Alpha), Placental Growth Factor (PLGF), Angiogenin, Interleukin-8 (IL-8), Hypoxia Inducible Factor-I (HIF-1), Angiotensin-Converting Enzyme (ACE) inhibitors Quinaprilat, Angiotropin, Thrombospondin, KGHK peptide, Hypoxic Tone, Lactate, Insulin, Leptin, Copper Sulfate, Estradiol, Prostaglandins, Cox Inhibitors, Endothelial Cell Binding Agents (e.g., Decorin or Vimentin), Glenipin, Hydrogen Peroxide, Nicotine, and Growth Hormone.Still other useful bioactive agents include enzymes, cytotoxic or necrotic agents (e.g., pactataxil, actinomycin, doxorubicin, daunorubicin, epirubicin, bleomycin, plicamycin, mitomycin), cyclophosphamide, chlorambucil, uramustine, melphalan, bryostatin, inflammatory bacterial cell wall components, histamine, pro-inflammatory factors, and the like.

[0269] Bioactive agents may be added during the manufacture of the sensor by incorporating the desired bioactive agent in the manufacturing materials for one or more sensor layers or in an external biomaterial such as a porous silicone membrane. For example, the bioactive agent may be mixed with a solution during membrane formation and then applied onto the sensor during manufacture. Alternatively, the completed sensor may be immersed in a solution of the bioactive agent or sprayed with a solution of the bioactive agent, for example. The amount of bioactive agent may be controlled by varying its concentration, varying the dwell time during immersion, applying multiple layers until a desired thickness is reached, etc., as disclosed elsewhere herein. In an alternative embodiment, the bioactive agent is microencapsulated prior to application to the sensor. For example, the microencapsulated bioactive agent may be sprayed onto the completed sensor or incorporated within a structure such as an outer mesh layer or a shedding layer. Microencapsulation may provide increased flexibility in controlling the bioactive agent release rate, the time of release occurrence, and / or the duration of release.

[0270] Stimulation chemistry / methods can be incorporated into external sensor structures such as a biointerface membrane (described elsewhere herein) or a decidual layer that releases stimulants into the local environment. For example, in some embodiments, a "decidual layer" can release (e.g., shed or leach) molecules into the local vicinity of the sensor to accelerate the osmotic fluid shift. In some embodiments, the decidual layer can provide a mild stimulus and promote a mild inflammatory / foreign body response, thereby preventing cells from stabilizing and building an ordered fibrous cap and promoting fluid pocket formation.

[0271] The decidual layer can be constructed from any convenient biocompatible material, including, but not limited to, hydrophilic degradable materials such as polyvinyl alcohol (PVA), PGC, polyethylene oxide (PEO), polyethylene glycol-polyvinylpyrrolidone (PEG-PVP) blends, PEG-sucrose blends, hydrogels such as polyhydroxyethyl methacrylate (pHEMA), polymethyl methacrylate (PMMA), or other polymers with rapidly degrading ester linkages. In certain embodiments, absorbable suture materials can be used that degrade into compounds with acid residues. Acid residues are chemical irritants that stimulate inflammation and wound healing. In certain embodiments, these compounds include glycolic and butyric acid based polymers, polyglactin, polydioxanone, polyziconate, poly(dioxanone), poly(trimethylene carbonate) copolymers, and poly(caprolactone) homopolymers and copolymers.

[0272] In other exemplary embodiments, the shedding layer may be a layer of materials listed elsewhere herein for the first domain, including polyethers such as polyvinylpyrrolidone (PVP), polyhydroxyethylmethacrylate, polyvinylalcohol, polyacrylic acid, polyethylene glycol, and copolymers or blends thereof with hydrophilic polymers, such as, for example, diblock, triblock, alternating, random, and graft copolymers (block copolymers are described in U.S. Pat. No. 4,803,243 and U.S. patents). In one preferred embodiment, the shedding layer is composed of polyurethane and a hydrophilic polymer. For example, the hydrophilic polymer may be polyvinylpyrrolidone. In one preferred embodiment, the shedding layer is a polyurethane comprising 5 weight percent or more polyvinylpyrrolidone and 45 weight percent or less polyvinylpyrrolidone. Preferably, the shedding layer comprises a polyurethane comprising 20 weight percent or more polyvinylpyrrolidone and 35 weight percent or less polyvinylpyrrolidone, and most preferably, about 27 weight percent polyvinylpyrrolidone.

[0273] In other exemplary embodiments, the shedding layer may comprise a silicone elastomer, such as a silicone elastomer and a poly(ethylene oxide) and poly(propylene oxide) copolymer blend, as disclosed in co-pending U.S. Patent Application Serial No. 11 / 404,417, filed April 14, 2006. In one embodiment, the silicone elastomer is a dimethyl- and methylhydrogen-siloxane copolymer. In one embodiment, the silicone elastomer comprises vinyl substituents. In one embodiment, the silicone elastomer is an elastomer produced by curing a MED-4840 mixture. In one embodiment, the copolymer comprises hydroxyl substituents. In one embodiment, the copolymer is a triblock poly(ethylene oxide)-poly(propylene oxide)-poly(ethylene oxide) polymer. In one embodiment, the copolymer is a triblock poly(propylene oxide)-poly(ethylene oxide)-poly(propylene oxide) polymer. In one embodiment, the copolymer is a PLURONIC® polymer. In one embodiment, the copolymer is PLURONIC® F-127. In one embodiment, at least a portion of the copolymer is crosslinked. In one embodiment, about 5% w / w to about 30% w / w of the membrane is the copolymer.

[0274] The shedding layer can be of any shape or geometry, symmetric or asymmetric, to promote fluid inflow at a desired location of the sensor, such as, for example, the sensor head or the electrochemically reactive surface. The shedding layer can be disposed on one or both sides of the sensor. In another example, the shedding layer can be applied to only a small portion of the sensor or to the entire sensor.

[0275] In one exemplary embodiment, a decidual layer comprising polyethylene oxide (PEO) is applied to the exterior of the sensor, allowing the tissue surrounding the sensor direct access to the decidual layer. The PEO leaches out of the decidual layer and is taken up by local cells which release pro-inflammatory factors. The pro-inflammatory factors diffuse through the surrounding tissue and stimulate an inflammatory response including fluid influx. Thus, premature noise can be reduced or eliminated and sensor function can be improved.

[0276] In another exemplary embodiment, the decidual layer is applied to the sensor in combination with an outer porous layer, such as a mesh or porous biointerface as disclosed elsewhere herein. In one embodiment, the local cells access the decidual layer through the through-holes of the porous silicone biointerface. In one example, the decidual layer material is applied to the sensor prior to application of the porous silicone. In another example, the decidual layer material can be absorbed into the bottom of the porous silicone (e.g., the portion of the porous silicone that will be proximal to the sensor after the porous silicone is applied to the sensor) prior to application of the porous silicone to the sensor.

[0277] n. Wound control Non-stationary noise can be reduced in some embodiments by wound suppression (e.g., during sensor insertion). Wound suppression includes any system or method that reduces and / or eliminates the amount of wound that occurs during sensor insertion. Without being bound by theory, it is believed that when wounding is suppressed or at least significantly reduced, the sensor is surrounded by substantially normal tissue (e.g., tissue substantially similar to the tissue before the sensor is inserted). It is believed that substantially normal tissue has a lower metabolism than wounded tissue and produces fewer interferents, reducing early noise.

[0278] Wounding can be inhibited or minimized by adapting the sensor architecture to an architecture that inhibits wounding or promotes rapid healing, such as an architecture that does not cause substantial wounding (e.g., an architecture configured to prevent wounding), an architecture that promotes wound healing, an anti-inflammatory architecture, etc. In one exemplary embodiment, the sensor is configured to have a low profile, zero footprint, or smooth surface. For example, the sensor can be formed from a substantially thin wire, such as a wire with a diameter of about 50 μm to about 150 μm. Preferably, the sensor is small enough to fit within a very small gauge needle, such as, for example, a 30, 31, 32, 33, 34, or 35 gauge needle (or smaller) on the stub scale. Generally, the smaller the needle, the less the amount of wounding during insertion. For example, a very small needle can reduce the amount of tissue disruption, thereby reducing the subsequent wound healing response. In an alternative embodiment, the surface of the sensor is smoothed with a lubricious coating to reduce wounding upon sensor insertion.

[0279] Wounds can also be reduced by including a wound inhibitor (bioactive agent), which reduces the amount of initial wound or inhibits the wound healing process. Without being bound by theory, it is believed that applying a wound inhibitor, such as an anti-inflammatory, immunosuppressant, anti-infective, or removal agent, to the sensor can create a local quiescent environment and inhibit wound healing. In a quiescent environment, bodily processes, such as increased cellular metabolism associated with wound healing, have minimal impact on the sensor. If the tissue surrounding the sensor is undisturbed, it can continue its normal metabolism and facilitate sensor function.

[0280] In some embodiments, compounds and / or agents useful for inhibiting wound healing include first generation H.sub.1 receptor antagonists, i.e., ethylenediamines (e.g., mepyramine (pyrilamine), antazoline), ethanolamines (e.g., diphenhydramine, carbinoxamine, doxylamine, clemastine, and dimenhydrinate), alkylamines (pheniramine, chlorpheniramine (chlorpheniramine), dexchlorpheniramine, brompheniramine, and triprolidine), piperazines (cyclizine, hydroxyzine, and meclizine), and tricyclics (promethazine, alimemazine (trimeprazine), cyproheptadine, , and azatadine); second generation H.sub.1 receptor antagonists (e.g., acrivastine, astemizole, cetirizine, loratadine, mizolastine, azelastine, levocabastine, and olopatadine); mast cell stabilizers (e.g., cromoglycate (cromolyn) and nedocromil); anti-inflammatory agents (e.g., acetophenone, aminosalicylic acid, aspirin, celecoxib, choline magnesium trisalicylate, diclofenac potassium, diclofenac sodium, diflunisal, etodolac, fenoprofen, flurbiprofen, ibuprofen, indomethacin, interleukin (IL)-10, IL-6 muteins, anti-IL-6 iNOS inhibitors (e.g., L-NMDA), interferons, ketoprofen, ketorolac, leflunomide, melenamic acid, mycophenolic acid, mizoribine, nabumetone, naproxen, naproxen sodium, oxaprozin, piroxicam, rofecoxib, salsalate, sulindac, and tolmetin);Corticosteroids (e.g., cortisone, hydrocortisone, methylprednisolone, prednisone, prednisolone, betamethasone, beclomethasone dipropionate, budesonide, dexamethasone sodium phosphate, flunisolide, fluticasone propionate, paclitaxel, tacrolimus, tranilast, triamcinolone acetonide, betamethasone, fluocinolone, fluocinonide, betamethasone dipropionate, betamethasone valerate, desonide, desoximetasone, fluocinolone, triamcinolone acetonides, clobetasol propionate, and dexamethasone; immunosuppressants and / or immunomodulators (e.g., antiproliferative agents, cell cycle inhibitors (e.g., paclitaxel, cytochalasin D, infiximab), taxol, actinomycin, mitomycin, thospromote VEGF, estradiol, NO donors, QP-2, tacrolimus, tranilast, actinomycin, everolimus, methotrexate, mycophenolic acid, angiopeptin, vincristine, mitomycin, statins, C MYC antisense, sirolimus (and analogs), restenase, 2-chlorodeoxyadenosine, PCNA ribozyme, batimustat, prolyl hydroxylase inhibitors, PPAR gamma ligands (e.g., troglitazone, rosiglitazone, pioglitazone), halofuginone, C-proteinase inhibitors, probucol, BCP671, EPC antibodies, catchin, glycation agents, endothelin inhibitors (e.g., ambrisentan, tesosentan, bosentan), statins (e.g., cerivastatin), E. coli burn enterotoxin, and advanced coatings; anti-infectives (e.g., anthelmintics (mebendazole));Antibiotics, such as aminoglycosides (gentamicin, neomycin, tobramycin), antifungal antibiotics (amphotericin b, fluconazole, griseofulvin, itraconazole, ketoconazole, nystatin, micatin, tolnaftate), cephalosporins (cefachlor, cefazolin, cefotaxime, ceftazidime, ceftriaxone, cefuroxime, cephalexin), beta-lactam antibiotics (cefotetan, meropenem), chloramphenicol, macrolide antibiotics (azithromycin, clarithromycin, erythromycin), penicillin antibiotics (penicillin G sodium salt, amoxicillin, ampicillin, dicloxacillin, nafcillin, piperacillin, thiamine ... calcillin), tetracyclines (doxycycline, minocycline, tetracycline), bacitracin);clindamycin;colistin metasodium;polymyxin b sulfate;vancomycin;antivirals including acyclovir, amantadine, didanosine, efavirenz, foscarnet, ganciclovir, indinavir, lamivudine, nelfinavir, ritonavir, saquinavir, silver, stavudine, valacyclovir, valganciclovir, zidovudine;quinolone antibiotics (ciprofloxacin, levofloxacin);sulfonamides (sulfadiazine, sulfisoxazole);sulfones (dapsone);furazolidone;metronidazole;pentamidine;sulfanilamidum crystallinum crystallinum); gatifloxacin; and sulfamethoxazole / trimethoprim; interference scavengers (e.g., superoxide dismutase (SOD), thioredoxin, glutathione peroxidase and catalase, antioxidants such as uric acid and vitamin C, iron compounds, heme compounds, and some heavy metals); artificial protective coating components (e.g., albumin, fibrin, collagen, endothelial cells, wound closure chemicals, blood products, platelet-rich plasma, growth factors);

[0281] Without being bound by theory, it is believed that the application of a lubricious coating to the sensor, in addition to the analyte sensor configurations described elsewhere herein, can substantially reduce and / or suppress noise generation by substantially preventing injury to the host. Thus, in some embodiments, a lubricious coating can be applied to the in vivo portion of the sensor to reduce foreign body reaction to the implanted sensor. As used herein, the term "lubricious coating" is a broad term and is used in its ordinary sense, including, but not limited to, surface treatments that reduce surface friction. A variety of polymers are suitable for use as lubricious sensor coatings, including, but not limited to, Teflon, polyethylene, polycarbonate, polyurethane, poly(ethylene oxide), poly(ethylene oxide)-poly(propylene oxide) copolymers. In one exemplary embodiment, one or more layers of HydroMed, a polyether-polyurethane manufactured by CardioTech International, Inc. (Wilmington, Massachusetts, USA), are applied to the sensor (e.g., over the resistive domain).

[0282] In some embodiments, wounding can be inhibited by including a silicone coating (e.g., a silicone-hydrophilic polymer blend) or a hydrophilic shedding layer can be applied to the sensor. Without being bound by theory, it is believed that the silicone bioprotective coating or decidual layer can promote the formation and maintenance of a fluid pocket around the sensor to improve glucose and fluid transport and clearance of interferents. The silicone bioprotective coating can create a local environment with enhanced vascular permeability and / or vascularization. Such a coating is believed to accelerate the inflammatory response to achieve a substantially consistent wound environment more quickly than in the absence of the coating. Furthermore, it is believed that the silicone bioprotective coating can reduce the inflammatory response and reduce the production of cellular by-products that are believed to be electrochemical interferents.

[0283] In one embodiment, the silicone bioprotective coating may be comprised of one or more layers formed from compositions that provide high oxygen solubility as well as allowing the transport of glucose or other such water-soluble molecules (e.g., drugs). In one embodiment, the layers comprise a blend of a silicone polymer and a hydrophilic polymer. For further description, see the section entitled "Silicon / Hydrophilic Polymer Blend Materials" herein, as well as U.S. Patent Application No. 11 / 404,417, filed April 14, 2006; U.S. Patent Application No. 11 / 675,063; U.S. Patent Application Publication No. 2005-0090607(A1); U.S. Patent Application Publication No. 2006-0270923(A1); and U.S. Patent Application Publication No. 2007-0027370(A1).

[0284] Many of the methods and structures disclosed above for forming fluid pockets, diluting interferents, reducing noise, etc., can be used in combination to promote a desired effect or outcome. For example, in one embodiment, a decidual layer composed of a hydrophilic silicone film and a necrotic agent can be applied in combination to at least a portion of the sensor. The silicone film can inhibit protein adhesion to the sensor surface, while the necrotic agent can devitalize a small portion of the tissue adjacent to the sensor and stimulate the formation of a fluid pocket around the hydrophilic silicone film. Preferably, the increased volume of fluid surrounding the sensor dilutes interferents, while the decidual layer provides a physical separation between the sensor and the surrounding tissue.

[0285] In another exemplary embodiment, a mesh sprayed with dexamethasone is wrapped around the exterior of the sensor. This mesh can provide a physical spacer for the fluid pocket while the dexamethasone inhibits inflammation. Preferably, the fluid can fill the mesh, and the dexamethasone can promote normal tissue metabolism around the sensor by inhibiting the influx of inflammatory cells. As a result, glucose and oxygen can freely travel between the tissue and the sensor through the fluid-filled mesh without accumulation of interferents, even during periods of tissue compression, thereby promoting sensor sensitivity and reducing noise.

[0286] Further description of increasing fluid bulk by adapting the sensor configuration can be found in U.S. Patent Application Publication No. 2006-0229512(A1) and U.S. Patent Application No. 11 / 654,140, ​​filed January 17, 2007.

[0287] o.Auxiliary electrode In some circumstances, non-stationary noise can be reduced by incorporating an auxiliary electrode into the sensor system that is configured to electrochemically modify (e.g., oxidize or reduce) the electrochemical interferents to overcome the effect of the interferents on the working electrode, rendering them substantially non-electroactively reactive at the electroactive sensing surface. Many electrochemical interferents can be reduced at potentials of about +0.1 V to about +1.2 V or higher, and it is known that, for example, acetaminophen is reduced at a potential of about +0.4 V. It should be noted that one challenge to electrochemically generating oxygen in this manner is that while the auxiliary electrode generates excess oxygen, placing the auxiliary electrode in close proximity to the analyte-measuring working electrode can cause oxidation of hydrogen peroxide at the auxiliary electrode, resulting in a reduced signal at the working electrode. Thus, the sensor of the preferred embodiment places the auxiliary electrode over the electrode system or other electroactive sensing surface, thereby reducing or eliminating the problem of inaccurate signals as described above.

[0288] Preferably, the auxiliary electrode is located within or adjacent to the membrane system, e.g., between the enzyme and other domains. However, the auxiliary electrode can be located anywhere between the electroactive sensing surface and the external fluid. The auxiliary electrode is formed from known working electrode materials (e.g., platinum, palladium, graphite, gold, carbon, conductive polymers, etc.) and has a voltage setting that generates oxygen (e.g., about +0.6V to about +1.2V or more) and / or a voltage setting that electrochemically modifies (e.g., reduces) electrochemical interferents to make them substantially non-reactive at the electroactive sensing surface (e.g., about +0.1V to about +1.2V or more). The auxiliary electrode can be a mesh, a grid, a plurality of spaced wires or a conductive polymer, or other configuration designed to allow the analyte to permeate therethrough.

[0289] In another aspect of the preferred embodiment, the auxiliary electrode is configured to electrochemically modify (e.g., oxidize or reduce) electrochemical interferents to render them substantially unreactive at the electroactive sensing surface. These embodiments can be in addition to or in place of the oxygen generating embodiments described above, where the polymer coating is selected to selectively allow interferents (e.g., urate, ascorbate, and / or acetaminophen as described in U.S. Pat. No. 6,579,690) to pass through the coating and react electrochemically with the auxiliary electrode, which effectively pre-oxidizes the interferent and renders it substantially unreactive at the working electrode. In one exemplary embodiment, the silicone material can be synthesized to allow transport of oxygen, acetaminophen, and other interferents, but not glucose. In some embodiments, the polymer coating material can be selected to have a molecular weight that blocks glucose and allows transport of oxygen, urate, ascorbate, and acetaminophen. In another exemplary embodiment, the silicone material can be synthesized to allow transport of oxygen, glucose, acetaminophen, and other interferents. In some embodiments, the polymer coating material is selected with a molecular weight that allows for the transport of oxygen, glucose, urate, ascorbate, and acetaminophen. The voltage setting required to react with the interfering species depends on the target electrochemical interferent, for example, from about +0.1 V to about +1.2 V. In some embodiments where the auxiliary electrode is set to a potential of about +0.6 to about +1.2 V, both oxygen generation and electrochemical interferent modification can be achieved. In some embodiments where the auxiliary electrode is set to a potential of less than about +0.6 V, the auxiliary electrode functions, for example, primarily to electrochemically modify the interferent. Further description can be found in U.S. Pat. No. 7,074,307.

[0290] p. Modeling factors affecting sensor signals to improve accuracy As mentioned above, CGM systems can suffer from calibration errors that cause them to provide inaccurate clinical blood glucose values. For example, some calibration codes may apply factory information to the sensor design, but this code is not tailored to specific patient, environment, or situation information. This is because factory calibration generally considers information derived from the manufacturing lot of the sensor during manufacturing, such as how a sensor with a particular lot characteristic (e.g., membrane thickness or electrode dimensions) interacts with a typical patient physiology under nominal conditions. However, the transformation of the sensor signal based on the factory-derived code is not personalized to the patient. With BG fingerstick calibration, for example, a broad correlation between a person's BG clinical value and the sensor signal can be established at one or more time points. However, the correlation is not static over time. Similarly, the application of BG fingerstick values ​​cannot distinguish the individualized effects associated with the sensor design and its interaction with the in vivo person's physiological environment throughout sensor wear.

[0291] To address this issue, a CGM system as described herein may include a differential modifier that increases or decreases the signal based on discrete factors that affect the signal over time, whether calibrated by factory information and / or real-time BG fingerstick meter values. This modifier may modify the sensor signal based on information associated with the sensor design and / or the sensor's interaction with the in vivo human physiological environment throughout sensor wear. In this manner, the sensor signal may be accurately transformed to provide an accurate clinical blood glucose value for a particular sensor and a particular patient in a particular physiological environment.

[0292] Modeling of factors affecting the sensor signal can be performed using information about the individualized sensor characteristics and behavior, as well as the individualized physiological characteristics of the patient. These factors can be modeled and used to process the sensor signal and remove their effects on the signal to provide accurate clinical data for further processing, output, and / or display. The systems and methods described herein that provide modeling of different factors affecting the signal can improve overall accuracy for all patients during the entire sensor session. For example, differentiated correction of differently affected aspects of the signal, such as calibration codes, can be applied to components of the signal affected by the sensor design, while other components of the signal affected by physiology can be left unaffected by the calibration codes and / or specifically modeled based on the specific characteristics of the patient. Independent modeling of different factors affecting the signal allows separate implementations of modeling (e.g., specific aspects of the sensor design and specific aspects of the physiological environment) to be isolated and applied only when appropriate, without affecting other components of the signal and their unique modeling.

[0293] In general, factors that affect the signal may include anything that affects the signal measured by the analyte sensor (e.g., causes an increase or decrease in the signal) other than the clinical blood analyte concentration itself. In other words, a complete analyte signal from a complete analyte sensor in a fully controlled environment provides a glucose-only signal that is not affected by any other factors. Such factors may result from a variety of sources, including factors resulting from individual sensor variability (possibly due to imperfect and variable manufacturing processes) and / or variability in the physiological characteristics of a particular patient using the analyte sensor. That is, these factors that affect the signal may be different factors from sensor to sensor and from patient to patient. These factors may also change over time.

[0294] Modeling the factors that affect the signal involves calculations resulting from a mathematical equation or mathematical terms that take the signal as an input and generate an output signal, and the algorithmic model provides a mathematical representation of the influence so that the factors can be represented independently, thereby allowing a final clinical value or representation of the analyte concentration to be reported without the influence of the factors.

[0295] Independent modeling of the factors allows for very specific representation of one or more of each factor, allowing for personalized fine tuning of the final analyte concentration based on the patient's unique physiological variations and the sensor's unique manufacturing variations. Because the effects on the signal are generally caused by a variety of different physiological and sensor variations, which generally do not increase or decrease in the same way, at the same rate, or may not even be correlated at all, personalized modeling has the ability to more accurately correct for each unique factor over time for each patient and each sensor.

[0296] Figure 6 is a flow chart illustrating one example of a method for providing clinical data representative of a concentration of a blood analyte in a patient. In some embodiments, the method may be implemented on a computing device, such as the display device shown in Figure 1 or one of the electronic devices shown in Figures 2 and 3. In some cases, the method may be performed by a continuous analyte monitoring application installed on the computing device.

[0297] At block 601, a signal is received from a continuous analyte sensor located in the interstitial fluid of a patient. The continuous analyte sensor may be any suitable type of analyte sensor, including but not limited to those described herein. At block 603, two or more factors affecting the signal are independently modeled. The factors arise from individualized characteristics of the sensor and / or individualized physiological characteristics of the patient. Individualized characteristic data associated with the individualized sensor characteristics and / or the individualized physiological patient characteristics is then received at block 605. At block 607, one or more models of at least one of the independently modeled factors are modified in response to receiving the individualized characteristic data. In response to receiving the individualized characteristic data, at block 609, clinical data representative of an analyte concentration in the patient's blood is output based at least in part on the modified model.

[0298] q. Factors affecting sensor signals Generally, any number of factors can be independently modeled; that is, in some embodiments, two, three, four, or even five or more factors may be independently modeled. Examples of factors that may be modeled are provided below for illustrative purposes only, and not as a limitation on the techniques described herein. Furthermore, for illustrative purposes only, the factors are described in the context of a diffusion-based glucose oxidase electrochemical sensor of the type described above. More generally, of course, any factor applicable to any other type of analyte sensor may be independently modeled using the techniques described herein.

[0299] The sensor's individualized factors that may be modeled are those of an individual sensor that are determined by its characteristic physical properties, such as the material from which it is formed and its chemical properties. These properties may be obtained, for example, using factory-derived sensor information, field-derived sensor information, algorithmically-derived sensor information, etc., which are described in more detail elsewhere herein.

[0300] The individualized physiological characteristics are characteristics of the patient's individual physiology, such as differences in the patient's interstitial and intradermal compartments and / or local wound healing response, as described in more detail elsewhere herein.

[0301] In the context of a diffusion-based glucose oxidase electrochemical sensor, some of the factors that can be modeled independently include, but are not limited to, H2O2 kinetics; enzyme kinetics and glucose diffusion kinetics through the resistive layer; electrochemical break-in kinetics; steady-state background components and changes in steady-state background components; IG-BG kinetics (which can be modeled independently as local cellular consumption at the insertion site); early sensor session wound healing response kinetics drop and recovery, and gradual decline in wound healing over time. Many of the aforementioned factors are phenomena that affect the signal and vary based in part on the individual physiological characteristics of the patient and the individual sensor characteristics.

[0302] In some embodiments, some of the factors modeled independently are factors that primarily affect the analyte (e.g., glucose) component of the raw electronic signal, while other factors primarily affect the background (non-analyte) component. These factors may result from various characteristics of the sensor. For example, the analyte signal components that depend on the physiochemical properties of the sensor may be modeled in whole or in part depending on the sensor design and available factory information. For example, as described in more detail below, the properties of the resistive layer control the diffusion rates of glucose and oxygen, the properties of the enzyme layer determine the rate of peroxide production, and the properties of the electrode layer determine the rate of peroxide reduction to generate a current. In one implementation, factors such as the kinetics of the measured species (H2O2), the enzyme reaction, and / or glucose diffusion through the sensor membrane may be modeled separately.

[0303] With regard to modeling the kinetics of H2O2, various sources of hydrogen peroxide can be modeled (e.g., using nonlinear partial differential equations), which can include parameters that can be fitted such as membrane thickness, resistivity / sensitivity of the membrane system, amount of enzyme in the membrane system, etc., as described in more detail elsewhere herein.

[0304] With respect to modeling the enzymatic (GOX) reaction with glucose, the enzyme activity may be modeled using the Michaelis-Menten equation, which may include parameters that are adapted based on the levels of glucose or oxygen exposure over the life of the sensor, as described in more detail elsewhere herein.

[0305] In some implementations of the in vivo sensor, the diffusion of glucose through the membrane layer can be modeled as a time-lag component using methods such as Fick's second law, which can be simplified and approximated by a first-order polynomial, with one parameter representing the time lag. This model can be applied to determine how glucose diffuses from the capillaries into the interstitial fluid and how glucose diffuses from the interstitial fluid into the interior of the sensor. In any of these models, parameters that can be fitted to the sensor characteristics can be measured or calculated, particularly during sensor fabrication. For example, measurements of physical dimensions may be related to the membrane deposition and / or final sensor sensitivity measurements.

[0306] In some embodiments, the sensor-dependent time lag factors that affect the analyte signal components may be caused by the physical and thermodynamic properties of the sensor. These properties may be modeled, for example, based on a simple time shift model, a transfer function of a diffusion process, or deconvolution. In one implementation, several different time lag sensitivity models may be defined, for example, bi-exponential, mono-exponential, and linear models. However, as will be appreciated by those skilled in the art, a continuum of models may also be used. In other implementations, the sensor-dependent time lag factors may be related to individualized sensor characteristics (e.g., film thickness) and may be modeled using parameters that may change over time, as described in more detail elsewhere herein.

[0307] In some embodiments, factors affecting the analyte signal components may be modeled based on factory derived sensor characteristics to obtain a sensitivity curve for the sensor over a sensor session, for example as described in more detail in U.S. Patent Application Serial No. 14 / 144,343, incorporated herein by reference in its entirety. These disclosures describe a method of factory calibration that uses sensitivity measurements from sensor manufacture to predict sensitivity values ​​over the in vivo life of the sensor based on a (e.g., linear) transformation. This method can be used herein to provide a linear combination of parameters measured at the factory (during or after sensor manufacture) and parameters measured in real time (measured or calculated during in vivo use). Using a combination of factory and real time measurements in modeling factors affecting the analyte signal components can improve the performance of the system, as the linear combination allows for a link between factory information and in vivo sensor measurements.

[0308] In one exemplary implementation, diffusion, enzyme activity, and H2O2 activity are individually modeled based on individualized sensor characteristic data such as manufacturing parameters including (but not limited to) wire dimensions, film thickness, calibration check waveform, and calibration check temperature, thereby improving signal modeling during warm-up, end of life, and temperature compensation. Mass conservation law may be used, which dictates that in unit space, the rate of change of glucose concentration is equal to glucose concentration inflow-glucose concentration outflow-glucose concentration consumed. This conservation law and model may be complex for practical use, but given knowledge of the sensor design, it can be significantly simplified for use in practical models. For example, the second derivative from the original model may be approximated by a first order polynomial, given the sensor resistive layer design. In addition, the model may also account for the activity of H2O2 diffusion to the working electrode, and properties of the working electrode that affect how it absorbs electrons.

[0309] In some embodiments, factors influencing the background (non-analyte) signal component result from electrochemical break-in, which is an individualized sensor characteristic resulting from electrochemical precipitation during start-up. Electrochemical break-in can be modeled using parameters that can vary due to factory-derived sensor measurements, time since sensor insertion, etc. For example, in one exemplary implementation, break-in may be modeled for the first hour after sensor insertion to allow for faster sensor start-up. In this example, the break-in model is based on background or non-enzymatic sensor signals. High frequency sampling data can be used immediately after sensor insertion to model hydration and non-glucose signal break-in and predict future non-glucose signal break-in. The model and optimization can be trained with non-enzymatic sensor data from the human body. This is because within a limited time immediately after insertion, the sensor and enzymes are not fully functional and the majority of the signal is dominated by the non-glucose signal.

[0310] The resulting break-in signal can be subtracted from the analyte signal, which can be further processed based on modeling of sensitivity (from individualized sensor data) and compartment bias (from individualized physiological data) to adaptively modify the output, thereby reducing the influence of individual factors. The observed break-in can be modeled as a function of hydration (a function of time to model hydration behavior, transportability of ions) and break-in (a function of time to model ions and corresponding currents generated since the transmitter was powered on). The functions (hydration and break-in) can be any function, including but not limited to multiplication, convolution, and / or nested integrals. In the example of hydration, the model can be an increasing function over time from 0% to 100%, including but not limited to any cumulative density function, such as the cumulative distribution function (CDF) of normal, logistic, exponential, log-logistic distribution. In one example, the model can be a decreasing function over time from a large real value to a constant value.

[0311] In some implementations of the in vivo sensor, factors influencing the background signal may result from systemic and / or local responses associated with personalized physiological characteristics produced by the physiological species, which may be modeled individually or collectively, for example, based on personalized modeling from previous sessions that may be adaptively modified based on the personalized physiological characteristic information.

[0312] In some implementations of H2O2 measurement sensors, systemic levels of interfering effects or species (depending on the sensor design) can be modeled as a steady-state signal, e.g., including parameters that change over time due to changes in the level of the interfering effect or species. Interfering effects or species may also be considered as a steady-state background signal (impact) from a timing perspective, which can be modeled after sensor break-in. In some cases, an increase or decrease in the steady-state background signal can be modeled.

[0313] In some implementations of in vivo sensors, the individualized physiological characteristics may cause a localized response to the sensor (e.g., oxidative species found at the insertion site induced by a foreign body / wound healing response) that may include parameters that may be modeled and matched to an individual's metabolic or wound healing response at the insertion site. For example, some patients have a stronger physiological response at a particular or any sensor insertion site compared to the average population.

[0314] In some implementations, factors resulting from electrochemical break-in can be modeled separately as factors associated with the break-in of the interference layer and factors associated with the break-in of the platinum surface.

[0315] In general, modeling of the raw electronic signal relative to interstitial glucose level can be collectively and / or individually modeled, and the modeling may be based on information derived from sensor manufacturing. Specifically, the background signal (including break-in effects, steady-state background, and / or changes in steady-state background) and / or the analyte signal (including kinetics of the measured species, enzymatic reactions, and / or diffusion of glucose through the membrane) can be modeled based on a priori information measured or calculated from sensor factory derived information.

[0316] In some implementations, modeling of error and validation parameters may also be provided. For example, wedges representing the range of acceptable slope values ​​at any given time, previously modeled based on a priori data, e.g., clinical data from trials, may be used to model factors that affect the signal.

[0317] In some embodiments, the modeling of interstitial glucose (IG) versus blood glucose (BG) includes one or more components influenced by individualized differences in diffusion of glucose from the interstitium to the blood compartment and / or cellular consumption of glucose around the sensor site, which may be modeled in whole or in part, as described in more detail elsewhere herein. In some embodiments, the models and model parameters are optimized based on parameters related to physiology, which may be captured using averages across population data, particularly averages of large populations across different ages, BMIs, etc. For example, effects such as time lag, temperature regulation, wound healing, and foreign body reactions may be correlated across specific age and / or BMI ranges. In some implementations, the time lag associated with diffusion of glucose between the interstitial and blood compartments is a factor resulting from individualized physiological characteristics that may be modeled using parameters that may vary due to, for example, the sensor insertion location. The temperature of the body at the insertion site is also a factor resulting from individualized physiological characteristics that may be modeled.

[0318] Another individualized physiological characteristic that gives rise to a factor that can be modeled is the cellular consumption of glucose around the sensor site. This characteristic can arise, for example, due to wound healing (or foreign body) response to sensor insertion, which can be modeled and has parameters that can vary, for example, due to patient BMI and / or sensor insertion site. In some cases, the sudden decrease in signal during the first 24 hours after sensor insertion caused by the body's initial response to sensor insertion and the eventual recovery is an individualized physiological characteristic that gives rise to a factor that can be modeled and has parameters that can vary, for example, due to BMI and / or sensor insertion site. In some cases, the gradual decline in signal associated with the end of the sensor's life is an individualized physiological characteristic that can be modeled and has parameters that can vary due to an individual's wound healing response.

[0319] As mentioned above, one factor that may be modeled independently includes IG-BG kinetics. In some embodiments, IG-BG kinetics may be modeled in its entirety or broken down into one or more additional factors (i.e., sub-factors). Each of these factors (or sub-factors) represents a different characteristic of the sensor that affects signal accuracy in different ways depending on a wide variety of root causes, and they may be modeled independently and / or combined together in various ways, for example, to balance efficiency of processing and optimality of accuracy improvement for a particular implementation, as can be understood by one of ordinary skill in the art.

[0320] The relationship between blood and interstitial glucose in a given subject, or even at a given location, can be related to how well vascularized the tissue area is, or how quickly glucose diffuses from the vasculature to the sensor site, which may be influenced by other anatomical or physiological factors. These factors, which arise from the subject's individual characteristics, can affect the relationship between blood glucose and interstitial glucose, both in terms of time lag and the observed relationship ("slope") between blood glucose and the sensor signal. In an exemplary model of blood glucose to interstitial glucose kinetics, interstitial glucose can include a low-pass filtered version of blood glucose.

[0321] Another factor related to IG-BG kinetics that may be modeled relates to compartment bias, a term that may be used to describe one or more effects that provide a certain bias between the interstitial space and the blood (e.g., capillary) space in an individual. Thus, the combined compartment bias may be used to describe interstitial glucose and convert it to a clinical blood glucose value based on data obtained from clinical studies or other reference sources. However, the modeling of this factor may be adaptive based on the received characteristic data.

[0322] In some embodiments, IG-BG kinetics can be modeled using a modified two-compartment model. For example, compartment biases may be defined and transformed by both a steady-state "baseline" compartment bias (a constant bias between BG and IG) and a time lag compensation. In particular, blood glucose is

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[0323] In some embodiments, the steady-state "baseline" compartment bias can be modeled as a dynamic baseline rather than a steady-state baseline to minimize the multiplicative or additive bias of the simple steady-state model. In this case, blood glucose is:

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[0324] In yet another embodiment, a Bayesian structure and prior distribution of sensitivities and other parameters (to estimate posterior sensitivity and other parameter distributions) may be useful in converting analyte signals to clinical blood analyte values. One additional parameter that may be applied as a prior distribution is compartment bias, which may be used to estimate posterior compartment bias. In these embodiments, the prior distribution variance determines how much the algorithm trusts the blood glucose value entered by the user against prior knowledge received from feasibility studies and manufacturing.

[0325] In another embodiment, as described in more detail in U.S. Patent Application No. 15 / 606,954, gradual sensor degradation is modeled based on the knowledge that the sensor signal will eventually degrade over time as the sensor approaches the end of its life.

[0326] In some embodiments, the modeling of the analyte signal is based on at least four pre-set parameters: initial sensitivity and final sensitivity (derived from factory calibration information and used to predict the sensitivity profile over time), as well as drop recovery compensation (lasting for the first 24 hours) and steady-state baseline, as described elsewhere herein. In one implementation, the 24-hour drop recovery compensation model is pre-optimized based on population data. In other embodiments, small time windows are pre-set, which may be in the time domain or may overlap, and may be defaulted to zero.

[0327] In some embodiments, neural network models may be applied to compartment bias profile estimates, compartment bias parameter estimates, factory-calibrated sensitivity profile estimates, and / or non-enzymatic and steady-state bias estimates (profiles or parameters).

[0328] r. Receiving characteristic sensor data The characteristic data regarding the sensor and the patient or subject can be received at the start of the sensor session prior to sensor insertion. Alternatively, or in addition, the characteristic data can be received at any time after the in vivo sensor session has begun.

[0329] In some embodiments, the characteristic sensor data may include factory-derived information. For example, the factor-derived information may include factory-provided calibration information for each individual sensor or for a group of sensors (e.g., a brick) manufactured together or under the same conditions.

[0330] Factory derived sensor data may be provided in a sensor code that can be transferred directly to the transmitter, which may be disposable and paired with only one sensor. If the transmitter is intended to be durable (the transmitter is used across multiple sensors), transfer of the sensor code can be provided through the sensor wire and the transmitter contacts (e.g., in some implementations, the conductive puck) by providing a disposable housing with contacts (e.g., the conductive puck) with different conductivities. The impedance response (or transient response to a step change in bias potential) of the contacts is distinguishable. For example, if three distinct contact materials are available - P1, P2, P3 - then for a given disposable housing, the two contacts may be selected from the following combinations: (i) P1, P1 corresponds to sensor code 1; (ii) P1, P2 corresponds to sensor code 2; (iii) P1, P3 corresponds to sensor code 3; and (iv) P2, P3 corresponds to sensor code 4. The static resistance / conductivity of the contacts may be consistent or different across the various combinations. In this approach, the contacts can be modeled as parallel resistors and capacitors. Thus, during regimes in which the bias potential is dynamic, capacitive behavior dominates the response. During regimes in which the bias potential is static, resistive behavior dominates. During or after manufacturing, dimensional differences in sensor characteristics that may result from variations in manufacturing can be measured.

[0331] Factory derived information that can be used as characteristic sensor data can also include set points and measurements obtained from manufacturing process lines in the factory. This information can define sensitivity profile variations that can be used in addition to or instead of the final calibration checks typically used in factory calibration of CGM systems. Some examples of set points and measurements from manufacturing process lines in the sensor factory can include, for example, raw material lot and characteristic information such as concentration, temperature, viscosity, and lot number; film deposition settings such as deposition rate and volume; electrode dimensions such as length, width, and diameter before and after film layer deposition; age of material or final sensor product; film dimensions such as thickness and surface area; positioning in the chamber; staging time; environmental conditions such as room temperature and relative humidity during staging; temperature and relative humidity in the chamber; curing characteristics such as environmental conditions before, during, and after curing; test conditions, etc.

[0332] In some implementations, neural networks can be used to calculate underlying relationships (especially non-linear relationships) between process parameters and model parameters such as sensitivity. After fabrication, the final sensor characteristics, such as sensor sensitivity to an analyte, can be measured and recorded.

[0333] In some embodiments, impedance is used to measure an individualized sensor characteristic related to the sensitivity (slope) of the sensor's membrane. In particular, in vivo measurement of RL membrane impedance (RRL) has been shown to provide real-time glucose sensitivity information (mt). An example of the use of impedance is described, for example, in U.S. Pat. No. 10,624,568. However, since many physical factors affect impedance (e.g., temperature fluctuations and factory-derived slope (sensitivity value)), the following impedance measurement and interpretation measurements / calculations may be implemented in conjunction with impedance measurement to overcome the above challenges.

[0334] In one embodiment, membrane impedance is measured using the impulse response, i.e., the integrated pulse current (PI) as driven by a square wave pulse supplied by the on-board sensor electronics, and glucose sensitivity information (mt) may be measured as follows, and the relationship between PI and RRL is calculated using any one of the following: (i) Using only impedance (hereinafter referred to as Impd).

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[0335] In all the above equations, the integrated pulse current PI can potentially be replaced directly by the membrane resistance RL (in kΩ).

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[0336] In some embodiments, if the factory information is provided via a code or the like, the personalized sensor characteristic information may be embedded therein.

[0337] In some implementations, the factory information can be encoded on the wire itself. In some implementations, the packaging includes a factory code based on the factory derived information. For example, in one exemplary embodiment, the sensor is sterilized in ethylene oxide (EtO) and then after sterilization, a desiccant is added to the packaging. The desiccant is refillable and can be present in the package during EtO and "baked out" after EtO to refill its drying capacity. The sensor can then be packaged in a final package with a moisture barrier and / or additional desiccant. Alternatively, bulk packaging can include multiple sensors with desiccant that can be e-beam sterilized like fingerstick sensor vials.

[0338] s. Receiving characteristic physiological patient data In some implementations, the characteristic physiological patient data can be obtained from a previous sensor session or from on-board measurements of the current sensor session.

[0339] For example, one type of characteristic physiological patient data that may be obtained is an individual's compartment bias, which is the difference between interstitial glucose and blood glucose. An individual's compartment bias may be tracked from session to session or from one day to the next within a session. In some cases, the compartment bias may be a term with a separate model profile of a larger conversion algorithm and / or the conversion itself. In one example, the compartment bias may be measured on day 1 and applied to that individual for the remainder of the sensor session. In another example, the compartment bias may be measured during a sensor session and applied to subsequent sensor sessions for the same individual. Information that may be obtained for use in determining an individual's compartment bias may include BG (blood glucose) values ​​entered by the user or a connected device; field data / big data on similar patients; exercise data including heart rate data; glucose exposure over time; and the patient's temperature over time.

[0340] In some cases, individualized physiological characteristic data may be obtained from a sensor insertion verification process, which may be performed using known techniques or using on-board systems and methods. In one example, a tunnel magnetoresistance (TMR) magnet sensor detects the deployment, and then a data log ensures that the deployment is true. The sensor insertion verification may be detected or measured in real time and used as part of the sensor and / or physiological characteristic data to adjust break-in parameters.

[0341] In some cases, the personalized physiological characteristic data may include individual physiological information, including the patient's age and BMI (body mass index), which may be manually entered by the patient or automatically pulled from a connected system. The patient's age and BMI may be useful, for example, to adaptively modify parameters of the IG-BG model, as described in more detail elsewhere herein.

[0342] Individual physiological characteristic data can be calculated from the frequency components of the low pass filter and used to model IG-BG dynamics including the typical observed lag and relative glucose concentrations in blood and interstitial fluid (which affect the observed "slope"). Using the individual physiological characteristics calculated from the frequency components, both the slope and time delay are predicted and this can be used to improve glucose accuracy, for example, by refining the predicted slope, changing the aggressiveness of the glucose prediction, or changing the BG value that is matched to the sensor signal during calibration.

[0343] Although the sensors and physiological characteristic data are described in separate paragraphs above, those skilled in the art will understand that certain characteristic data may include effects that are not strictly attributable to one or the other. One example of this is in vivo impedance measurements, which may have effects from both sensor and physiological characteristics. Furthermore, while calculations and measurements are described in various examples above, those skilled in the art will understand that the measurements and calculations may be algorithmically processed, for example, using deep learning techniques such as neural networks, to estimate biases based on measured / derived variables (e.g., time, raw signal, FC sensitivity, temperature, etc.), which biases may be applied as individualized sensor and / or physiological characteristic data.

[0344] t. Modifying the model of factors affecting the sensor signal Modification of the model of factors affecting the sensor signal can be performed when the sensor characteristic data is received at the start of a sensor session and / or during a sensor session as personalized physiological and / or sensor characteristic data is collected and / or processed from the system.

[0345] In some embodiments, factory-derived sensor characterization information may be used to model the sensitivity curve of the sensor over a sensor session, as described in more detail in U.S. Patent Application Serial No. 13 / 446,848, which is incorporated herein by reference in its entirety. In those embodiments in which the analyte signal may be modeled based on factory-derived code that interprets the signal based on pre-set model parameters (e.g., generally designed for an average patient population), the systems and methods described herein modify the pre-set model parameters based on received individualized sensor characterization data, such as set points and measurements from a manufacturing process line at the factory, as described in more detail elsewhere herein.

[0346] In some implementations of diffusion-based glucose oxidase electrochemical sensors, where the diffusion or loss of hydrogen peroxide in and around the sensor may be modeled based on pre-set model parameters, the systems and methods described herein may adaptively modify those pre-set model parameters based on received individualized sensor characteristics, such as measures of cumulative exposure to a particular temperature, glucose, and / or oxygen concentration (in vivo) over time. Models that may be used include linear regression models to estimate hydrogen peroxide loss, where a diffusion model (Fick's law), diffusion coefficient, and electrochemical reaction model may be applied to the working electrode.

[0347] FIG. 7 shows the sensitivity change from temperature excursions resulting in an increase and decrease in sensor sensitivity before sensor insertion. The data shows how a sensitivity change of approximately 3.2% / C occurs when the temperature is temporarily deviated from 36° C. Without being bound by theory, it is believed that a similar sensitivity change appears in vivo when the patient's body temperature changes over time, for example due to a fever. This is further illustrated in FIG. 8, which shows how the model of sensitivity over time changes based on temperature over time. The data shows a sensitivity change of approximately 4% per degree Celsius when the temperature deviates from 36° C., which is higher than the temporary temperature excursion. In some implementations, the sensor takes longer to stabilize when exposed to lower temperatures, which can be modeled as part of the sensitivity component. In some implementations, the sensor sensitivity is lower when exposed to higher glucose concentrations over time, which can be modeled as part of the sensitivity component.

[0348] During in vitro testing of the sensor for sensitivity to glucose exposure over time, it was found that when the sensor was moved from high to low glucose, the initial sensitivity was higher than steady state. In contrast, when the sensor was moved from low to high glucose, the initial sensitivity was lower than steady state.

[0349] As previously described with respect to modeling the enzyme (GOX) reaction with glucose, the enzyme activity may be modeled using the Michaelis-Menten equation, which may include parameters that are adapted based on the level of glucose or oxygen exposure over the life of the sensor, as described in more detail elsewhere herein. In this case, the systems and methods described herein may adaptively modify their pre-set model parameters based on received personalized sensor characteristic data, such as data collected from a factory calibration check, which may be used to estimate enzyme performance. Personalized physiological characteristic data may also be useful in adaptively modifying model parameters, for example, measurements of cumulative exposure to glucose and / or oxygen concentrations (in vivo) over time.

[0350] In some implementations of the in vivo sensor, the diffusion of glucose through the membrane layer (particularly the resistive layer) can be modeled based on pre-set model parameters. The systems and methods described herein can adaptively modify those pre-set model parameters based on received individualized sensor characteristic data such as the simplified Fick's second law, Michaelis-Menten rate, diffusion coefficient, Michaelis-Menten constant, maximum enzyme reaction rate, sensor resistive layer thickness, sensor enzyme layer thickness, sensor wire interference layer thickness, sensor wire dimensions, etc., using elemental charge and Avogadro's number to convert the signal in pA to a molar amount of glucose.

[0351] In embodiments where the sensor-dependent time lag is modeled, the physical and thermodynamic properties of the sensor can be used to adaptively modify the sensor-dependent time lag model based on measurements or calculations taken during or after manufacturing that are related to dimensional differences in the sensor design. As can be appreciated by one skilled in the art, this adaptive modification can scale, bias, or modify the sensor-dependent time lag model in response to receiving characteristic information. For example, the physical dimensional properties can be used to scale the magnitude and / or thermodynamic properties of the model curve, which can result in a modification of the "starting point" on the model curve.

[0352] In some implementations where electrochemical break-in may be modeled based on pre-set model parameters, the systems and methods described herein can adaptively modify those pre-set model parameters based on received individualized sensor characteristic data, such as factory-derived sensor measurements, field data, time since sensor insertion, and information obtained during sensor insertion verification.

[0353] In some implementations, the error and validation models may include pre-set parameters, and the systems and methods described herein adaptively modify those pre-set model parameters based, for example, on received sensor and / or physiological characteristic data. In these implementations, one or more of the following parameters may be adaptively modified: weighting of BG calibration input data, selection of error model (e.g., drop and recovery), range of acceptable sensitivity and / or background signal, determination of how to match real-time patient input with sensor-derived data (e.g., BG-to-sensor match), and number of acceptable BG values ​​available for calibration or validation. Additionally or alternatively, aberrations, noise conditions, PSD likelihood, time lag, and high ROC may be considered as well, allowing for a more individualized (and accurate) assessment of sensor health and improving confidence in identifying outlier BG values.

[0354] In some implementations of in vivo sensors, the systemic and / or local responses generated by physiological species can be modeled based on pre-set model parameters. The systems and methods described herein can adaptively modify these pre-set model parameters based on received physiological characteristic data, e.g., oxygen concentration over time, which may be measured in vivo. High frequency measurements, such as those obtained during the sensor insertion validation process, can be used to modify the hydration model and the non-glucose break-in model. In some implementations, the systemic level of an interfering species (e.g., ascorbic acid or uric acid for some H2O2 measuring sensors) may be modeled as a steady-state signal, which may include parameters that change over time, e.g., due to changes in the levels of ascorbic acid, uric acid, or other measurements or calculations of personal characteristic data. This systemic effect can be considered, from a timing perspective, as a steady-state background signal that settles after sensor break-in, which can be modeled with parameters that can change due to calculated and / or measured individual sensor characteristic data, e.g., oxygen concentration. The increase or decrease in steady state can be modeled, and the model can be adaptively modified based on BMI, age, sex, etc. to adjust for the steady-state compartment baseline.

[0355] In some implementations of in vivo sensors, localized responses to the sensor (e.g., oxidative species found at the insertion site induced by a foreign body / wound healing response) may be modeled, including parameters that may vary due to an individual's metabolic or wound healing response at the insertion site, which may be calculated and / or measured in the model and adaptively adjusted. For example, some patients have stronger physiological responses at certain sensor insertion / sites. Interstitial glucose (...

Claims

1. 1. A method for providing data indicative of a concentration of an analyte in a patient, comprising: (ii) receiving a signal from an analyte sensor located within the patient; and (iii) independently modeling at least one factor influencing the signal, the at least one factor resulting from individualized characteristics of the sensor and / or individualized physiological characteristics of the patient; (iv) receiving personalized characteristic data associated with the personalized characteristic of the sensor and / or the personalized physiological characteristic of the patient; (v) modifying a model of one or more of the at least one factor that is independently modeled based on said receiving the individualized characteristic data; and (vi) outputting data representative of the concentration of the analyte in the patient based at least in part on the modified one or more models.

2. The method of claim 1 , wherein the factors that are modeled independently include diffusion time lag, diffusion enzyme activity, and / or IG-BG kinetics.

3. The method of claim 1 , wherein the factors that are independently modeled include the sensitivity of the sensor and / or the baseline response of the sensor.

4. The method of claim 1, wherein the analyte sensor is an enzyme-based electrochemical sensor, and the enzyme-based electrochemical sensor uses a glucose oxidase enzyme.

5. The method of claim 1 , wherein the individualized sensor characteristic comprises a physiochemical characteristic of the analyte sensor.

6. 2. The method of claim 1 , wherein modeling electrochemical break-in factors comprises modeling factors associated with an interference layer of the analyte sensor independent of factors associated with a catalytic surface of the analyte sensor.

7. The method of claim 1 , wherein the individualized physiological patient characteristic that modifies one or more of the models is a change in the signal during cellular consumption of the analyte surrounding an insertion site of the sensor.

8. The method of claim 2, wherein modeling the IG-BG dynamics comprises modeling compartment bias.

9. 9. The method of claim 8, wherein modeling the compartment biases comprises modeling steady-state compartment bias components separately from time-lag compartment bias components.

10. 2. The method of claim 1, wherein the factors that are modeled independently include gradual sensor degradation, which causes the signal to degrade as the analyte sensor approaches the end of its life.

11. The method of claim 1 , wherein the individualized characteristic data comprises data received prior to sensor insertion.

12. The method of claim 1 , wherein at least a portion of the individualized characteristic data is received after an in-vivo sensor session begins.

13. The method of claim 1 , wherein the personalized characteristic data associated with the personalized sensor characteristic comprises factory-derived information.

14. The method of claim 1 , wherein one of the factors modeled is diffusion of the analyte through one or more membrane layers of the analyte sensor.

15. The method of claim 2 , wherein the diffusion time lag factor is modeled by physical and thermodynamic sensor properties.

16. The method of claim 2 , wherein the diffusion time lag factor is modeled based on a time shift model, a transfer function of a diffusion process, or deconvolution.

17. The method of claim 1 , wherein modifying the one or more models comprises modifying a model of diffusion or loss of hydrogen peroxide in and around the sensor.

18. 20. The method of claim 17, wherein the model of diffusion or loss of hydrogen peroxide in and around the sensor is based on pre-set model parameters.

19. 20. The method of claim 18, wherein modifying the model of hydrogen peroxide diffusion or loss in and around the sensor comprises adaptively modifying the pre-set model parameters based on the individualized sensor characteristics.

20. 20. The method of claim 19, wherein the individualized sensor characteristic comprises a measure of the analyte sensor's cumulative exposure over time to in vivo temperature, glucose, and / or oxygen concentration.

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