US20170184527A1 - Sensor systems, devices, and methods for continuous glucose monitoring - Google Patents

Sensor systems, devices, and methods for continuous glucose monitoring Download PDF

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US20170184527A1
US20170184527A1 US14/980,185 US201514980185A US2017184527A1 US 20170184527 A1 US20170184527 A1 US 20170184527A1 US 201514980185 A US201514980185 A US 201514980185A US 2017184527 A1 US2017184527 A1 US 2017184527A1
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sensor
voltage
impedance
invention
eis
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US14/980,185
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Keith Nogueira
Taly G. Engel
Xiaolong Li
Bradley C. Liang
Rajiv Shah
Jaeho Kim
Mike C. Liu
Andy Y. Tsai
Andrea Varsavsky
Fei Yu
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Medtronic Minimed Inc
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Medtronic Minimed Inc
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Priority to US14/980,185 priority Critical patent/US20170184527A1/en
Assigned to MEDTRONIC MINIMED, INC. reassignment MEDTRONIC MINIMED, INC. ASSIGNMENT OF ASSIGNORS INTEREST (SEE DOCUMENT FOR DETAILS). Assignors: SHAH, RAJIV
Assigned to MEDTRONIC MINIMED, INC. reassignment MEDTRONIC MINIMED, INC. ASSIGNMENT OF ASSIGNORS INTEREST (SEE DOCUMENT FOR DETAILS). Assignors: TSAI, ANDY Y., YU, FEI, ENGEL, TALY G., VARSAVSKY, ANDREA, KIM, JAEHO, LIU, MIKE C., NOGUEIRA, KEITH, LIANG, BRADLEY C., LI, XIAOLONG
Priority claimed from EP16751061.9A external-priority patent/EP3397159A1/en
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    • GPHYSICS
    • G01MEASURING; TESTING
    • G01NINVESTIGATING OR ANALYSING MATERIALS BY DETERMINING THEIR CHEMICAL OR PHYSICAL PROPERTIES
    • G01N27/00Investigating or analysing materials by the use of electric, electro-chemical, or magnetic means
    • G01N27/02Investigating or analysing materials by the use of electric, electro-chemical, or magnetic means by investigating the impedance of the material
    • G01N27/026Dielectric impedance spectroscopy
    • GPHYSICS
    • G01MEASURING; TESTING
    • G01NINVESTIGATING OR ANALYSING MATERIALS BY DETERMINING THEIR CHEMICAL OR PHYSICAL PROPERTIES
    • G01N27/00Investigating or analysing materials by the use of electric, electro-chemical, or magnetic means
    • G01N27/26Investigating or analysing materials by the use of electric, electro-chemical, or magnetic means by investigating electrochemical variables; by using electrolysis or electrophoresis
    • G01N27/416Systems
    • G01N27/4163Systems checking the operation of, or calibrating, the measuring apparatus
    • GPHYSICS
    • G01MEASURING; TESTING
    • G01NINVESTIGATING OR ANALYSING MATERIALS BY DETERMINING THEIR CHEMICAL OR PHYSICAL PROPERTIES
    • G01N33/00Investigating or analysing materials by specific methods not covered by groups G01N1/00 - G01N31/00
    • G01N33/48Biological material, e.g. blood, urine; Haemocytometers
    • G01N33/483Physical analysis of biological material
    • G01N33/487Physical analysis of biological material of liquid biological material
    • G01N33/48707Physical analysis of biological material of liquid biological material by electrical means
    • GPHYSICS
    • G01MEASURING; TESTING
    • G01NINVESTIGATING OR ANALYSING MATERIALS BY DETERMINING THEIR CHEMICAL OR PHYSICAL PROPERTIES
    • G01N33/00Investigating or analysing materials by specific methods not covered by groups G01N1/00 - G01N31/00
    • G01N33/48Biological material, e.g. blood, urine; Haemocytometers
    • G01N33/483Physical analysis of biological material
    • G01N33/487Physical analysis of biological material of liquid biological material
    • G01N33/49Blood

Abstract

Electrochemical impedance spectroscopy (EIS) may be used in conjunction with continuous glucose monitoring (CGM) to enable identification of valid and reliable sensor data, as well implementation of Smart Calibration algorithms.

Description

    FIELD OF THE INVENTION
  • Embodiments of this invention are related generally to subcutaneous and implantable sensor devices and, in particular embodiments, to systems, devices, and methods for continuous glucose monitoring (CGM).
  • BACKGROUND OF THE INVENTION
  • Over the years, a variety of sensors have been developed for detecting and/or quantifying specific agents or compositions in a patient's blood, which enable patients and medical personnel to monitor physiological conditions within the patient's body. Illustratively, subjects may wish to monitor blood glucose levels in a subject's body on a continuing basis. Thus, glucose sensors have been developed for use in obtaining an indication of blood glucose levels in a diabetic patient. Such readings are useful in monitoring and/or adjusting a treatment regimen which typically includes the regular administration of insulin to the patient.
  • Presently, a patient can measure his/her blood glucose (BG) using a BG measurement device (i.e., glucose meter), such as a test strip meter, a continuous glucose measurement system (or a continuous glucose monitor), or a hospital hemacue. BG measurement devices use various methods to measure the BG level of a patient, such as a sample of the patient's blood, a sensor in contact with a bodily fluid, an optical sensor, an enzymatic sensor, or a fluorescent sensor. When the BG measurement device has generated a BG measurement, the measurement is displayed on the BG measurement device.
  • Current continuous glucose measurement systems include subcutaneous (or short-term) sensors and implantable (or long-term) sensors. Sensors have been applied in a telemetered characteristic monitor system. As described, e.g., in commonly-assigned U.S. Pat. No. 6,809,653, the entire contents of which are incorporated herein by reference, a telemetered system using an electrochemical sensor includes a remotely located data receiving device, a sensor for producing signals indicative of a characteristic of a user, and a transmitter device for processing signals received from the sensor and for wirelessly transmitting the processed signals to the remotely located data receiving device. The data receiving device may be a characteristic monitor, a data receiver that provides data to another device, an RF programmer, a medication delivery device (such as an infusion pump), or the like.
  • Regardless of whether the data receiving device (e.g., a glucose monitor), the transmitter device, and the sensor (e.g., a glucose sensor) communicate wirelessly or via an electrical wire connection, a characteristic monitoring system of the type described above is of practical use only after it has been calibrated based on the unique characteristics of the individual user. According to the current state of the art, the user is required to externally calibrate the sensor. More specifically, and in connection with the illustrative example of a diabetic patient, the latter is required to utilize a finger-stick blood glucose meter reading an average of two-four times per day for the duration that the characteristic monitor system is used. Each time, blood is drawn from the user's finger and analyzed by the blood glucose meter to provide a real-time blood sugar level for the user. The user then inputs this data into the glucose monitor as the user's current blood sugar level which is used to calibrate the glucose monitoring system.
  • Such external calibrations, however, are disadvantageous for various reasons. For example, blood glucose meters are not perfectly accurate and include inherent margins of error. Moreover, even if completely accurate, blood glucose meters are susceptible to improper use; for example, if the user has handled candy or other sugar-containing substance immediately prior to performing the finger stick, with some of the sugar sticking to the user's fingers, the blood sugar analysis will result in an inaccurate blood sugar level indication. Furthermore, there is a cost, not to mention pain and discomfort, associated with each application of the finger stick.
  • The current state of the art in continuous glucose monitoring (CGM) is largely adjunctive, meaning that the readings provided by a CGM device (including, e.g., an implantable or subcutaneous sensor) cannot be used without a reference value in order to make a clinical decision. The reference value, in turn, must be obtained from a finger stick using, e.g., a BG meter. The reference value is needed because there is a limited amount of information that is available from the sensor/sensing component. Specifically, the only pieces of information that are currently provided by the sensing component for processing are the raw sensor value (i.e., the sensor current or Isig) and the counter voltage. Therefore, during analysis, if it appears that the raw sensor signal is abnormal (e.g., if the signal is decreasing), the only way one can distinguish between a sensor failure and a physiological change within the user/patient (i.e., glucose level changing in the body) is by acquiring a reference glucose value via a finger stick. As is known, the reference finger stick is also used for calibrating the sensor.
  • The art has searched for ways to eliminate or, at the very least, minimize, the number of finger sticks that are necessary for calibration and for assessing sensor health. However, given the number and level of complexity of the multitude of sensor failure modes, no satisfactory solution has been found. At most, diagnostics have been developed that are based on either direct assessment of the Isig, or on comparison of multiple Isigs, e.g., from redundant and/or orthogonally redundant, sensors and/or electrodes. In either case, because the Isig tracks the level of glucose in the body, by definition, it is not analyte independent. As such, by itself, the Isig is not a reliable source of information for sensor diagnostics, nor is it a reliable predictor for continued sensor performance.
  • Another limitation that has existed in the art thus far has been the lack of sensor electronics that can not only run the sensor, but also perform real-time sensor and electrode diagnostics, and do so for redundant electrodes, all while managing the sensor's power supply. To be sure, the concept of electrode redundancy has been around for quite some time. However, up until now, there has been little to no success in using electrode redundancy not only for obtaining more than one reading at a time, but also for assessing the relative health of the redundant electrodes, the overall reliability of the sensor, and the frequency of the need, if at all, for calibration reference values.
  • The art has also searched for more accurate and reliable means for providing self-calibrating sensors, and for performing sensor diagnostics by developing a variety of circuit models. In such models, an attempt is generally made to correlate circuit elements to parameters that may be used in intelligent diagnostics, gross failure analysis, and real-time self-calibrations. However, most such models have had limited success thus far.
  • For each of the short-term sensors and the long-term sensors, a patient has to wait a certain amount of time in order for the continuous glucose sensor to stabilize and to provide accurate readings. In many continuous glucose sensors, the subject must wait three hours for the continuous glucose sensor to stabilize before any glucose measurements are utilized. This is an inconvenience for the patient and in some cases may cause the patient not to utilize a continuous glucose measurement system.
  • Further, when a glucose sensor is first inserted into a patient's skin or subcutaneous layer, the glucose sensor does not operate in a stable state. The electrical readings from the sensor, which represent the glucose level of the patient, vary over a wide range of readings. Thus, the sensor must first be stabilized. It is also desirable to allow electrodes of the sensor to be sufficiently “wetted” or hydrated before utilization of the electrodes of the sensor. If the electrodes of the sensor are not sufficiently hydrated, the result may be inaccurate readings of the patient's physiological condition. A user of current blood glucose sensors is instructed to not power up the sensors immediately. If they are utilized too early, current blood glucose sensors do not operate in an optimal or efficient fashion.
  • SUMMARY
  • According to an embodiment of the invention, a method for real-time calibration of a glucose sensor for measuring the level of glucose in a body of a user, the sensor having physical sensor electronics, a microcontroller, and a working electrode, comprises: measuring, by the physical sensor electronics, the electrode current (Isig) for the working electrode; obtaining a blood glucose (BG) value for the user; calculating, by the microcontroller, an expected calibration factor (CF) value based on the glucose sensor's age; and calculating, by the microcontroller, a calibrated sensor glucose (SG) value associated with the Isig based on the CF and BG values.
  • In accordance with another embodiment of the invention, a method for determining the validity of glucose sensor data, the glucose sensor having physical sensor electronics, a microcontroller, and a working electrode, and being in operational contact with a display device configured to display the data to a user, comprises: performing, by the microcontroller, an electrochemical impedance spectroscopy (EIS) procedure to obtain real impedance values for the electrode; filtering, by the microcontroller, the real impedance values; analyzing the real impedance values by the microcontroller to determine whether the values are stable; if the real impedance values are stable, comparing the most-recent real impedance value to a first threshold value; and based on the comparison, determining whether the sensor data is valid.
  • In accordance with another embodiment of the invention, a method for determining the validity of glucose sensor data, the glucose sensor including physical sensor electronics, a microcontroller, and a working electrode, comprises: performing, by the microcontroller, an electrochemical impedance spectroscopy (EIS) procedure to obtain imaginary impedance values for the electrode; setting a threshold reference for the imaginary impedance values; calculating a change value as a difference between the threshold reference and the most-recent imaginary impedance value; obtaining measurements of the calibration factor for the sensor; comparing, by the microcontroller, the change value to a first threshold and the calibration factor to a second threshold; and based on the comparison, determining whether the sensor data is valid, such that the sensor can continue to operate, or the data is invalid, such that the sensor should be terminated.
  • In yet another embodiment of the invention, a method for signal dip detection during the first 4-12 hours of glucose sensor data, the glucose sensor including physical sensor electronics, a microcontroller, and a working electrode, and being in operational contact with a display device configured to display the data to a user, comprises: performing, by the microcontroller, an electrochemical impedance spectroscopy (EIS) procedure to obtain real impedance values for the electrode; periodically measuring, by the physical sensor electronics, values of the electrode current (Isig) for the working electrode; calculating, by the microcontroller, sensor glucose (SG) values associated with the Isig values; comparing a current value of the Isig to a first threshold and the current value of the SG to a second threshold; evaluating, by the microcontroller, a trend of the real impedance values; and based on the comparison and the evaluation, determining whether a dip event exists.
  • In a further embodiment of the invention, a method for signal dip detection during the first 4 hours of glucose sensor data, the glucose sensor including physical sensor electronics, a microcontroller, and a working electrode, and being in operational contact with a display device configured to display the data to a user, comprises: performing, by the microcontroller, an electrochemical impedance spectroscopy (EIS) procedure to obtain real impedance values for the electrode; periodically measuring, by the physical sensor electronics, values of the electrode current (Isig) for the working electrode; comparing a current value of the Isig to a first threshold; evaluating, by the microcontroller, a trend of the real impedance values; and based on the comparison and the evaluation, determining whether a dip event exists.
  • In another embodiment of the invention, a method for first day calibration (FDC) of a glucose sensor for measuring the level of glucose in a body of a user, the sensor including physical sensor electronics, a microcontroller, and a working electrode, comprises: measuring, by the physical sensor electronics, the electrode current (Isig) for the working electrode; calculating, by the microcontroller, a calibration ratio (CR); comparing the calibration ratio to a threshold range; and based on the comparison, calculating a time interval until the next calibration.
  • BRIEF DESCRIPTION OF THE DRAWINGS
  • A detailed description of embodiments of the invention will be made with reference to the accompanying drawings, wherein like numerals designate corresponding parts in the figures.
  • FIG. 1 is a perspective view of a subcutaneous sensor insertion set and block diagram of a sensor electronics device according to an embodiment of the invention.
  • FIG. 2A illustrates a substrate having two sides, a first side which contains an electrode configuration and a second side which contains electronic circuitry.
  • FIG. 2B illustrates a general block diagram of an electronic circuit for sensing an output of a sensor.
  • FIG. 3 illustrates a block diagram of a sensor electronics device and a sensor including a plurality of electrodes according to an embodiment of the invention.
  • FIG. 4 illustrates an alternative embodiment of the invention including a sensor and a sensor electronics device according to an embodiment of the invention.
  • FIG. 5 illustrates an electronic block diagram of the sensor electrodes and a voltage being applied to the sensor electrodes according to an embodiment of the invention.
  • FIG. 6A illustrates a method of applying pulses during a stabilization timeframe in order to reduce the stabilization timeframe according to an embodiment of the invention.
  • FIG. 6B illustrates a method of stabilizing sensors according to an embodiment of the invention.
  • FIG. 6C illustrates utilization of feedback in stabilizing the sensors according to an embodiment of the invention.
  • FIG. 7 illustrates an effect of stabilizing a sensor according to an embodiment of the invention.
  • FIG. 8A illustrates a block diagram of a sensor electronics device and a sensor including a voltage generation device according to an embodiment of the invention.
  • FIG. 8B illustrates a voltage generation device to implement this embodiment of the invention.
  • FIG. 8C illustrates a voltage generation device to generate two voltage values according to an embodiment of the invention.
  • FIG. 8D illustrates a voltage generation device having three voltage generation systems, according to embodiments of the invention.
  • FIG. 9A illustrates a sensor electronics device including a microcontroller for generating voltage pulses according to an embodiment of the invention.
  • FIG. 9B illustrates a sensor electronics device including an analyzation module according to an embodiment of the invention.
  • FIG. 10 illustrates a block diagram of a sensor system including hydration electronics according to an embodiment of the invention.
  • FIG. 11 illustrates an embodiment of the invention including a mechanical switch to assist in determining a hydration time.
  • FIG. 12 illustrates a method of detection of hydration according to an embodiment of the invention.
  • FIG. 13A illustrates a method of hydrating a sensor according to an embodiment of the present invention.
  • FIG. 13B illustrates an additional method for verifying hydration of a sensor according to an embodiment of the invention.
  • FIGS. 14A, 14B, and 14C illustrate methods of combining hydrating of a sensor with stabilizing a sensor according to an embodiment of the invention.
  • FIG. 15A illustrates EIS-based analysis of system response to the application of a periodic AC signal in accordance with embodiments of the invention.
  • FIG. 15B illustrates a known circuit model for electrochemical impedance spectroscopy.
  • FIG. 16A illustrates an example of a Nyquist plot where, for a selected frequency spectrum from 0.1 Hz to 1000 Mhz, AC voltages plus a DC voltage (DC bias) are applied to the working electrode in accordance with embodiments of the invention.
  • FIG. 16B shows another example of a Nyquist plot with a linear fit for the relatively-lower frequencies and the intercept approximating the value of real impedance at the relatively-higher frequencies.
  • FIGS. 16C and 16D show, respectively, infinite and finite glucose sensor response to a sinusoidal working potential.
  • FIG. 16E shows a Bode plot for magnitude in accordance with embodiments of the invention.
  • FIG. 16F shows a Bode plot for phase in accordance with embodiments of the invention.
  • FIG. 17 illustrates the changing Nyquist plot of sensor impedance as the sensor ages in accordance with embodiments of the invention.
  • FIG. 18 illustrates methods of applying EIS technique in stabilizing and detecting the age of the sensor in accordance with embodiments of the invention.
  • FIG. 19 illustrates a schedule for performing the EIS procedure in accordance with embodiments of the invention.
  • FIG. 20 illustrates a method of detecting and repairing a sensor using EIS procedures in conjunction with remedial action in accordance with embodiments of the invention.
  • FIGS. 21A and 21B illustrate examples of a sensor remedial action in accordance with embodiments of the invention.
  • FIG. 22 shows a Nyquist plot for a normally-functioning sensor where the Nyquist slope gradually increases, and the intercept gradually decreases, as the sensor wear-time progresses.
  • FIG. 23A shows raw current signal (Isig) from two redundant working electrodes, and the electrodes' respective real impedances at 1 kHz, in accordance with embodiments of the invention.
  • FIG. 23B shows the Nyquist plot for the first working electrode (WE1) of FIG. 23A.
  • FIG. 23C shows the Nyquist plot for the second working electrode (WE2) of FIG. 23A.
  • FIG. 24 illustrates examples of signal dip for two redundant working electrodes, and the electrodes' respective real impedances at 1 kHz, in accordance with embodiments of the invention.
  • FIG. 25A illustrates substantial glucose independence of real impedance, imaginary impedance, and phase at relatively-higher frequencies for a normally-functioning glucose sensor in accordance with embodiments of the invention.
  • FIG. 25B shows illustrative examples of varying levels of glucose dependence of real impedance at the relatively-lower frequencies in accordance with embodiments of the invention.
  • FIG. 25C shows illustrative examples of varying levels of glucose dependence of phase at the relatively-lower frequencies in accordance with embodiments of the invention.
  • FIG. 26 shows the trending for 1 kHz real impedance, 1 kHz imaginary impedance, and relatively-higher frequency phase as a glucose sensor loses sensitivity as a result of oxygen deficiency at the sensor insertion site, according to embodiments of the invention.
  • FIG. 27 shows Isig and phase for an in-vitro simulation of oxygen deficit at different glucose concentrations in accordance with embodiments of the invention.
  • FIGS. 28A-28C show an example of oxygen deficiency-led sensitivity loss with redundant working electrodes WE1 and WE2, as well as the electrodes' EIS-based parameters, in accordance with embodiments of the invention.
  • FIG. 28D shows EIS-induced spikes in the raw Isig for the example of FIG. 28A-28C.
  • FIG. 29 shows an example of sensitivity loss due to oxygen deficiency that is caused by an occlusion, in accordance with embodiments of the invention.
  • FIGS. 30A-30C show an example of sensitivity loss due to bio-fouling, with redundant working electrodes WE1 and WE2, as well as the electrodes' EIS-based parameters, in accordance with embodiments of the invention.
  • FIG. 30D shows EIS-induced spikes in the raw Isig for the example of FIG. 30A-30C.
  • FIG. 31 shows a diagnostic procedure for sensor fault detection in accordance with embodiments of the invention.
  • FIGS. 32A and 32B show another diagnostic procedure for sensor fault detection in accordance with embodiments of the invention.
  • FIG. 33A shows a top-level flowchart involving a current (Isig)-based fusion algorithm in accordance with embodiments of the invention.
  • FIG. 33B shows a top-level flowchart involving a sensor glucose (SG)-based fusion algorithm in accordance with embodiments of the invention.
  • FIG. 34 shows details of the sensor glucose (SG)-based fusion algorithm of FIG. 33B in accordance with embodiments of the invention.
  • FIG. 35 shows details of the current (Isig)-based fusion algorithm of FIG. 33A in accordance with embodiments of the invention.
  • FIG. 36 is an illustration of calibration for a sensor in steady state, in accordance with embodiments of the invention.
  • FIG. 37 is an illustration of calibration for a sensor in transition, in accordance with embodiments of the invention.
  • FIG. 38A is an illustration of EIS-based dynamic slope (with slope adjustment) in accordance with embodiments of the invention for sensor calibration.
  • FIG. 38B shows an EIS-assisted sensor calibration flowchart involving low start-up detection in accordance with embodiments of the invention.
  • FIG. 39 shows sensor current (Isig) and 1 kHz impedance magnitude for an in-vitro simulation of an interferent being in close proximity to a sensor in accordance with embodiments of the invention.
  • FIGS. 40A and 40B show Bode plots for phase and impedance, respectively, for the simulation shown in FIG. 39.
  • FIG. 40C shows a Nyquist plot for the simulation shown in FIG. 39.
  • FIG. 41 shows another in-vitro simulation with an interferent in accordance to embodiments of the invention.
  • FIGS. 42A and 42B illustrate an ASIC block diagram in accordance with embodiments of the invention.
  • FIG. 43 shows a potentiostat configuration for a sensor with redundant working electrodes in accordance with embodiments of the invention.
  • FIG. 44 shows an equivalent AC inter-electrode circuit for a sensor with the potentiostat configuration shown in FIG. 43.
  • FIG. 45 shows some of the main blocks of the EIS circuitry in the analog front end IC of a glucose sensor in accordance with embodiments of the invention.
  • FIGS. 46A-46F show a simulation of the signals of the EIS circuitry shown in FIG. 45 for a current of 0-degree phase with a 0-degree phase multiply.
  • FIGS. 47A-47F show a simulation of the signals of the EIS circuitry shown in FIG. 45 for a current of 0-degree phase with a 90-degree phase multiply.
  • FIG. 48 shows a circuit model in accordance with embodiments of the invention.
  • FIGS. 49A-49C show illustrations of circuit models in accordance with alternative embodiments of the invention.
  • FIG. 50A is a Nyquist plot overlaying an equivalent circuit simulation in accordance with embodiments of the invention.
  • FIG. 50B is an enlarged diagram of the high-frequency portion of FIG. 50A.
  • FIG. 51 shows a Nyquist plot with increasing Cdl in the direction of Arrow A, in accordance with embodiments of the invention.
  • FIG. 52 shows a Nyquist plot with increasing a in the direction of Arrow A, in accordance with embodiments of the invention.
  • FIG. 53 shows a Nyquist plot with increasing Rp in the direction of Arrow A, in accordance with embodiments of the invention.
  • FIG. 54 shows a Nyquist plot with increasing Warburg admittance in the direction of Arrow A, in accordance with embodiments of the invention.
  • FIG. 55 shows a Nyquist plot with increasing λ in the direction of Arrow A, in accordance with embodiments of the invention.
  • FIG. 56 shows the effect of membrane capacitance on the Nyquist plot, in accordance with embodiments of the invention.
  • FIG. 57 shows a Nyquist plot with increasing membrane resistance in the direction of Arrow A, in accordance with embodiments of the invention.
  • FIG. 58 shows a Nyquist plot with increasing Rsol in the direction of Arrow A, in accordance with embodiments of the invention.
  • FIGS. 59A-59C show changes in EIS parameters relating to circuit elements during start-up and calibration in accordance with embodiments of the invention.
  • FIGS. 60A-60C show changes in a different set of EIS parameters relating to circuit elements during start-up and calibration in accordance with embodiments of the invention.
  • FIGS. 61A-61C show changes in yet a different set of EIS parameters relating to circuit elements during start-up and calibration in accordance with embodiments of the invention.
  • FIG. 62 shows the EIS response for multiple electrodes in accordance with embodiments of the invention.
  • FIG. 63 is a Nyquist plot showing the effect of Isig calibration via an increase in glucose in accordance with embodiments of the invention.
  • FIG. 64 shows the effect of oxygen (Vcntr) response on the Nyquist plot, in accordance with embodiments of the invention.
  • FIG. 65 shows a shift in the Nyquist plot due to temperature changes, in accordance with embodiments of the invention.
  • FIG. 66 shows the relationship between Isig and blood glucose in accordance with embodiments of the invention.
  • FIGS. 67A-67B show sensor drift in accordance with embodiments of the invention.
  • FIG. 68 shows an increase in membrane resistance during sensitivity loss, in accordance with embodiments of the invention.
  • FIG. 69 shows a drop in Warburg Admittance during sensitivity loss, in accordance with embodiments of the invention.
  • FIG. 70 shows calibration curves in accordance with embodiments of the invention.
  • FIG. 71 shows a higher-frequency semicircle becoming visible on a Nyquist plot in accordance with embodiments of the invention.
  • FIGS. 72A and 72B show Vcntr rail and Cdl decrease in accordance with embodiments of the invention.
  • FIG. 73 shows the changing slope of calibration curves in accordance with embodiments of the invention
  • FIG. 74 shows the changing length of the Nyquist plot in accordance with embodiments of the invention.
  • FIG. 75 shows enlarged views of the lower-frequency and the higher-frequency regions of the Nyquist plot of FIG. 74.
  • FIGS. 76A and 76B show the combined effect of increase in membrane resistance, decrease in Cdl, and Vcntr rail in accordance with embodiments of the invention.
  • FIG. 77 shows relative Cdl values for two working electrodes in accordance with embodiments of the invention.
  • FIG. 78 shows relative Rp values for two working electrodes in accordance with embodiments of the invention.
  • FIG. 79 shows the combined effect of changing EIS parameters on calibration curves in accordance with embodiments of the invention.
  • FIG. 80 shows that, in accordance with embodiments of the invention, the length of the Nyquist plot in the lower-frequency region is longer where there is sensitivity loss.
  • FIG. 81 is a flow diagram for sensor self-calibration based on the detection of sensitivity change in accordance with embodiments of the invention.
  • FIG. 82 illustrates a horizontal shift in Nyquist plot as a result of sensitivity loss, in accordance with embodiments of the invention.
  • FIG. 83 shows a method of developing a heuristic EIS metric based on a Nyquist plot in accordance with embodiments of the invention.
  • FIG. 84 shows the relationship between Rm and Calibration Factor in accordance with embodiments of the invention.
  • FIG. 85 shows the relationship between Rm and normalized Isig in accordance with embodiments of the invention.
  • FIG. 86 shows Isig plots for various glucose levels as a function of time, in accordance with embodiments of the invention.
  • FIG. 87 shows Cdl plots for various glucose levels as a function of time, in accordance with embodiments of the invention.
  • FIG. 88 shows a second inflection point for the plots of FIG. 86, in accordance with embodiments of the invention.
  • FIG. 89 shows a second inflection point for Rm corresponding to the peak in FIG. 88, in accordance with embodiments of the invention.
  • FIG. 90 shows one illustration of the relationship between Calibration Factor (CF) and Rmem+Rsol in accordance with embodiments of the invention.
  • FIG. 91A is a chart showing in-vivo results for MARD over all valid BGs in approximately the first 8 hours of sensor life, in accordance with embodiments of the invention.
  • FIG. 91B is a chart showing median ARD numbers over all valid BGs in approximately the first 8 hours of sensor life, in accordance with embodiments of the invention.
  • FIGS. 92A-92C show Calibration Factor adjustment in accordance with embodiments of the invention.
  • FIGS. 93A-93C show Calibration Factor adjustment in accordance with embodiments of the invention.
  • FIGS. 94A-94C show Calibration Factor adjustment in accordance with embodiments of the invention.
  • FIG. 95 shows an illustrative example of initial decay in Cdl in accordance with embodiments of the invention.
  • FIG. 96 shows the effects on Isig of removal of the non-Faradaic current, in accordance with embodiments of the invention.
  • FIG. 97A shows the Calibration Factor before removal of the non-Faradaic current for two working electrodes, in accordance with embodiments of the invention.
  • FIG. 97B shows the Calibration Factor after removal of the non-Faradaic current for two working electrodes, in accordance with embodiments of the invention.
  • FIGS. 98A and 98B show the effect on MARD of the removal of the non-Faradaic current, in accordance with embodiments of the invention.
  • FIG. 99 is an illustration of double layer capacitance over time, in accordance with embodiments of the invention.
  • FIG. 100 shows a shift in Rmem+Rsol and the appearance of the higher-frequency semicircle during sensitivity loss, in accordance with embodiments of the invention.
  • FIG. 101A shows a flow diagram for detection of sensitivity loss using combinatory logic, in accordance with an embodiment of the invention.
  • FIG. 101B shows a flow diagram for detection of sensitivity loss using combinatory logic, in accordance with another embodiment of the invention.
  • FIG. 102 shows an illustrative method for using Nyquist slope as a marker to differentiate between new and used sensors, in accordance with embodiments of the invention.
  • FIGS. 103A-103C show an illustrative example of Nyquist plots having different lengths for different sensor configurations, in accordance with embodiments of the invention.
  • FIG. 104 shows Nyquist plot length as a function of time, for the sensors of FIGS. 103A-103C.
  • FIG. 105 shows a flow diagram for blanking sensor data or terminating a sensor in accordance with an embodiment of the invention.
  • FIG. 106 shows a flow diagram for sensor termination in accordance with an embodiment of the invention.
  • FIG. 107 shows a flow diagram for signal dip detection in accordance with an embodiment of the invention.
  • FIG. 108A shows Isig and Vcntr as a function of time, and FIG. 108B shows glucose as a function of time, in accordance with an embodiment of the invention.
  • FIG. 109A calibration ratio as a function of time, and FIG. 109B show glucose as a function of time, in accordance with an embodiment of the invention.
  • FIGS. 110A and 110B show calibration factor trends as a function of time in accordance with embodiments of the invention.
  • FIG. 111 shows a flow diagram for First Day Calibration (FDC) in accordance with an embodiment of the invention.
  • FIG. 112 shows a flow diagram for EIS-based calibration in accordance with an embodiment of the invention.
  • DETAILED DESCRIPTION
  • In the following description, reference is made to the accompanying drawings which form a part hereof and which illustrate several embodiments of the present inventions. It is understood that other embodiments may be utilized and structural and operational changes may be made without departing from the scope of the present inventions.
  • The inventions herein are described below with reference to flowchart illustrations of methods, systems, devices, apparatus, and programming and computer program products. It will be understood that each block of the flowchart illustrations, and combinations of blocks in the flowchart illustrations, can be implemented by programming instructions, including computer program instructions (as can any menu screens described in the figures). These computer program instructions may be loaded onto a computer or other programmable data processing apparatus (such as a controller, microcontroller, or processor in a sensor electronics device) to produce a machine, such that the instructions which execute on the computer or other programmable data processing apparatus create instructions for implementing the functions specified in the flowchart block or blocks. These computer program instructions may also be stored in a computer-readable memory that can direct a computer or other programmable data processing apparatus to function in a particular manner, such that the instructions stored in the computer-readable memory produce an article of manufacture including instructions which implement the function specified in the flowchart block or blocks. The computer program instructions may also be loaded onto a computer or other programmable data processing apparatus to cause a series of operational steps to be performed on the computer or other programmable apparatus to produce a computer implemented process such that the instructions which execute on the computer or other programmable apparatus provide steps for implementing the functions specified in the flowchart block or blocks, and/or menus presented herein. Programming instructions may also be stored in and/or implemented via electronic circuitry, including integrated circuits (ICs) and Application Specific Integrated Circuits (ASICs) used in conjunction with sensor devices, apparatuses, and systems.
  • FIG. 1 is a perspective view of a subcutaneous sensor insertion set and a block diagram of a sensor electronics device according to an embodiment of the invention. As illustrated in FIG. 1, a subcutaneous sensor set 10 is provided for subcutaneous placement of an active portion of a flexible sensor 12 (see, e.g., FIG. 2), or the like, at a selected site in the body of a user. The subcutaneous or percutaneous portion of the sensor set 10 includes a hollow, slotted insertion needle 14, and a cannula 16. The needle 14 is used to facilitate quick and easy subcutaneous placement of the cannula 16 at the subcutaneous insertion site. Inside the cannula 16 is a sensing portion 18 of the sensor 12 to expose one or more sensor electrodes 20 to the user's bodily fluids through a window 22 formed in the cannula 16. In an embodiment of the invention, the one or more sensor electrodes 20 may include a counter electrode, a reference electrode, and one or more working electrodes. After insertion, the insertion needle 14 is withdrawn to leave the cannula 16 with the sensing portion 18 and the sensor electrodes 20 in place at the selected insertion site.
  • In particular embodiments, the subcutaneous sensor set 10 facilitates accurate placement of a flexible thin film electrochemical sensor 12 of the type used for monitoring specific blood parameters representative of a user's condition. The sensor 12 monitors glucose levels in the body, and may be used in conjunction with automated or semi-automated medication infusion pumps of the external or implantable type as described, e.g., in U.S. Pat. Nos. 4,562,751; 4,678,408; 4,685,903 or 4,573,994, to control delivery of insulin to a diabetic patient.
  • Particular embodiments of the flexible electrochemical sensor 12 are constructed in accordance with thin film mask techniques to include elongated thin film conductors embedded or encased between layers of a selected insulative material such as polyimide film or sheet, and membranes. The sensor electrodes 20 at a tip end of the sensing portion 18 are exposed through one of the insulative layers for direct contact with patient blood or other body fluids, when the sensing portion 18 (or active portion) of the sensor 12 is subcutaneously placed at an insertion site. The sensing portion 18 is joined to a connection portion 24 that terminates in conductive contact pads, or the like, which are also exposed through one of the insulative layers. In alternative embodiments, other types of implantable sensors, such as chemical based, optical based, or the like, may be used.
  • As is known in the art, the connection portion 24 and the contact pads are generally adapted for a direct wired electrical connection to a suitable monitor or sensor electronics device 100 for monitoring a user's condition in response to signals derived from the sensor electrodes 20. Further description of flexible thin film sensors of this general type are be found in U.S. Pat. No. 5,391,250, entitled METHOD OF FABRICATING THIN FILM SENSORS, which is herein incorporated by reference. The connection portion 24 may be conveniently connected electrically to the monitor or sensor electronics device 100 or by a connector block 28 (or the like) as shown and described in U.S. Pat. No. 5,482,473, entitled FLEX CIRCUIT CONNECTOR, which is also herein incorporated by reference. Thus, in accordance with embodiments of the present invention, subcutaneous sensor sets 10 may be configured or formed to work with either a wired or a wireless characteristic monitor system.
  • The sensor electrodes 20 may be used in a variety of sensing applications and may be configured in a variety of ways. For example, the sensor electrodes 20 may be used in physiological parameter sensing applications in which some type of biomolecule is used as a catalytic agent. For example, the sensor electrodes 20 may be used in a glucose and oxygen sensor having a glucose oxidase (GOx) enzyme catalyzing a reaction with the sensor electrodes 20. The sensor electrodes 20, along with a biomolecule or some other catalytic agent, may be placed in a human body in a vascular or non-vascular environment. For example, the sensor electrodes 20 and biomolecule may be placed in a vein and be subjected to a blood stream, or may be placed in a subcutaneous or peritoneal region of the human body.
  • The monitor 100 may also be referred to as a sensor electronics device 100. The monitor 100 may include a power source 110, a sensor interface 122, processing electronics 124, and data formatting electronics 128. The monitor 100 may be coupled to the sensor set 10 by a cable 102 through a connector that is electrically coupled to the connector block 28 of the connection portion 24. In an alternative embodiment, the cable may be omitted. In this embodiment of the invention, the monitor 100 may include an appropriate connector for direct connection to the connection portion 104 of the sensor set 10. The sensor set 10 may be modified to have the connector portion 104 positioned at a different location, e.g., on top of the sensor set to facilitate placement of the monitor 100 over the sensor set.
  • In embodiments of the invention, the sensor interface 122, the processing electronics 124, and the data formatting electronics 128 are formed as separate semiconductor chips, however, alternative embodiments may combine the various semiconductor chips into a single or multiple customized semiconductor chips. The sensor interface 122 connects with the cable 102 that is connected with the sensor set 10.
  • The power source 110 may be a battery. The battery can include three series silver oxide 357 battery cells. In alternative embodiments, different battery chemistries may be utilized, such as lithium based chemistries, alkaline batteries, nickel metalhydride, or the like, and a different number of batteries may be used. The monitor 100 provides power to the sensor set via the power source 110, through the cable 102 and cable connector 104. In an embodiment of the invention, the power is a voltage provided to the sensor set 10. In an embodiment of the invention, the power is a current provided to the sensor set 10. In an embodiment of the invention, the power is a voltage provided at a specific voltage to the sensor set 10.
  • FIGS. 2A and 2B illustrate an implantable sensor and electronics for driving the implantable sensor according to an embodiment of the present invention. FIG. 2A shows a substrate 220 having two sides, a first side 222 of which contains an electrode configuration and a second side 224 of which contains electronic circuitry. As may be seen in FIG. 2A, a first side 222 of the substrate comprises two counter electrode-working electrode pairs 240, 242, 244, 246 on opposite sides of a reference electrode 248. A second side 224 of the substrate comprises electronic circuitry. As shown, the electronic circuitry may be enclosed in a hermetically sealed casing 226, providing a protective housing for the electronic circuitry. This allows the sensor substrate 220 to be inserted into a vascular environment or other environment which may subject the electronic circuitry to fluids. By sealing the electronic circuitry in a hermetically sealed casing 226, the electronic circuitry may operate without risk of short circuiting by the surrounding fluids. Also shown in FIG. 2A are pads 228 to which the input and output lines of the electronic circuitry may be connected. The electronic circuitry itself may be fabricated in a variety of ways. According to an embodiment of the present invention, the electronic circuitry may be fabricated as an integrated circuit using techniques common in the industry.
  • FIG. 2B illustrates a general block diagram of an electronic circuit for sensing an output of a sensor according to an embodiment of the present invention. At least one pair of sensor electrodes 310 may interface to a data converter 312, the output of which may interface to a counter 314. The counter 314 may be controlled by control logic 316. The output of the counter 314 may connect to a line interface 318. The line interface 318 may be connected to input and output lines 320 and may also connect to the control logic 316. The input and output lines 320 may also be connected to a power rectifier 322.
  • The sensor electrodes 310 may be used in a variety of sensing applications and may be configured in a variety of ways. For example, the sensor electrodes 310 may be used in physiological parameter sensing applications in which some type of biomolecule is used as a catalytic agent. For example, the sensor electrodes 310 may be used in a glucose and oxygen sensor having a glucose oxidase (GOx) enzyme catalyzing a reaction with the sensor electrodes 310. The sensor electrodes 310, along with a biomolecule or some other catalytic agent, may be placed in a human body in a vascular or non-vascular environment. For example, the sensor electrodes 310 and biomolecule may be placed in a vein and be subjected to a blood stream.
  • FIG. 3 illustrates a block diagram of a sensor electronics device and a sensor including a plurality of electrodes according to an embodiment of the invention. The sensor set or system 350 includes a sensor 355 and a sensor electronics device 360. The sensor 355 includes a counter electrode 365, a reference electrode 370, and a working electrode 375. The sensor electronics device 360 includes a power supply 380, a regulator 385, a signal processor 390, a measurement processor 395, and a display/transmission module 397. The power supply 380 provides power (in the form of either a voltage, a current, or a voltage including a current) to the regulator 385. The regulator 385 transmits a regulated voltage to the sensor 355. In an embodiment of the invention, the regulator 385 transmits a voltage to the counter electrode 365 of the sensor 355.
  • The sensor 355 creates a sensor signal indicative of a concentration of a physiological characteristic being measured. For example, the sensor signal may be indicative of a blood glucose reading. In an embodiment of the invention utilizing subcutaneous sensors, the sensor signal may represent a level of hydrogen peroxide in a subject. In an embodiment of the invention where blood or cranial sensors are utilized, the amount of oxygen is being measured by the sensor and is represented by the sensor signal. In an embodiment of the invention utilizing implantable or long-term sensors, the sensor signal may represent a level of oxygen in the subject. The sensor signal is measured at the working electrode 375. In an embodiment of the invention, the sensor signal may be a current measured at the working electrode. In an embodiment of the invention, the sensor signal may be a voltage measured at the working electrode.
  • The signal processor 390 receives the sensor signal (e.g., a measured current or voltage) after the sensor signal is measured at the sensor 355 (e.g., the working electrode). The signal processor 390 processes the sensor signal and generates a processed sensor signal. The measurement processor 395 receives the processed sensor signal and calibrates the processed sensor signal utilizing reference values. In an embodiment of the invention, the reference values are stored in a reference memory and provided to the measurement processor 395. The measurement processor 395 generates sensor measurements. The sensor measurements may be stored in a measurement memory (not shown). The sensor measurements may be sent to a display/transmission device to be either displayed on a display in a housing with the sensor electronics or transmitted to an external device.
  • The sensor electronics device 360 may be a monitor which includes a display to display physiological characteristics readings. The sensor electronics device 360 may also be installed in a desktop computer, a pager, a television including communications capabilities, a laptop computer, a server, a network computer, a personal digital assistant (PDA), a portable telephone including computer functions, an infusion pump including a display, a glucose sensor including a display, and/or a combination infusion pump/glucose sensor. The sensor electronics device 360 may be housed in a blackberry, a network device, a home network device, or an appliance connected to a home network.
  • FIG. 4 illustrates an alternative embodiment of the invention including a sensor and a sensor electronics device according to an embodiment of the invention. The sensor set or sensor system 400 includes a sensor electronics device 360 and a sensor 355. The sensor includes a counter electrode 365, a reference electrode 370, and a working electrode 375. The sensor electronics device 360 includes a microcontroller 410 and a digital-to-analog converter (DAC) 420. The sensor electronics device 360 may also include a current-to-frequency converter (I/F converter) 430.
  • The microcontroller 410 includes software program code, which when executed, or programmable logic which, causes the microcontroller 410 to transmit a signal to the DAC 420, where the signal is representative of a voltage level or value that is to be applied to the sensor 355. The DAC 420 receives the signal and generates the voltage value at the level instructed by the microcontroller 410. In embodiments of the invention, the microcontroller 410 may change the representation of the voltage level in the signal frequently or infrequently. Illustratively, the signal from the microcontroller 410 may instruct the DAC 420 to apply a first voltage value for one second and a second voltage value for two seconds.
  • The sensor 355 may receive the voltage level or value. In an embodiment of the invention, the counter electrode 365 may receive the output of an operational amplifier which has as inputs the reference voltage and the voltage value from the DAC 420. The application of the voltage level causes the sensor 355 to create a sensor signal indicative of a concentration of a physiological characteristic being measured. In an embodiment of the invention, the microcontroller 410 may measure the sensor signal (e.g., a current value) from the working electrode. Illustratively, a sensor signal measurement circuit 431 may measure the sensor signal. In an embodiment of the invention, the sensor signal measurement circuit 431 may include a resistor and the current may be passed through the resistor to measure the value of the sensor signal. In an embodiment of the invention, the sensor signal may be a current level signal and the sensor signal measurement circuit 431 may be a current-to-frequency (I/F) converter 430. The current-to-frequency converter 430 may measure the sensor signal in terms of a current reading, convert it to a frequency-based sensor signal, and transmit the frequency-based sensor signal to the microcontroller 410. In embodiments of the invention, the microcontroller 410 may be able to receive frequency-based sensor signals easier than non-frequency-based sensor signals. The microcontroller 410 receives the sensor signal, whether frequency-based or non frequency-based, and determines a value for the physiological characteristic of a subject, such as a blood glucose level. The microcontroller 410 may include program code, which when executed or run, is able to receive the sensor signal and convert the sensor signal to a physiological characteristic value. In an embodiment of the invention, the microcontroller 410 may convert the sensor signal to a blood glucose level. In an embodiment of the invention, the microcontroller 410 may utilize measurements stored within an internal memory in order to determine the blood glucose level of the subject. In an embodiment of the invention, the microcontroller 410 may utilize measurements stored within a memory external to the microcontroller 410 to assist in determining the blood glucose level of the subject.
  • After the physiological characteristic value is determined by the microcontroller 410, the microcontroller 410 may store measurements of the physiological characteristic values for a number of time periods. For example, a blood glucose value may be sent to the microcontroller 410 from the sensor every second or five seconds, and the microcontroller may save sensor measurements for five minutes or ten minutes of BG readings. The microcontroller 410 may transfer the measurements of the physiological characteristic values to a display on the sensor electronics device 360. For example, the sensor electronics device 360 may be a monitor which includes a display that provides a blood glucose reading for a subject. In an embodiment of the invention, the microcontroller 410 may transfer the measurements of the physiological characteristic values to an output interface of the microcontroller 410. The output interface of the microcontroller 410 may transfer the measurements of the physiological characteristic values, e.g., blood glucose values, to an external device, e.g., an infusion pump, a combined infusion pump/glucose meter, a computer, a personal digital assistant, a pager, a network appliance, a server, a cellular phone, or any computing device.
  • FIG. 5 illustrates an electronic block diagram of the sensor electrodes and a voltage being applied to the sensor electrodes according to an embodiment of the present invention. In the embodiment of the invention illustrated in FIG. 5, an op amp 530 or other servo controlled device may connect to sensor electrodes 510 through a circuit/electrode interface 538. The op amp 530, utilizing feedback through the sensor electrodes, attempts to maintain a prescribed voltage (what the DAC may desire the applied voltage to be) between a reference electrode 532 and a working electrode 534 by adjusting the voltage at a counter electrode 536. Current may then flow from a counter electrode 536 to a working electrode 534. Such current may be measured to ascertain the electrochemical reaction between the sensor electrodes 510 and the biomolecule of a sensor that has been placed in the vicinity of the sensor electrodes 510 and used as a catalyzing agent. The circuitry disclosed in FIG. 5 may be utilized in a long-term or implantable sensor or may be utilized in a short-term or subcutaneous sensor.
  • In a long-term sensor embodiment, where a glucose oxidase (GOx) enzyme is used as a catalytic agent in a sensor, current may flow from the counter electrode 536 to a working electrode 534 only if there is oxygen in the vicinity of the enzyme and the sensor electrodes 510. Illustratively, if the voltage set at the reference electrode 532 is maintained at about 0.5 volts, the amount of current flowing from the counter electrode 536 to a working electrode 534 has a fairly linear relationship with unity slope to the amount of oxygen present in the area surrounding the enzyme and the electrodes. Thus, increased accuracy in determining an amount of oxygen in the blood may be achieved by maintaining the reference electrode 532 at about 0.5 volts and utilizing this region of the current-voltage curve for varying levels of blood oxygen. Different embodiments of the present invention may utilize different sensors having biomolecules other than a glucose oxidase enzyme and may, therefore, have voltages other than 0.5 volts set at the reference electrode.
  • As discussed above, during initial implantation or insertion of the sensor 510, the sensor 510 may provide inaccurate readings due to the adjusting of the subject to the sensor and also electrochemical byproducts caused by the catalyst utilized in the sensor. A stabilization period is needed for many sensors in order for the sensor 510 to provide accurate readings of the physiological parameter of the subject. During the stabilization period, the sensor 510 does not provide accurate blood glucose measurements. Users and manufacturers of the sensors may desire to improve the stabilization timeframe for the sensor so that the sensors can be utilized quickly after insertion into the subject's body or a subcutaneous layer of the subject.
  • In previous sensor electrode systems, the stabilization period or timeframe was one hour to three hours. In order to decrease the stabilization period or timeframe and increase the timeliness of accuracy of the sensor, a sensor (or electrodes of a sensor) may be subjected to a number of pulses rather than the application of one pulse followed by the application of another voltage. FIG. 6A illustrates a method of applying pulses during a stabilization timeframe in order to reduce the stabilization timeframe according to an embodiment of the present invention. In this embodiment of the invention, a voltage application device applies 600 a first voltage to an electrode for a first time or time period. In an embodiment of the invention, the first voltage may be a DC constant voltage. This results in an anodic current being generated. In an alternative embodiment of the invention, a digital-to-analog converter or another voltage source may supply the voltage to the electrode for a first time period. The anodic current means that electrons are being driven towards the electrode to which the voltage is applied. In an embodiment of the invention, an application device may apply a current instead of a voltage. In an embodiment of the invention where a voltage is applied to a sensor, after the application of the first voltage to the electrode, the voltage regulator may wait (i.e., not apply a voltage) for a second time, timeframe, or time period 605. In other words, the voltage application device waits until a second time period elapses. The non-application of voltage results in a cathodic current, which results in the gaining of electrons by the electrode to which the voltage is not applied. The application of the first voltage to the electrode for a first time period followed by the non-application of voltage for a second time period is repeated 610 for a number of iterations. This may be referred to as an anodic and cathodic cycle. In an embodiment of the invention, the number of total iterations of the stabilization method is three, i.e., three applications of the voltage for the first time period, each followed by no application of the voltage for the second time period. In an embodiment of the invention, the first voltage may be 1.07 volts. In an embodiment of the invention, the first voltage may be 0.535 volts. In an embodiment of the invention, the first voltage may be approximately 0.7 volts.
  • The repeated application of the voltage and the non-application of the voltage results in the sensor (and thus the electrodes) being subjected to an anodic-cathodic cycle. The anodic-cathodic cycle results in the reduction of electrochemical byproducts which are generated by a patient's body reacting to the insertion of the sensor or the implanting of the sensor. In an embodiment of the invention, the electrochemical byproducts cause generation of a background current, which results in inaccurate measurements of the physiological parameter of the subject. In an embodiment of the invention, the electrochemical byproduct may be eliminated. Under other operating conditions, the electrochemical byproducts may be reduced or significantly reduced. A successful stabilization method results in the anodic-cathodic cycle reaching equilibrium, electrochemical byproducts being significantly reduced, and background current being minimized.
  • In an embodiment of the invention, the first voltage being applied to the electrode of the sensor may be a positive voltage. In an embodiment of the invention, the first voltage being applied may be a negative voltage. In an embodiment of the invention, the first voltage may be applied to a working electrode. In an embodiment of the invention, the first voltage may be applied to the counter electrode or the reference electrode.
  • In embodiments of the invention, the duration of the voltage pulse and the non-application of voltage may be equal, e.g., such as three minutes each. In embodiments of the invention, the duration of the voltage application or voltage pulse may be different values, e.g., the first time and the second time may be different. In an embodiment of the invention, the first time period may be five minutes and the waiting period may be two minutes. In an embodiment of the invention, the first time period may be two minutes and the waiting period (or second timeframe) may be five minutes. In other words, the duration for the application of the first voltage may be two minutes and there may be no voltage applied for five minutes. This timeframe is only meant to be illustrative and should not be limiting. For example, a first timeframe may be two, three, five or ten minutes and the second timeframe may be five minutes, ten minutes, twenty minutes, or the like. The timeframes (e.g., the first time and the second time) may depend on unique characteristics of different electrodes, the sensors, and/or the patient's physiological characteristics.
  • In embodiments of the invention, more or less than three pulses may be utilized to stabilize the glucose sensor. In other words, the number of iterations may be greater than 3 or less than three. For example, four voltage pulses (e.g., a high voltage followed by no voltage) may be applied to one of the electrodes or six voltage pulses may be applied to one of the electrodes.
  • Illustratively, three consecutive pulses of 1.07 volts (followed by respective waiting periods) may be sufficient for a sensor implanted subcutaneously. In an embodiment of the invention, three consecutive voltage pulses of 0.7 volts may be utilized. The three consecutive pulses may have a higher or lower voltage value, either negative or positive, for a sensor implanted in blood or cranial fluid, e.g., the long-term or permanent sensors. In addition, more than three pulses (e.g., five, eight, twelve) may be utilized to create the anodic-cathodic cycling between anodic and cathodic currents in any of the subcutaneous, blood, or cranial fluid sensors.
  • FIG. 6B illustrates a method of stabilizing sensors according to an embodiment of the invention. In the embodiment of the invention illustrated in FIG. 6B, a voltage application device may apply 630 a first voltage to the sensor for a first time to initiate an anodic cycle at an electrode of the sensor. The voltage application device may be a DC power supply, a digital-to-analog converter, or a voltage regulator. After the first time period has elapsed, a second voltage is applied 635 to the sensor for a second time to initiate a cathodic cycle at an electrode of the sensor. Illustratively, rather than no voltage being applied, as is illustrated in the method of FIG. 6A, a different voltage (from the first voltage) is applied to the sensor during the second timeframe. In an embodiment of the invention, the application of the first voltage for the first time and the application of the second voltage for the second time is repeated 640 for a number of iterations. In an embodiment of the invention, the application of the first voltage for the first time and the application of the second voltage for the second time may each be applied for a stabilization timeframe, e.g., 10 minutes, 15 minutes, or 20 minutes rather than for a number of iterations. This stabilization timeframe is the entire timeframe for the stabilization sequence, e.g., until the sensor (and electrodes) are stabilized. The benefit of this stabilization methodology is a faster run-in of the sensors, less background current (in other words a suppression of some the background current), and a better glucose response.
  • In an embodiment of the invention, the first voltage may be 0.535 volts applied for five minutes, the second voltage may be 1.070 volts applied for two minutes, the first voltage of 0.535 volts may be applied for five minutes, the second voltage of 1.070 volts may be applied for two minutes, the first voltage of 0.535 volts may be applied for five minutes, and the second voltage of 1.070 volts may be applied for two minutes. In other words, in this embodiment, there are three iterations of the voltage pulsing scheme. The pulsing methodology may be changed in that the second timeframe, e.g., the timeframe of the application of the second voltage may be lengthened from two minutes to five minutes, ten minutes, fifteen minutes, or twenty minutes. In addition, after the three iterations are applied in this embodiment of the invention, a nominal working voltage of 0.535 volts may be applied.
  • The 1.070 and 0.535 volts are illustrative values. Other voltage values may be selected based on a variety of factors. These factors may include the type of enzyme utilized in the sensor, the membranes utilized in the sensor, the operating period of the sensor, the length of the pulse, and/or the magnitude of the pulse. Under certain operating conditions, the first voltage may be in a range of 1.00 to 1.09 volts and the second voltage may be in a range of 0.510 to 0.565 volts. In other operating embodiments, the ranges that bracket the first voltage and the second voltage may have a higher range, e.g., 0.3 volts, 0.6 volts, 0.9 volts, depending on the voltage sensitivity of the electrode in the sensor. Under other operating conditions, the voltage may be in a range of 0.8 volts to 1.34 volts and the other voltage may be in a range of 0.335 to 0.735. Under other operating conditions, the range of the higher voltage may be smaller than the range of the lower voltage. Illustratively, the higher voltage may be in a range of 0.9 to 1.09 volts and the lower voltage may be in a range of 0.235 to 0.835 volts.
  • In an embodiment of the invention, the first voltage and the second voltage may be positive voltages, or alternatively in other embodiments of the invention, negative voltages. In an embodiment of the invention, the first voltage may be positive and the second voltage may be negative, or alternatively, the first voltage may be negative and the second voltage may be positive. The first voltage may be different voltage levels for each of the iterations. In an embodiment of the invention, the first voltage may be a D.C. constant voltage. In other embodiments of the invention, the first voltage may be a ramp voltage, a sinusoid-shaped voltage, a stepped voltage, or other commonly utilized voltage waveforms. In an embodiment of the invention, the second voltage may be a D.C. constant voltage, a ramp voltage, a sinusoid-shaped voltage, a stepped voltage, or other commonly utilized voltage waveforms. In an embodiment of the invention, the first voltage or the second voltage may be an AC signal riding on a DC waveform. In an embodiment of the invention, the first voltage may be one type of voltage, e.g., a ramp voltage, and the second voltage may be a second type of voltage, e.g., a sinusoid-shaped voltage. In an embodiment of the invention, the first voltage (or the second voltage) may have different waveform shapes for each of the iterations. For example, if there are three cycles in a stabilization method, in a first cycle, the first voltage may be a ramp voltage, in the second cycle, the first voltage may be a constant voltage, and in the third cycle, the first voltage may be a sinusoidal voltage.
  • In an embodiment of the invention, a duration of the first timeframe and a duration of the second timeframe may have the same value, or alternatively, the duration of the first timeframe and the second timeframe may have different values. For example, the duration of the first timeframe may be two minutes and the duration of the second timeframe may be five minutes and the number of iterations may be three. As discussed above, the stabilization method may include a number of iterations. In embodiments of the invention, during different iterations of the stabilization method, the duration of each of the first timeframes may change and the duration of each of the second timeframes may change. Illustratively, during the first iteration of the anodic-cathodic cycling, the first timeframe may be 2 minutes and the second timeframe may be 5 minutes. During the second iteration, the first timeframe may be 1 minute and the second timeframe may be 3 minutes. During the third iteration, the first timeframe may be 3 minutes and the second timeframe may be 10 minutes.
  • In an embodiment of the invention, a first voltage of 0.535 volts is applied to an electrode in a sensor for two minutes to initiate an anodic cycle, then a second voltage of 1.07 volts is applied to the electrode for five minutes to initiate a cathodic cycle. The first voltage of 0.535 volts is then applied again for two minutes to initiate the anodic cycle and a second voltage of 1.07 volts is applied to the sensor for five minutes. In a third iteration, 0.535 volts is applied for two minutes to initiate the anodic cycle and then 1.07 volts is applied for five minutes. The voltage applied to the sensor is then 0.535 during the actual working timeframe of the sensor, e.g., when the sensor provides readings of a physiological characteristic of a subject.
  • Shorter duration voltage pulses may be utilized in the embodiment of FIGS. 6A and 6B. The shorter duration voltage pulses may be utilized to apply the first voltage, the second voltage, or both. In an embodiment of the present invention, the magnitude of the shorter duration voltage pulse for the first voltage is −1.07 volts and the magnitude of the shorter duration voltage pulse for the second voltage is approximately half of the high magnitude, e.g., −0.535 volts. Alternatively, the magnitude of the shorter duration pulse for the first voltage may be 0.535 volts and the magnitude of the shorter duration pulse for the second voltage is 1.07 volts.
  • In embodiments of the invention utilizing short duration pulses, the voltage may not be applied continuously for the entire first time period. Instead, the voltage application device may transmit a number of short duration pulses during the first time period. In other words, a number of mini-width or short duration voltage pulses may be applied to the electrodes of the sensor over the first time period. Each mini-width or short duration pulse may have a width of a number of milliseconds. Illustratively, this pulse width may be 30 milliseconds, 50 milliseconds, 70 milliseconds or 200 milliseconds. These values are meant to be illustrative and not limiting. In an embodiment of the invention, such as the embodiment illustrated in FIG. 6A, these short duration pulses are applied to the sensor (electrode) for the first time period and then no voltage is applied for the second time period.
  • In an embodiment of the invention, each short duration pulse may have the same time duration within the first time period. For example, each short duration voltage pulse may have a time width of 50 milliseconds and each pulse delay between the pulses may be 950 milliseconds. In this example, if two minutes is the measured time for the first timeframe, then 120 short duration voltage pulses may be applied to the sensor. In an embodiment of the invention, each of the short duration voltage pulses may have different time durations. In an embodiment of the invention, each of the short duration voltage pulses may have the same amplitude values. In an embodiment of the invention, each of the short duration voltage pulses may have different amplitude values. By utilizing short duration voltage pulses rather than a continuous application of voltage to the sensor, the same anodic and cathodic cycling may occur and the sensor (e.g., electrodes) is subjected to less total energy or charge over time. The use of short duration voltage pulses utilizes less power as compared to the application of continuous voltage to the electrodes because there is less energy applied to the sensors (and thus the electrodes).
  • FIG. 6C illustrates utilization of feedback in stabilizing the sensor according to an embodiment of the present invention. The sensor system may include a feedback mechanism to determine if additional pulses are needed to stabilize a sensor. In an embodiment of the invention, a sensor signal generated by an electrode (e.g., a working electrode) may be analyzed to determine if the sensor signal is stabilized. A first voltage is applied 630 to an electrode for a first timeframe to initiate an anodic cycle. A second voltage is applied 635 to an electrode for a second timeframe to initiate a cathodic cycle. In an embodiment of the invention, an analyzation module may analyze a sensor signal (e.g., the current emitted by the sensor signal, a resistance at a specific point in the sensor, an impedance at a specific node in the sensor) and determine if a threshold measurement has been reached 637 (e.g., determining if the sensor is providing accurate readings by comparing against the threshold measurement). If the sensor readings are determined to be accurate, which represents that the electrode (and thus the sensor) is stabilized 642, no additional application of the first voltage and/or the second voltage may be generated. If stability was not achieved, in an embodiment of the invention, then an additional anodic/cathodic cycle is initiated by the application 630 of a first voltage to an electrode for a first time period and then the application 635 of the second voltage to the electrode for a second time period.
  • In embodiments of the invention, the analyzation module may be employed after an anodic/cathodic cycle of three applications of the first voltage and the second voltage to an electrode of the sensor. In an embodiment of the invention, an analyzation module may be employed after one application of the first voltage and the second voltage, as is illustrated in FIG. 6C.
  • In an embodiment of the invention, the analyzation module may be utilized to measure a voltage emitted after a current has been introduced across an electrode or across two electrodes. The analyzation module may monitor a voltage level at the electrode or at the receiving level. In an embodiment of the invention, if the voltage level is above a certain threshold, this may mean that the sensor is stabilized. In an embodiment of the invention, if the voltage level falls below a threshold level, this may indicate that the sensor is stabilized and ready to provide readings. In an embodiment of the invention, a current may be introduced to an electrode or across a couple of electrodes. The analyzation module may monitor a current level emitted from the electrode. In this embodiment of the invention, the analyzation module may be able to monitor the current if the current is different by an order of magnitude from the sensor signal current. If the current is above or below a current threshold, this may signify that the sensor is stabilized.
  • In an embodiment of the invention, the analyzation module may measure an impedance between two electrodes of the sensor. The analyzation module may compare the impedance against a threshold or target impedance value and if the measured impedance is lower than the target or threshold impedance, the sensor (and hence the sensor signal) may be stabilized. In an embodiment of the invention, the analyzation module may measure a resistance between two electrodes of the sensor. In this embodiment of the invention, if the analyzation module compares the resistance against a threshold or target resistance value and the measured resistance value is less than the threshold or target resistance value, then the analyzation module may determine that the sensor is stabilized and that the sensor signal may be utilized.
  • FIG. 7 illustrates an effect of stabilizing a sensor according to an embodiment of the invention. Line 705 represents blood glucose sensor readings for a glucose sensor where a previous single pulse stabilization method was utilized. Line 710 represents blood glucose readings for a glucose sensor where three voltage pulses are applied (e.g., 3 voltage pulses having a duration of 2 minutes each followed by 5 minutes of no voltage being applied). The x-axis 715 represents an amount of time. The dots 720, 725, 730, and 735 represent measured glucose readings, taken utilizing a finger stick and then input into a glucose meter. As illustrated by the graph, the previous single pulse stabilization method took approximately 1 hour and 30 minutes in order to stabilize to the desired glucose reading, e.g., 100 units. In contrast, the three pulse stabilization method took only approximately 15 minutes to stabilize the glucose sensor and results in a drastically improved stabilization timeframe.
  • FIG. 8A illustrates a block diagram of a sensor electronics device and a sensor including a voltage generation device according to an embodiment of the invention. The voltage generation or application device 810 includes electronics, logic, or circuits which generate voltage pulses. The sensor electronics device 360 may also include an input device 820 to receive reference values and other useful data. In an embodiment of the invention, the sensor electronics device may include a measurement memory 830 to store sensor measurements. In this embodiment of the invention, the power supply 380 may supply power to the sensor electronics device. The power supply 380 may supply power to a regulator 385, which supplies a regulated voltage to the voltage generation or application device 810. The connection terminals 811 represent that in the illustrated embodiment of the invention, the connection terminal couples or connects the sensor 355 to the sensor electronics device 360.
  • In an embodiment of the invention illustrated in FIG. 8A, the voltage generation or application device 810 supplies a voltage, e.g., the first voltage or the second voltage, to an input terminal of an operational amplifier 840. The voltage generation or application device 810 may also supply the voltage to a working electrode 375 of the sensor 355. Another input terminal of the operational amplifier 840 is coupled to the reference electrode 370 of the sensor. The application of the voltage from the voltage generation or application device 810 to the operational amplifier 840 drives a voltage measured at the counter electrode 365 to be close to or equal to the voltage applied at the working electrode 375. In an embodiment of the invention, the voltage generation or application device 810 could be utilized to apply the desired voltage between the counter electrode and the working electrode. This may occur by the application of the fixed voltage to the counter electrode directly.
  • In an embodiment of the invention as illustrated in FIGS. 6A and 6B, the voltage generation device 810 generates a first voltage that is to be applied to the sensor during a first timeframe. The voltage generation device 810 transmits this first voltage to an op amp 840 which drives the voltage at a counter electrode 365 of the sensor 355 to the first voltage. In an embodiment of the invention, the voltage generation device 810 also could transmit the first voltage directly to the counter electrode 365 of the sensor 355. In the embodiment of the invention illustrated in FIG. 6A, the voltage generation device 810 then does not transmit the first voltage to the sensor 355 for a second timeframe. In other words, the voltage generation device 810 is turned off or switched off. The voltage generation device 810 may be programmed to continue cycling between applying the first voltage and not applying a voltage for either a number of iterations or for a stabilization timeframe, e.g., for twenty minutes. FIG. 8B illustrates a voltage generation device to implement this embodiment of the invention. The voltage regulator 385 transfers the regulated voltage to the voltage generation device 810. A control circuit 860 controls the closing and opening of a switch 850. If the switch 850 is closed, the voltage is applied. If the switch 850 is opened, the voltage is not applied. The timer 865 provides a signal to the control circuit 860 to instruct the control circuit 860 to turn on and off the switch 850. The control circuit 860 includes logic which can instruct the circuit to open and close the switch 850 a number of times (to match the necessary iterations). In an embodiment of the invention, the timer 865 may also transmit a stabilization signal to identify that the stabilization sequence is completed, i.e., that a stabilization timeframe has elapsed.
  • In an embodiment of the invention, the voltage generation device generates a first voltage for a first timeframe and generates a second voltage for a second timeframe. FIG. 8C illustrates a voltage generation device to generate two voltage values to implement this embodiment of the invention. In this embodiment of the invention, a two position switch 870 is utilized. Illustratively, if the first switch position 871 is turned on or closed by the timer 865 instructing the control circuit 860, then the voltage generation device 810 generates a first voltage for the first timeframe. After the first voltage has been applied for the first timeframe, the timer sends a signal to the control circuit 860 indicating the first timeframe has elapsed and the control circuit 860 directs the switch 870 to move to the second position 872. When the switch 870 is at the second position 872, the regulated voltage is directed to a voltage step-down or buck converter 880 to reduce the regulated voltage to a lesser value. The lesser value is then delivered to the op amp 840 for the second timeframe. After the timer 865 has sent a signal to the control circuit 860 that the second timeframe has elapsed, the control circuit 860 moves the switch 870 back to the first position. This continues until the desired number of iterations has been completed or the stabilization timeframe has elapsed. In an embodiment of the invention, after the sensor stabilization timeframe has elapsed, the sensor transmits a sensor signal 350 to the signal processor 390.
  • FIG. 8D illustrates a voltage application device 810 utilized to perform more complex applications of voltage to the sensor. The voltage application device 810 may include a control device 860, a switch 890, a sinusoid voltage generation device 891, a ramp voltage generation device 892, and a constant voltage generation device 893. In other embodiments of the invention, the voltage application may generate an AC wave on top of a DC signal or other various voltage pulse waveforms. In the embodiment of the invention illustrated in FIG. 8D, the control device 860 may cause the switch to move to one of the three voltage generation systems 891 (sinusoid), 892 (ramp), 893 (constant DC). This results in each of the voltage generation systems generating the identified voltage waveform. Under certain operating conditions, e.g., where a sinusoidal pulse is to be applied for three pulses, the control device 860 may cause the switch 890 to connect the voltage from the voltage regulator 385 to the sinusoid voltage generator 891 in order for the voltage application device 810 to generate a sinusoidal voltage. Under other operating conditions, e.g., when a ramp voltage is applied to the sensor as the first voltage for a first pulse of three pulses, a sinusoid voltage is applied to the sensor as the first voltage for a second pulse of the three pulses, and a constant DC voltage is applied to the sensor as the first voltage for a third pulse of the three pulses, the control device 860 may cause the switch 890, during the first timeframes in the anodic/cathodic cycles, to move between connecting the voltage from the voltage generation or application device 810 to the ramp voltage generation system 892, then to the sinusoidal voltage generation system 891, and then to the constant DC voltage generation system 893. In this embodiment of the invention, the control device 860 may also be directing or controlling the switch to connect certain ones of the voltage generation subsystems to the voltage from the regulator 385 during the second timeframe, e.g., during application of the second voltage.
  • FIG. 9A illustrates a sensor electronics device including a microcontroller for generating voltage pulses according to an embodiment of the invention. The advanced sensor electronics device may include a microcontroller 410 (see FIG. 4), a digital-to-analog converter (DAC) 420, an op amp 840, and a sensor signal measurement circuit 431. In an embodiment of the invention, the sensor signal measurement circuit may be a current-to-frequency (I/F) converter 430. In the embodiment of the invention illustrated in FIG. 9A, software or programmable logic in the microcontroller 410 provides instructions to transmit signals to the DAC 420, which in turn instructs the DAC 420 to output a specific voltage to the operational amplifier 840. The microcontroller 410 may also be instructed to output a specific voltage to the working electrode 375, as is illustrated by line 911 in FIG. 9A. As discussed above, the application of the specific voltage to operational amplifier 840 and the working electrode 375 may drive the voltage measured at the counter electrode to the specific voltage magnitude. In other words, the microcontroller 410 outputs a signal which is indicative of a voltage or a voltage waveform that is to be applied to the sensor 355 (e.g., the operational amplifier 840 coupled to the sensor 355). In an alternative embodiment of the invention, a fixed voltage may be set by applying a voltage directly from the DAC 420 between the reference electrode and the working electrode 375. A similar result may also be obtained by applying voltages to each of the electrodes with the difference equal to the fixed voltage applied between the reference and working electrode. In addition, the fixed voltage may be set by applying a voltage between the reference and the counter electrode. Under certain operating conditions, the microcontroller 410 may generate a pulse of a specific magnitude which the DAC 420 understands represents that a voltage of a specific magnitude is to be applied to the sensor. After a first timeframe, the microcontroller 410 (via the program or programmable logic) outputs a second signal which either instructs the DAC 420 to output no voltage (for a sensor electronics device 360 operating according to the method described in FIG. 6A) or to output a second voltage (for a sensor electronics device 360 operating according to the method described in FIG. 6B). The microcontroller 410, after the second timeframe has elapsed, then repeats the cycle of sending the signal indicative of a first voltage to be applied (for the first timeframe) and then sending the signal to instruct no voltage is to be applied or that a second voltage is to be applied (for the second timeframe).
  • Under other operating conditions, the microcontroller 410 may generate a signal to the DAC 420 which instructs the DAC to output a ramp voltage. Under other operating conditions, the microcontroller 410 may generate a signal to the DAC 420 which instructs the DAC 420 to output a voltage simulating a sinusoidal voltage. These signals could be incorporated into any of the pulsing methodologies discussed above in the preceding paragraph or earlier in the application. In an embodiment of the invention, the microcontroller 410 may generate a sequence of instructions and/or pulses, which the DAC 420 receives and understands to mean that a certain sequence of pulses is to be applied. For example, the microcontroller 410 may transmit a sequence of instructions (via signals and/or pulses) that instruct the DAC 420 to generate a constant voltage for a first iteration of a first timeframe, a ramp voltage for a first iteration of a second timeframe, a sinusoidal voltage for a second iteration of a first timeframe, and a squarewave having two values for a second iteration of the second timeframe.
  • The microcontroller 410 may include programmable logic or a program to continue this cycling for a stabilization timeframe or for a number of iterations. Illustratively, the microcontroller 410 may include counting logic to identify when the first timeframe or the second timeframe has elapsed. Additionally, the microcontroller 410 may include counting logic to identify that a stabilization timeframe has elapsed. After any of the preceding timeframes have elapsed, the counting logic may instruct the microcontroller to either send a new signal or to stop transmission of a signal to the DAC 420.
  • The use of the microcontroller 410 allows a variety of voltage magnitudes to be applied in a number of sequences for a number of time durations. In an embodiment of the invention, the microcontroller 410 may include control logic or a program to instruct the digital-to-analog converter 420 to transmit a voltage pulse having a magnitude of approximately 1.0 volt for a first time period of 1 minute, to then transmit a voltage pulse having a magnitude of approximately 0.5 volts for a second time period of 4 minutes, and to repeat this cycle for four iterations. In an embodiment of the invention, the microcontroller 420 may be programmed to transmit a signal to cause the DAC 420 to apply the same magnitude voltage pulse for each first voltage in each of the iterations. In an embodiment of the invention, the microcontroller 410 may be programmed to transmit a signal to cause the DAC to apply a different magnitude voltage pulse for each first voltage in each of the iterations. In this embodiment of the invention, the microcontroller 410 may also be programmed to transmit a signal to cause the DAC 420 to apply a different magnitude voltage pulse for each second voltage in each of the iterations. Illustratively, the microcontroller 410 may be programmed to transmit a signal to cause the DAC 420 to apply a first voltage pulse of approximately 1.0 volt in the first iteration, to apply a second voltage pulse of approximately 0.5 volts in the first iteration, to apply a first voltage of 0.7 volts and a second voltage of 0.4 volts in the second iteration, and to apply a first voltage of 1.2 volts and a second voltage of 0.8 volts in the third iteration.
  • The microcontroller 410 may also be programmed to instruct the DAC 420 to provide a number of short duration voltage pulses for a first timeframe. In this embodiment of the invention, rather than one voltage being applied for the entire first timeframe (e.g., two minutes), a number of shorter duration pulses may be applied to the sensor. In this embodiment, the microcontroller 410 may also be programmed to instruct the DAC 420 to provide a number of short duration voltage pulses for the second timeframe to the sensor. Illustratively, the microcontroller 410 may send a signal to cause the DAC to apply a number of short duration voltage pulses where the short duration is 50 milliseconds or 100 milliseconds. In between these short duration pulses the DAC may apply no voltage or the DAC may apply a minimal voltage. The microcontroller may cause the DAC 420 to apply the short duration voltage pulses for the first timeframe, e.g., two minutes. The microcontroller 410 may then send a signal to cause the DAC to either not apply any voltage or to apply the short duration voltage pulses at a magnitude of a second voltage for a second timeframe to the sensor, e.g., the second voltage may be 0.75 volts and the second timeframe may be 5 minutes. In an embodiment of the invention, the microcontroller 410 may send a signal to the DAC 420 to cause the DAC 420 to apply a different magnitude voltage for each of the short duration pulses in the first timeframe and/or in the second timeframe. In an embodiment of the invention, the microcontroller 410 may send a signal to the DAC 420 to cause the DAC 420 to apply a pattern of voltage magnitudes to the short durations voltage pulses for the first timeframe or the second timeframe. For example, the microcontroller may transmit a signal or pulses instructing the DAC 420 to apply thirty 20-millisecond pulses to the sensor during the first timeframe. Each of the thirty 20-millisecond pulses may have the same magnitude or may have a different magnitude. In this embodiment of the invention, the microcontroller 410 may instruct the DAC 420 to apply short duration pulses during the second timeframe or may instruct the DAC 420 to apply another voltage waveform during the second timeframe.
  • Although the disclosures in FIGS. 6-8 disclose the application of a voltage, a current may also be applied to the sensor to initiate the stabilization process. Illustratively, in the embodiment of the invention illustrated in FIG. 6B, a first current may be applied during a first timeframe to initiate an anodic or cathodic response and a second current may be applied during a second timeframe to initiate the opposite anodic or cathodic response. The application of the first current and the second current may continue for a number of iterations or may continue for a stabilization timeframe. In an embodiment of the invention, a first current may be applied during a first timeframe and a first voltage may be applied during a second timeframe. In other words, one of the anodic or cathodic cycles may be triggered by a current being applied to the sensor and the other of the anodic or cathodic cycles may be triggered by a voltage being applied to the sensor. As described above, a current applied may be a constant current, a ramp current, a stepped pulse current, or a sinusoidal current. Under certain operating conditions, the current may be applied as a sequence of short duration pulses during the first timeframe.
  • FIG. 9B illustrates a sensor and sensor electronics utilizing an analyzation module for feedback in a stabilization period according to an embodiment of the present invention. FIG. 9B introduces an analyzation module 950 to the sensor electronics device 360. The analyzation module 950 utilizes feedback from the sensor to determine whether or not the sensor is stabilized. In an embodiment of the invention, the microcontroller 410 may include instructions or commands to control the DAC 420 so that the DAC 420 applies a voltage or current to a part of the sensor 355. FIG. 9B illustrates that a voltage or current could be applied between a reference electrode 370 and a working electrode 375. However, the voltage or current can be applied in between electrodes or directly to one of the electrodes and the invention should not be limited by the embodiment illustrated in FIG. 9B. The application of the voltage or current is illustrated by dotted line 955. The analyzation module 950 may measure a voltage, a current, a resistance, or an impedance in the sensor 355. FIG. 9B illustrates that the measurement occurs at the working electrode 375, but this should not limit the invention because other embodiments of the invention may measure a voltage, a current, a resistance, or an impedance in between electrodes of the sensor or directly at either the reference electrode 370 or the counter electrode 365. The analyzation module 950 may receive the measured voltage, current, resistance, or impedance and may compare the measurement to a stored value (e.g., a threshold value). Dotted line 956 represents the analyzation module 950 reading or taking a measurement of the voltage, current, resistance, or impedance. Under certain operating conditions, if the measured voltage, current, resistance, or impedance is above the threshold, the sensor is stabilized and the sensor signal is providing accurate readings of a physiological condition of a patient. Under other operating conditions, if the measured voltage, current, resistance, or impedance is below the threshold, the sensor is stabilized. Under other operating conditions, the analyzation module 950 may verify that the measured voltage, current, resistance, or impedance is stable for a specific timeframe, e.g., one minute or two minutes. This may represent that the sensor 355 is stabilized and that the sensor signal is transmitting accurate measurements of a subject's physiological parameter, e.g., blood glucose level. After the analyzation module 950 has determined that the sensor is stabilized and the sensor signal is providing accurate measurements, the analyzation module 950 may transmit a signal (e.g., a sensor stabilization signal) to the microcontroller 410 indicating that the sensor is stabilized and that the microcontroller 410 can start using or receiving the sensor signal from the sensor 355. This is represented by dotted line 957.
  • FIG. 10 illustrates a block diagram of a sensor system including hydration electronics according to an embodiment of the invention. The sensor system includes a connector 1010, a sensor 1012, and a monitor or sensor electronics device 1025. The sensor 1012 includes electrodes 1020 and a connection portion 1024. In an embodiment of the invention, the sensor 1012 may be connected to the sensor electronics device 1025 via a connector 1010 and a cable. In other embodiments of the invention, the sensor 1012 may be directly connected to the sensor electronics device 1025. In other embodiments of the invention, the sensor 1012 may be incorporated into the same physical device as the sensor electronics device 1025. The monitor or sensor electronics device 1025 may include a power supply 1030, a regulator 1035, a signal processor 1040, a measurement processor 1045, and a processor 1050. The monitor or sensor electronics device 1025 may also include a hydration detection circuit 1060. The hydration detection circuit 1060 interfaces with the sensor 1012 to determine if the electrodes 1020 of the sensor 1012 are sufficiently hydrated. If the electrodes 1020 are not sufficiently hydrated, the electrodes 1020 do not provide accurate glucose readings, so it is important to know when the electrodes 1020 are sufficiently hydrated. Once the electrodes 1020 are sufficiently hydrated, accurate glucose readings may be obtained.
  • In an embodiment of the invention illustrated in FIG. 10, the hydration detection circuit 1060 may include a delay or timer module 1065 and a connection detection module 1070. In an embodiment of the invention utilizing the short term sensor or the subcutaneous sensor, after the sensor 1012 has been inserted into the subcutaneous tissue, the sensor electronics device or monitor 1025 is connected to the sensor 1012. The connection detection module 1070 identifies that the sensors electronics device 1025 has been connected to the sensor 1012 and sends a signal to the timer module 1065. This is illustrated in FIG. 10 by the arrow 1084 which represents a detector 1083 detecting a connection and sending a signal to the connection detection module 1070 indicating the sensor 1012 has been connected to the sensor electronics device 1025. In an embodiment of the invention where implantable or long-term sensors are utilized, a connection detection module 1070 identifies that the implantable sensor has been inserted into the body. The timer module 1065 receives the connection signal and waits a set or established hydration time. Illustratively, the hydration time may be two minutes, five minutes, ten minutes, or 20 minutes. These examples are meant to be illustrative and not to be limiting. The timeframe does not have to be a set number of minutes and can include any number of seconds. In an embodiment of the invention, after the timer module 1065 has waited for the set hydration time, the timer module 1065 may notify the processor 1050 that the sensor 1012 is hydrated by sending a hydration signal, which is illustrated by line 1086.
  • In this embodiment of the invention, the processor 1050 may receive the hydration signal and only start utilizing the sensor signal (e.g., sensor measurements) after the hydration signal has been received. In another embodiment of the invention, the hydration detection circuit 1060 may be coupled between the sensor (the sensor electrodes 1020) and the signal processor 1040. In this embodiment of the invention, the hydration detection circuit 1060 may prevent the sensor signal from being sent to signal processor 1040 until the timer module 1065 has notified the hydration detection circuit 1060 that the set hydration time has elapsed. This is illustrated by the dotted lines labeled with reference numerals 1080 and 1081. Illustratively, the timer module 1065 may transmit a connection signal to a switch (or transistor) to turn on the switch and let the sensor signal proceed to the signal processor 1040. In an alternative embodiment of the invention, the timer module 1065 may transmit a connection signal to turn on a switch 1088 (or close the switch 1088) in the hydration detection circuit 1060 to allow a voltage from the regulator 1035 to be applied to the sensor 1012 after the hydration time has elapsed. In other words, in this embodiment of the invention, the voltage from the regulator 1035 is not applied to the sensor 1012 until after the hydration time has elapsed.
  • FIG. 11 illustrates an embodiment of the invention including a mechanical switch to assist in determining a hydration time. In an embodiment of the invention, a single housing may include a sensor assembly 1120 and a sensor electronics device 1125. In an embodiment of the invention, the sensor assembly 1120 may be in one housing and the sensor electronics device 1125 may be in a separate housing, but the sensor assembly 1120 and the sensor electronics device 1125 may be connected together. In this embodiment of the invention, a connection detection mechanism 1160 may be a mechanical switch. The mechanical switch may detect that the sensor 1120 is physically connected to the sensor electronics device 1125. In an embodiment of the invention, a timer circuit 1135 may also be activated when the mechanical switch 1160 detects that the sensor 1120 is connected to the sensor electronics device 1125. In other words, the mechanical switch may close and a signal may be transferred to a timer circuit 1135. Once a hydration time has elapsed, the timer circuit 1135 transmits a signal to the switch 1140 to allow the regulator 1035 to apply a voltage to the sensor 1120. In other words, no voltage is applied until the hydration time has elapsed. In an embodiment of the invention, current may replace voltage as what is being applied to the sensor once the hydration time elapses. In an alternative embodiment of the invention, when the mechanical switch 1160 identifies that a sensor 1120 has been physically connected to the sensor electronics device 1125, power may initially be applied to the sensor 1120. Power being sent to the sensor 1120 results in a sensor signal being output from the working electrode in the sensor 1120. The sensor signal may be measured and sent to a processor 1175. The processor 1175 may include a counter input. Under certain operating conditions, after a set hydration time has elapsed from when the sensor signal was input into the processor 1175, the processor 1175 may start processing the sensor signal as an accurate measurement of the glucose in a subject's body. In other words, the processor 1170 has received the sensor signal from the potentiostat circuit 1170 for a certain amount of time, but will not process the signal until receiving an instruction from the counter input of the processor identifying that a hydration time has elapsed. In an embodiment of the invention, the potentiostat circuit 1170 may include a current-to-frequency converter 1180. In this embodiment of the invention, the current-to-frequency converter 1180 may receive the sensor signal as a current value and may convert the current value into a frequency value, which is easier for the processor 1175 to handle.
  • In an embodiment of the invention, the mechanical switch 1160 may also notify the processor 1175 when the sensor 1120 has been disconnected from the sensor electronics device 1125. This is represented by dotted line 1176 in FIG. 11. This may result in the processor 1170 powering down or reducing power to a number of components, chips, and/or circuits of the sensor electronics device 1125. If the sensor 1120 is not connected, the battery or power source may be drained if the components or circuits of the sensor electronics device 1125 are in a power on state. Accordingly, if the mechanical switch 1160 detects that the sensor 1120 has been disconnected from the sensor electronics device 1125, the mechanical switch may indicate this to the processor 1175, and the processor 1175 may power down or reduce power to one or more of the electronic circuits, chips, or components of the sensor electronics device 1125.
  • FIG. 12 illustrates an electrical method of detection of hydration according to an embodiment of the invention. In an embodiment of the invention, an electrical detecting mechanism for detecting connection of a sensor may be utilized. In this embodiment of the invention, the hydration detection electronics 1250 may include an AC source 1255 and a detection circuit 1260. The hydration detection electronics 1250 may be located in the sensor electronics device 1225. The sensor 1220 may include a counter electrode 1221, a reference electrode 1222, and a working electrode 1223. As illustrated in FIG. 12, the AC source 1255 is coupled to a voltage setting device 1275, the reference electrode 1222, and the detection circuit 1260. In this embodiment of the invention, an AC signal from the AC source is applied to the reference electrode connection, as illustrated by dotted line 1291 in FIG. 12. In an embodiment of the invention, the AC signal is coupled to the sensor 1220 through an impedance and the coupled signal is attenuated significantly if the sensor 1220 is connected to the sensor electronics device 1225. Thus, a low level AC signal is present at an input to the detection circuit 1260. This may also be referred to as a highly attenuated signal or a signal with a high level of attenuation. Under certain operating conditions, the voltage level of the AC signal may be Vapplied*(Ccoupling)/(Ccoupling+Csensor). If the detection circuit 1260 detects that a high level AC signal (lowly attenuated signal) is present at an input terminal of the detection circuit 1260, no interrupt is sent to the microcontroller 410 because the sensor 1220 has not been sufficiently hydrated or activated. For example, the input of the detection circuit 1260 may be a comparator. If the sensor 1220 is sufficiently hydrated (or wetted), an effective capacitance forms between the counter electrode and the reference electrode (e.g., capacitance Cr-c in FIG. 12), and an effective capacitance forms between the reference electrode and the working electrode (e.g., capacitance Cw-r in FIG. 12). In other words, an effective capacitance relates to capacitance being formed between two nodes and does not represent that an actual capacitor is placed in a circuit between the two electrodes. In an embodiment of the invention, the AC signal from the AC source 1255 is sufficiently attenuated by capacitances Cr-c and Cw-r and the detection circuit 1260 detects the presence of a low level or highly attenuated AC signal from the AC source 1255 at the input terminal of the detection circuit 1260. This embodiment of the invention is significant because the utilization of the existing connections between the sensor 1120 and the sensor electronics device 1125 reduces the number of connections to the sensor. In other words, the mechanical switch, disclosed in FIG. 11, requires a switch and associated connections between the sensor 1120 and the sensor electronics device 1125. It is advantageous to eliminate the mechanical switch because the sensor 1120 is continuously shrinking in size and the elimination of components helps achieve this size reduction. In alternative embodiments of the invention, the AC signal may be applied to different electrodes (e.g., the counter electrode or the working electrode) and the invention may operate in a similar fashion.
  • As noted above, after the detection circuit 1260 has detected that a low level AC signal is present at the input terminal of the detection circuit 1260, the detection circuit 1260 may later detect that a high level AC signal, with low attenuation, is present at the input terminal. This represents that the sensor 1220 has been disconnected from the sensor electronics device 1225 or that the sensor is not operating properly. If the sensor has been disconnected from the sensor electronics device 1225, the AC source may be coupled with little or low attenuation to the input of the detection circuit 1260. As noted above, the detection circuit 1260 may generate an interrupt to the microcontroller. This interrupt may be received by the microcontroller and the microcontroller may reduce or eliminate power to one or a number of components or circuits in the sensor electronics device 1225. This may be referred to as the second interrupt. Again, this helps reduce power consumption of the sensor electronics device 1225, specifically when the sensor 1220 is not connected to the sensor electronics device 1225.
  • In an alternative embodiment of the invention illustrated in FIG. 12, the AC signal may be applied to the reference electrode 1222, as is illustrated by reference numeral 1291, and an impedance measuring device 1277 may measure the impedance of an area in the sensor 1220. Illustratively, the area may be an area between the reference electrode and the working electrode, as illustrated by dotted line 1292 in FIG. 12. Under certain operating conditions, the impedance measuring device 1277 may transmit a signal to the detection circuit 1260 if a measured impedance has decreased to below an impedance threshold or other set criteria. This represents that the sensor is sufficiently hydrated. Under other operating conditions, the impedance measuring device 1277 may transmit a signal to the detection circuit 1260 once the impedance is above an impedance threshold. The detection circuit 1260 then transmits the interrupt to the microcontroller 410. In another embodiment of the invention, the impedance measuring device 1277 may transmit an interrupt or signal directly to the microcontroller.
  • In an alternative embodiment of the invention, the AC source 1255 may be replaced by a DC source. If a DC source is utilized, then a resistance measuring element may be utilized in place of an impedance measuring element 1277. In an embodiment of the invention utilizing the resistance measuring element, once the resistance drops below a resistance threshold or a set criteria, the resistance measuring element may transmit a signal to the detection circuit 1260 (represented by dotted line 1293) or directly to the microcontroller indicating that the sensor is sufficiently hydrated and that power may be applied to the sensor.
  • In the embodiment of the invention illustrated in FIG. 12, if the detection circuit 1260 detects a low level or highly attenuated AC signal from the AC source, an interrupt is generated to the microcontroller 410. This interrupt indicates that sensor is sufficiently hydrated. In this embodiment of the invention, in response to the interrupt, the microcontroller 410 generates a signal that is transferred to a digital-to-analog converter 420 to instruct or cause the digital-to-analog converter 420 to apply a voltage or current to the sensor 1220. Any of the different sequence of pulses or short duration pulses described above in FIG. 6A, 6B, or 6C or the associated text describing the application of pulses, may be applied to the sensor 1220. Illustratively, the voltage from the DAC 420 may be applied to an op-amp 1275, the output of which is applied to the counter electrode 1221 of the sensor 1220. This results in a sensor signal being generated by the sensor, e.g., the working electrode 1223 of the sensor. Because the sensor is sufficiently hydrated, as identified by the interrupt, the sensor signal created at the working electrode 1223 is accurately measuring glucose. The sensor signal is measured by a sensor signal measuring device 431 and the sensor signal measuring device 431 transmits the sensor signal to the microcontroller 410 where a parameter of a subject's physiological condition is measured. The generation of the interrupt represents that a sensor is sufficiently hydrated and that the sensor 1220 is now supplying accurate glucose measurements. In this embodiment of the invention, the hydration period may depend on the type and/or the manufacturer of the sensor and on the sensor's reaction to insertion or implantation in the subject. Illustratively, one sensor 1220 may have a hydration time of five minutes and one sensor 1220 may have a hydration time of one minute, two minutes, three minutes, six minutes, or 20 minutes. Again, any amount of time may be an acceptable amount of hydration time for the sensor, but smaller amounts of time are preferable.
  • If the sensor 1220 has been connected, but is not sufficiently hydrated or wetted, the effective capacitances Cr-c and Cw-r may not attenuate the AC signal from the AC source 1255. The electrodes in the sensor 1120 are dry before insertion and because the electrodes are dry, a good electrical path (or conductive path) does not exist between the two electrodes. Accordingly, a high level AC signal or lowly attenuated AC signal may still be detected by the detection circuit 1260 and no interrupt may be generated. Once the sensor has been inserted, the electrodes become immersed in the conductive body fluid. This results in a leakage path with lower DC resistance. Also, boundary layer capacitors form at the metal/fluid interface. In other words, a rather large capacitance forms between the metal/fluid interface and this large capacitance looks like two capacitors in series between the electrodes of the sensor. This may be referred to as an effective capacitance. In practice, a conductivity of an electrolyte above the electrode is being measured. In some embodiments of the invention, the glucose limiting membrane (GLM) also illustrates impedance blocking electrical efficiency. An unhydrated GLM results in high impedance, whereas a high moisture GLM results in low impedance. Low impedance is desired for accurate sensor measurements.
  • FIG. 13A illustrates a method of hydrating a sensor according to an embodiment of the present invention. In an embodiment of the invention, the sensor may be physically connected 1310 to the sensor electronics device. After the connection, in one embodiment of the invention, a timer or counter may be initiated to count 1320 a hydration time. After the hydration time has elapsed, a signal may be transmitted 1330 to a subsystem in the sensor electronics device to initiate the application of a voltage to the sensor. As discussed above, in an embodiment of the invention, a microcontroller may receive the signal and instruct the DAC to apply a voltage to the sensor or in another embodiment of the invention, a switch may receive a signal which allows a regulator to apply a voltage to the sensor. The hydration time may be five minutes, two minutes, ten minutes and may vary depending on the subject and also on the type of sensor.
  • In an alternative embodiment of the invention, after the connection of the sensor to the sensor electronics device, an AC signal (e.g., a low voltage AC signal) may be applied 1340 to the sensor, e.g., the reference electrode of the sensor. The AC signal may be applied because the connection of the sensor to the sensor electronics device allows the AC signal to be applied to the sensor. After application of the AC signal, an effective capacitance forms 1350 between the electrode in the sensor that the voltage is applied to and the other two electrodes. A detection circuit determines 1360 what level of the AC signal is present at the input of the detection circuit. If a low level AC signal (or highly attenuated AC signal) is present at the input of the detection circuit, due to the effective capacitance forming a good electrical conduit between the electrodes and the resulting attenuation of the AC signal, an interrupt is generated 1370 by the detection circuit and sent to a microcontroller.
  • The microcontroller receives the interrupt generated by the detection circuit and transmits 1380 a signal to a digital-to-analog converter instructing or causing the digital-to-analog converter to apply a voltage to an electrode of the sensor, e.g., the counter electrode. The application of the voltage to the electrode of the sensor results in the sensor creating or generating a sensor signal 1390. A sensor signal measurement device 431 measures the generated sensor signal and transmits the sensor signal to the microcontroller. The microcontroller receives 1395 the sensor signal from the sensor signal measurement device, which is coupled to the working electrode, and processes the sensor signal to extract a measurement of a physiological characteristic of the subject or patient.
  • FIG. 13B illustrates an additional method for verifying hydration of a sensor according to an embodiment of the present invention. In the embodiment of the invention illustrated in FIG. 13B, the sensor is physically connected 1310 to the sensor electronics device. In an embodiment of the invention, an AC signal is applied 1341 to an electrode, e.g., a reference electrode, in the sensor. Alternatively, in an embodiment of the invention, a DC signal is applied 1341 to an electrode in the sensor. If an AC signal is applied, an impedance measuring element measures 1351 an impedance at a point within the sensor. Alternatively, if a DC signal is applied, a resistance measuring element measures 1351 a resistance at a point within the sensor. If the resistance or impedance is lower than a resistance threshold or an impedance threshold, respectively, (or other set criteria), then the impedance (or resistance) measuring element transmits 1361 (or allows a signal to be transmitted) to the detection circuit, and the detection circuit transmits an interrupt to the microcontroller identifying that the sensor is hydrated. The reference numbers 1380, 1390, and 1395 are the same in FIGS. 13A and 13B because they represent the same action.
  • The microcontroller receives the interrupt and transmits 1380 a signal to a digital-to-analog converter to apply a voltage to the sensor. In an alternative embodiment of the invention, the digital-to-analog converter can apply a current to the sensor, as discussed above. The sensor, e.g., the working electrode, creates 1390 a sensor signal, which represents a physiological parameter of a patient. The microcontroller receives 1395 the sensor signal from a sensor signal measuring device, which measures the sensor signal at an electrode in the sensor, e.g., the working electrode. The microcontroller processes the sensor signal to extract a measurement of the physiological characteristic of the subject or patient, e.g., the blood glucose level of the patient.
  • FIGS. 14A and 14B illustrate methods of combining hydrating of a sensor with stabilizing of a sensor according to an embodiment of the present invention. In an embodiment of the invention illustrated in FIG. 14A, the sensor is connected 1405 to the sensor electronics device. The AC signal is applied 1410 to an electrode of the sensor. The detection circuit determines 1420 what level of the AC signal is present at an input of the detection circuit. If the detection circuit determines that a low level of the AC signal is present at the input (representing a high level of attenuation to the AC signal), an interrupt is sent 1430 to microcontroller. Once the interrupt is sent to the microcontroller, the microcontroller knows to begin or initiate 1440 a stabilization sequence, i.e., the application of a number of voltage pulses to an electrode of the sensors, as described above. For example, the microcontroller may cause a digital-to-analog converter to apply three voltage pulses (having a magnitude of +0.535 volts) to the sensor with each of the three voltage pulses followed by a period of three voltage pulses (having a magnitude of 1.07 volts to be applied). This may be referred to transmitting a stabilization sequence of voltages. The microcontroller may cause this by the execution of a software program in a read-only memory (ROM) or a random access memory. After the stabilization sequence has finished executing, the sensor may generate 1450 a sensor signal, which is measured and transmitted to a microcontroller.
  • In an embodiment of the invention, the detection circuit may determine 1432 that a high level AC signal has continued to be present at the input of the detection circuit (e.g., an input of a comparator), even after a hydration time threshold has elapsed. For example, the hydration time threshold may be 10 minutes. After 10 minutes has elapsed, the detection circuit may still be detecting that a high level AC signal is present. At this point in time, the detection circuit may transmit 1434 a hydration assist signal to the microcontroller. If the microcontroller receives the hydration assist signal, the microcontroller may transmit 1436 a signal to cause a DAC to apply a voltage pulse or a series of voltage pulses to assist the sensor in hydration. In an embodiment of the invention, the microcontroller may transmit a signal to cause the DAC to apply a portion of the stabilization sequence or other voltage pulses to assist in hydrating the sensor. In this embodiment of the invention, the application of voltage pulses may result in the low level AC signal (or highly attenuated signal) being detected 1438 at the detection circuit. At this point, the detection circuit may transmit an interrupt, as is disclosed in step 1430, and the microcontroller may initiate a stabilization sequence.
  • FIG. 14B illustrates a second embodiment of a combination of a hydration method and a stabilization method where feedback is utilized in the stabilization process. A sensor is connected 1405 to a sensor electronics device. An AC signal (or a DC signal) is applied 1411 to the sensor. In an embodiment of the invention, the AC signal (or the DC signal) is applied to an electrode of the sensor, e.g. the reference electrode. An impedance measuring device (or resistance measuring device) measures 1416 the impedance (or resistance) within a specified area of the sensor. In an embodiment of the invention, the impedance (or resistance) may be measured between the reference electrode and the working electrode. The measured impedance (or resistance) may be compared 1421 to an impedance or resistance value to see if the impedance (or resistance) is low enough in the sensor, which indicates the sensor is hydrated. If the impedance (or resistance) is below the impedance (or resistance) value or other set criteria, (which may be a threshold value), an interrupt is transmitted 1431 to the microcontroller. After receiving the interrupt, the microcontroller transmits 1440 a signal to the DAC instructing the DAC to apply a stabilization sequence of voltages (or currents) to the sensor. After the stabilization sequence has been applied to the sensor, a sensor signal is created in the sensor (e.g., at the working electrode), is measured by a sensor signal measuring device, is transmitted by the sensor signal measuring device, and is received 1450 by the microcontroller. Because the sensor is hydrated and the stabilization sequence of voltages has been applied to the sensor, the sensor signal is accurately measuring a physiological parameter (i.e., blood glucose).
  • FIG. 14C illustrates a third embodiment of the invention where a stabilization method and hydration method are combined. In this embodiment of the invention, the sensor is connected 1500 to the sensor electronics device. After the sensor is physically connected to the sensor electronics device, an AC signal (or DC signal) is applied 1510 to an electrode (e.g., reference electrode) of the sensor. At the same time, or around the same time, the microcontroller transmits a signal to cause the DAC to apply 1520 a stabilization voltage sequence to the sensor. In an alternative embodiment of the invention, a stabilization current sequence may be applied to the sensor instead of a stabilization voltage sequence. The detection circuit determines 1530 what level of an AC signal (or DC signal) is present at an input terminal of the detection circuit. If there is a low level AC signal (or DC signal), representing a highly attenuated AC signal (or DC signal), present at the input terminal of the detection circuit, an interrupt is transmitted 1540 to the microcontroller. Because the microcontroller has already initiated the stabilization sequence, the microcontroller receives the interrupt and sets 1550 a first indicator that the sensor is sufficiently hydrated. After the stabilization sequence is complete, the microcontroller sets 1555 a second indicator indicating the completion of the stabilization sequence. The application of the stabilization sequence voltages results in the sensor, e.g., the working electrode, creating 1560 a sensor signal, which is measured by a sensor signal measuring circuit, and sent to the microcontroller. If the second indicator that the stabilization sequence is complete is set and the first indicator that the hydration is complete is set, the microcontroller is able to utilize 1570 the sensor signal. If one or both of the indicators are not set, the microcontroller may not utilize the sensor signal because the sensor signal may not represent accurate measurements of the physiological measurements of the subject.
  • The above-described hydration and stabilization processes may be used, in general, as part of a larger continuous glucose monitoring (CGM) methodology. The current state of the art in continuous glucose monitoring is largely adjunctive, meaning that the readings provided by a CGM device (including, e.g., an implantable or subcutaneous sensor) cannot be used without a reference value in order to make a clinical decision. The reference value, in turn, must be obtained from a finger stick using, e.g., a BG meter. The reference value is needed because there is a limited amount of information that is available from the sensor/sensing component. Specifically, the only pieces of information that are currently provided by the sensing component for processing are the raw sensor value (i.e., the sensor current or Isig) and the counter voltage, which is the voltage between the counter electrode and the reference electrode (see, e.g., FIG. 5). Therefore, during analysis, if it appears that the raw sensor signal is abnormal (e.g., if the signal is decreasing), the only way one can distinguish between a sensor failure and a physiological change within the user/patient (i.e., glucose level changing in the body) is by acquiring a reference glucose value via a finger stick. As is known, the reference finger stick is also used for calibrating the sensor.
  • Embodiments of the inventions described herein are directed to advancements and improvements in continuous glucose monitoring resulting in a more autonomous system, as well as related devices and methodologies, wherein the requirement of reference finger sticks may be minimized, or eliminated, and whereby clinical decisions may be made based on information derived from the sensor signal alone, with a high level of reliability. From a sensor-design standpoint, in accordance with embodiments of the invention, such autonomy may be achieved through electrode redundancy, sensor diagnostics, and Isig and/or sensor glucose (SG) fusion.
  • As will be explored further hereinbelow, redundancy may be achieved through the use of multiple working electrodes (e.g., in addition to a counter electrode and a reference electrode) to produce multiple signals indicative of the patient's blood glucose (BG) level. The multiple signals, in turn, may be used to assess the relative health of the (working) electrodes, the overall reliability of the sensor, and the frequency of the need, if at all, for calibration reference values.
  • Sensor diagnostics includes the use of additional (diagnostic) information which can provide a real-time insight into the health of the sensor. In this regard, it has been discovered that Electrochemical Impedance Spectroscopy (EIS) provides such additional information in the form of sensor impedance and impedance-related parameters at different frequencies. Moreover, advantageously, it has been further discovered that, for certain ranges of frequencies, impedance and/or impedance-related data are substantially glucose independent. Such glucose independence enables the use of a variety of EIS-based markers or indicators for not only producing a robust, highly-reliable sensor glucose value (through fusion methodologies), but also assessing the condition, health, age, and efficiency of individual electrode(s) and of the overall sensor substantially independently of the glucose-dependent Isig.
  • For example, analysis of the glucose-independent impedance data provides information on the efficiency of the sensor with respect to how quickly it hydrates and is ready for data acquisition using, e.g., values for 1 kHz real-impedance, 1 kHz imaginary impedance, and Nyquist Slope (to be described in more detail hereinbelow). Moreover, glucose-independent impedance data provides information on potential occlusion(s) that may exist on the sensor membrane surface, which occlusion(s) may temporarily block passage of glucose into the sensor and thus cause the signal to dip (using, e.g., values for 1 kHz real impedance). In addition, glucose-independent impedance data provides information on loss of sensor sensitivity during extended wear—potentially due to local oxygen deficit at the insertion site—using, e.g., values for phase angle and/or imaginary impedance at 1 kHz and higher frequencies.
  • Within the context of electrode redundancy and EIS, as well as other contexts, as will be described in further detail hereinbelow, a fusion algorithm may be used to take the diagnostic information provided by EIS for each redundant electrode and assess the reliability of each electrode independently. Weights, which are a measure of reliability, may then be added for each independent signal, and a single fused signal may be calculated that can be used to generate sensor glucose values as seen by the patient/subject.
  • As can be seen from the above, the combined use of redundancy, sensor diagnostics using EIS, and EIS-based fusion algorithms allows for an overall CGM system that is more reliable than what is currently available. Redundancy is advantageous in at least two respects. First, redundancy removes the risk of a single point of failure by providing multiple signals. Second, providing multiple (working) electrodes where a single electrode may be sufficient allows the output of the redundant electrode to be used as a check against the primary electrode, thereby reducing, and perhaps eliminating, the need for frequent calibrations. In addition, EIS diagnostics scrutinize the health of each electrode autonomously without the need for a reference glucose value (finger stick), thereby reducing the number of reference values required. However, the use of EIS technology and EIS diagnostic methods is not limited to redundant systems, i.e., those having more than one working electrode. Rather, as is discussed below in connection with embodiments of the invention, EIS may be advantageously used in connection with single- and/or multiple-electrode sensors.
  • EIS, or AC impedance methods, study the system response to the application of a periodic small amplitude AC signal. This is shown illustratively in FIG. 15A, where E is the applied potential, I is the current, and impedance (Z) is defined as AE/AI. However, although impedance, per se, may be mathematically simply defined as AE/AI, heretofore, there has been no commercialization success in application of EIS technology to continuous glucose monitoring. This has been due, in part, to the fact that glucose sensors are very complicated systems and, so far, no mathematical models have been developed which can completely explain the complexity of the EIS output for a glucose sensor.
  • One simplified electrical circuit model that has been used to describe electrochemical impedance spectroscopy is shown in FIG. 15B. In this illustration, IHP stands for Inner Helmholtz Plane, OHP stands for Outer Helmholtz Plane, CE is the counter electrode, WE is the working electrode, Cd is double layer capacitance, Rp is polarization resistance, Zw is Warburg impedance, and Rs is solution resistance. Each of the latter four components—double layer capacitance (Cd), Warburg impedance (Zw), polarization resistance (Rp), and solution resistance (Rs)—may play a significant role in sensor performance, and can be measured separately by applying low- or high-frequency alternating working potential. For example, Warburg impedance is closely related to diffusional impedance of electrochemical systems—which is primarily a low-frequency impedance—and, as such, exists in all diffusion-limited electrochemical sensors. Thus, by correlating one or more of these components with one or more components and/or layers of a glucose sensor, one may use EIS technology as a sensor-diagnostics tool.
  • As is known, impedance may be defined in terms of its magnitude and phase, where the magnitude (|Z|) is the ratio of the voltage difference amplitude to the current amplitude, and the phase (θ) is the phase shift by which the current is ahead of the voltage. When a circuit is driven solely with direct current (DC), the impedance is the same as the resistant, i.e., resistance is a special case of impedance with zero phase angle. However, as a complex quantity, impedance may also be represented by its real and imaginary parts. In this regard, the real and imaginary impedance can be derived from the impedance magnitude and phase using the following equations:

  • Real Impedance(ω)=Magnitude(ω)×cos(Phase(ω)/180×π)

  • Imaginary Impedance(ω)=Magnitude(ω)×sin(Phase(ω)/180×π)
  • where ω represents the input frequency at which the magnitude (in ohms) and the phase (in degrees) are measured. The relationship between impedance, on the one hand, and current and voltage on the other—including how the former may be calculated based on measurement of the latter—will be explored more fully below in connection with the sensor electronics, including the Application Specific Integrated Circuit (ASIC), that has been developed for use in embodiments of the invention.
  • Continuing with the circuit model shown in FIG. 15B, total system impedance may be simplified as:
  • Z t ( ω ) = Z w ( ω ) + R s + R p 1 + ω 2 R p 2 C d 2 - j ω R p 2 C d 1 + ω 2 R p 2 C d 2
  • where Zw(ω) is the Warburg impedance, ω is the angular velocity, j is the imaginary unit (used instead of the traditional “i” so as not to be confused with electric current), and Cd, Rp, and Rs are the double layer capacitance, the polarization resistance, and the solution resistance, respectively (as defined previously). Warburg impedance can be calculated as
  • Z w ( ω ) = Z 0 tanh ( ( js ) m ) ( js ) m s = L 2 ω / D = ( Membrane Thickness Frequency Dependent Diffusion Length ) 2 Z 0 = RTL n 2 F 2 D C
  • where D is diffusivity, L is the sensor membrane thickness, C is Peroxide concentration, and m: ½ corresponds to a 45° Nyquist slope.
  • A Nyquist plot is a graphical representation, wherein the real part of impedance (Real Z) is plotted against its imaginary part (Img Z) across a spectrum of frequencies. FIG. 16A shows a generalized example of a Nyquist Plot, where the X value is the real part of the impedance and the Y value is the imaginary part of the impedance. The phase angle is the angle between the impedance point (X,Y)—which defines a vector having magnitude |Z|—and the X axis.
  • The Nyquist plot of FIG. 16A is generated by applying AC voltages plus a DC voltage (DC bias) between the working electrode and the counter electrode at selected frequencies from 0.1 Hz to 1000 MHz (i.e., a frequency sweep). Starting from the right, the frequency increases from 0.1 Hz. With each frequency, the real and imaginary impedance can be calculated and plotted. As shown, a typical Nyquist plot of an electrochemical system may look like a semicircle joined with a straight line at an inflection point, wherein the semicircle and the line indicate the plotted impedance. In certain embodiments, the impedance at the inflection point is of particular interest since it is easiest to identify in the Nyquist plot and may define an intercept. Typically, the inflection point is close to the X axis, and the X value of the inflection point approximates the sum of the polarization resistance and solution resistance (Rp+Rb).
  • With reference to FIG. 16B, a Nyquist plot may typically be described in terms of a lower-frequency region 1610 and a higher-frequency region 1620, where the labels “higher frequency” and “lower frequency” are used in a relative sense, and are not meant to be limiting. Thus, for example, the lower-frequency region 1610 may illustratively include data points obtained for a frequency range between about 0.1 Hz and about 100 Hz (or higher), and the higher-frequency region 1620 may illustratively include data points obtained for a frequency range between about 1 kHz (or lower) and about 8 kHz (and higher). In the lower-frequency region 1610, the Nyquist slope represents the gradient of the linear fit 1630 of the lower-frequency data points in the Nyquist plot. As shown, in the higher-frequencies region 1620, the value of imaginary impedance is minimal, and may become negligible. As such, the intercept 1600 is essentially the value of the real impedance at the higher frequencies (e.g., approximately in the 1 kHz to 8 kHz range in this case). In FIG. 16B, the intercept 1600 is at about 25 kOhms.
  • FIGS. 16C and 16D demonstrate how a glucose sensor responds to a sinusoidal (i.e., alternating) working potential. In these figures, GLM is the sensor's glucose limiting membrane, AP is the adhesion promoter, HSA is human serum albumin, GOX is glucose oxidase enzyme (layer), Ede is DC potential, Eac is AC potential, and Cperoxide′ is peroxide concentration during AC application. As shown in FIG. 16C, if the sensor diffusion length, which is a function of AC potential frequency, molecular diffusivity, and membrane thickness, is small compared to the membrane (GOX) length, the system gives a relatively linear response with a constant phase angle (i.e., infinite). In contrast, if the diffusion length is equal to the membrane (GOX) length, the system response will become finite, resulting in a semi-circle Nyquist plot, as shown in FIG. 16D. The latter usually holds true for low-frequency EIS, where the non-Faradaic process is negligible.
  • In performing an EIS analysis, an AC voltage of various frequencies and a DC bias may be applied between, e.g., the working and reference electrodes. In this regard, EIS is an improvement over previous methodologies that may have limited the application to a simple DC current or an AC voltage of single frequency. Although, generally, EIS may be performed at frequencies in the μHz to MHz range, in embodiments of the invention, a narrower range of frequencies (e.g., between about 0.1 Hz and about 8 kHz) may be sufficient. Thus, in embodiments of the invention, AC potentials may be applied that fall within a frequency range of between about 0.1 Hz and about 8 kHz, with a programmable amplitude of up to at least 100 mV, and preferably at about 50 mV.
  • Within the above-mentioned frequency range, the relatively-higher frequencies—i.e., those that fall generally between about 1 kHz and about 8 kHz—are used to scrutinize the capacitive nature of the sensor. Depending on the thickness and permeability of membranes, a typical range of impedance at the relatively-higher frequencies may be, e.g., between about 500 Ohms and 25 kOhms, and a typical range for the phase may be, e.g., between 0 degrees and −40 degrees. The relatively-lower frequencies—i.e., those that fall generally between about 0.1 Hz and about 100 Hz—on the other hand, are used to scrutinize the resistive nature of the sensor. Here, depending on electrode design and the extent of metallization, a typical functioning range for output real impedance may be, e.g., between about 50 kOhms and 300 kOhms, and a typical range for the phase may be between about −50 degrees to about −90 degrees. The above illustrative ranges are shown, e.g., in the Bode plots of FIGS. 16E and 16F.
  • As noted previously, the phrases “higher frequencies” and “lower frequencies” are meant to be used relative to one another, rather than in an absolute sense, and they, as well as the typical impedance and phase ranges mentioned above, are meant to be illustrative, and not limiting. Nevertheless, the underlying principle remains the same: the capacitive and resistive behavior of a sensor can be scrutinized by analyzing the impedance data across a frequency spectrum, wherein, typically, the lower frequencies provide information about the more resistive components (e.g., the electrode, etc.), while the higher frequencies provide information about the capacitive components (e.g., membranes). However, the actual frequency range in each case is dependent on the overall design, including, e.g., the type(s) of electrode(s), the surface area of the electrode(s), membrane thickness, the permeability of the membrane, and the like. See also FIG. 15B regarding general correspondence between high-frequency circuit components and the sensor membrane, as well as between low-frequency circuit components and the Faradaic process, including, e.g., the electrode(s).
  • EIS may be used in sensor systems where the sensor includes a single working electrode, as well those in which the sensor includes multiple (redundant) working electrodes. In one embodiment, EIS provides valuable information regarding the age (or aging) of the sensor. Specifically, at different frequencies, the magnitude and the phase angle of the impedance vary. As seen in FIG. 17, the sensor impedance—in particular, the sum of Rp and Rs—reflects the sensor age as well as the sensor's operating conditions. Thus, a new sensor normally has higher impedance than a used sensor as seen from the different plots in FIG. 17. In this way, by considering the X-value of the sum of Rp and Rs, a threshold can be used to determine when the sensor's age has exceeded the specified operating life of the sensor. It is noted that, although for the illustrative examples shown in FIGS. 17-21 and discussed below, the value of real impedance at the inflection point (i.e., Rp+Rs) is used to determine the aging, status, stabilization, and hydration of the sensor, alternative embodiments may use other EIS-based parameters, such as, e.g., imaginary impedance, phase angle, Nyquist slope, etc. in addition to, or in place of, real impedance.
  • FIG. 17 illustrates an example of a Nyquist plot over the life time of a sensor. The points indicated by arrows are the respective inflection points for each of the sweeps across the frequency spectrum. For example, before initialization (at time t=0), Rs+Rp is higher than 8.5 kOhms, and after initialization (at time t=0.5 hr), the value of Rs+Rp dropped to below 8 kOhms. Over the next six days, Rs+Rp continues to decrease, such that, at the end of the specified sensor life, Rs+Rp dropped to below 6.5 kOhms. Based on such examples, a threshold value can be set to specify when the Rs+Rp value would indicate the end of the specified operating life of the sensor. Therefore, the EIS technique allows closure of the loophole of allowing a sensor to be re-used beyond the specified operating time. In other words, if the patient attempts to re-use a sensor after the sensor has reached its specified operating time by disconnecting and then re-connecting the sensor again, the EIS will measure abnormally-low impedance, thereby enabling the system to reject the sensor and prompt the patient for a new sensor.
  • Additionally, EIS may enable detection of sensor failure by detecting when the sensor's impedance drops below a low impedance threshold level indicating that the sensor may be too worn to operate normally. The system may then terminate the sensor before the specified operating life. As will be explored in more detail below, sensor impedance can also be used to detect other sensor failure (modes). For example, when a sensor goes into a low-current state (i.e., sensor failure) due to any variety of reasons, the sensor impedance may also increase beyond a certain high impedance threshold. If the impedance becomes abnormally high during sensor operation, due, e.g., to protein or polypeptide fouling, macrophage attachment or any other factor, the system may also terminate the sensor before the specified sensor operating life.
  • FIG. 18 illustrates how the EIS technique can be applied during sensor stabilization and in detecting the age of the sensor in accordance with embodiments of the invention. The logic of FIG. 18 begins at 1800 after the hydration procedure and sensor initialization procedure described previously has been completed. In other words, the sensor has been deemed to be sufficiently hydrated, and the first initialization procedure has been applied to initialize the sensor. The initialization procedure may preferably be in the form of voltage pulses as described previously in the detailed description. However, in alternative embodiments, different waveforms can be used for the initialization procedure. For example, a sine wave can be used, instead of the pulses, to accelerate the wetting or conditioning of the sensor. In addition, it may be necessary for some portion of the waveform to be greater than the normal operating voltage of the sensor, i.e., 0.535 volt.
  • At block 1810, an EIS procedure is applied and the impedance is compared to both a first high and a first low threshold. An example of a first high and first low threshold value would be 7 kOhms and 8.5 kOhms, respectively, although the values can be set higher or lower as needed. If the impedance, for example, Rp+Rs, is higher than the first high threshold, the sensor undergoes an additional initialization procedure (e.g., the application of one or more additional pulses) at block 1820. Ideally, the number of total initialization procedures applied to initialize the sensor would be optimized to limit the impact on both the battery life of the sensor and the overall amount of time needed to stabilize a sensor. Thus, by applying EIS, fewer initializations can be initially performed, and the number of initializations can be incrementally added to give just the right amount of initializations to ready the sensor for use. Similarly, in an alternative embodiment, EIS can be applied to the hydration procedure to minimize the number of initializations needed to aid the hydration process as described in FIGS. 13-14.
  • On the other hand, if the impedance, for example, Rp+Rs, is below the first low threshold, the sensor will be determined to be faulty and would be terminated immediately at block 1860. A message will be given to the user to replace the sensor and to begin the hydration process again. If the impedance is within the high and low thresholds, the sensor will begin to operate normally at block 1830. The logic then proceeds to block 1840 where an additional EIS is performed to check the age of the sensor. The first time the logic reaches block 1840, the microcontroller will perform an EIS to gauge the age of the sensor to close the loophole of the user being able to plug in and plug out the same sensor. In future iterations of the EIS procedure as the logic returns to block 1840, the microprocessor will perform an EIS at fixed intervals during the specified life of the sensor. In one preferred embodiment, the fixed interval is set for every 2 hours, however, longer or shorter periods of time can easily be used.
  • At block 1850, the impedance is compared to a second set of high and low thresholds. An example of such second high and low threshold values may be 5.5 kOhms and 8.5 kOhms, respectively, although the values can be set higher or lower as needed. As long as the impedance values stay within a second high and low threshold, the logic proceeds to block 1830, where the sensor operates normally until the specified sensor life, for example, 5 days, is reached. Of course, as described with respect to block 1840, EIS will be performed at the regularly scheduled intervals throughout the specified sensor life. However, if, after the EIS is performed, the impedance is determined to have dropped below a second lower threshold or risen above a second higher threshold at block 1850, the sensor is terminated at block 1860. In further alternative embodiments, a secondary check can be implemented of a faulty sensor reading. For example, if the EIS indicates that the impedance is out of the range of the second high and low thresholds, the logic can perform a second EIS to confirm that the second set of thresholds is indeed not met (and confirm that the first EIS was correctly performed) before determining the end of sensor at block 1860.
  • FIG. 19 builds upon the above description and details a possible schedule for performing diagnostic EIS procedures in accordance with preferred embodiments of the present invention. Each diagnostic EIS procedure is optional and it is possible to not schedule any diagnostic EIS procedure or to have any combination of one or more diagnostic EIS procedures, as deemed needed. The schedule of FIG. 19 begins at sensor insertion at point 1900. Following sensor insertion, the sensor undergoes a hydration period 1910. This hydration period is important because a sensor that is not sufficiently hydrated may give the user inaccurate readings, as described previously. The first optional diagnostic EIS procedure at point 1920 is scheduled during this hydration period 1910 to ensure that the sensor is sufficiently hydrated. The first diagnostic EIS procedure 1920 measures the sensor impedance value to determine if the sensor has been sufficiently hydrated. If the first diagnostic EIS procedure 1920 determines impedance is within a set high and low threshold, indicating sufficient hydration, the sensor controller will allow the sensor power-up at point 1930. Conversely, if the first diagnostic EIS procedure 1920 determines impedance is outside a set high and low threshold, indicating insufficient hydration, the sensor hydration period 1910 may be prolonged. After prolonged hydration, once a certain capacitance has been reached between the sensor's electrodes, meaning the sensor is sufficiently hydrated, power-up at point 1930 can occur.
  • A second optional diagnostic EIS procedure 1940 is scheduled after sensor power-up at point 1930, but before sensor initialization starts at point 1950. Scheduled here, the second diagnostic EIS procedure 1940 can detect if a sensor is being re-used prior to the start of initialization at 1950. The test to determine if the sensor is being reused was detailed in the description of FIG. 18. However, unlike the previous description with respect to FIG. 18, where the aging test is performed after initialization is completed, the aging test is shown in FIG. 19 as being performed before initialization. It is important to appreciate that the timeline of EIS procedures described in FIG. 19 can be rearranged without affecting the overall teaching of the application, and that the order of some of the steps can be interchanged. As explained previously, the second diagnostic EIS procedure 1940 detects a re-used sensor by determining the sensor's impedance value and then comparing it to a set high and low threshold. If impedance falls outside of the set threshold, indicating the sensor is being re-used, the sensor may then be rejected and the user prompted to replace it with a new sensor. This prevents the complications that may arise out of re-use of an old sensor. Conversely, if impedance falls within a set threshold, sensor initialization 1950 can start with the confidence that a new sensor is being used.
  • A third optional diagnostic EIS procedure 1960 is scheduled after initialization starts at point 1950. The third diagnostic EIS procedure 1960 tests the sensor's impedance value to determine if the sensor is fully initialized. The third diagnostic EIS procedure 1960 should be performed at the minimum amount of time needed for any sensor to be fully initialized. When performed at this time, sensor life is maximized by limiting the time a fully initialized sensor goes unused, and over-initialization is averted by confirming full initialization of the sensor before too much initialization occurs. Preventing over-initialization is important because over-initialization results in a suppressed current which can cause inaccurate readings. However, under-initialization is also a problem, so if the third diagnostic EIS procedure 1960 indicates the sensor is under-initialized, an optional initialization at point 1970 may be performed in order to fully initialize the sensor. Under-initialization is disadvantageous because an excessive current results that does not relate to the actual glucose concentration. Because of the danger of under- and over-initialization, the third diagnostic EIS procedure plays an important role in ensuring the sensor functions properly when used.
  • In addition, optional periodic diagnostic EIS procedures 1980 can be scheduled for the time after the sensor is fully initialized. The EIS procedures 1980 can be scheduled at any set interval. As will be discussed in more detail below, EIS procedures 1980 may also be triggered by other sensor signals, such as an abnormal current or an abnormal counter electrode voltage. Additionally, as few or as many EIS procedures 1980 can be scheduled as desired. In preferred embodiments, the EIS procedure used during the hydration process, sensor life check, initialization process, or the periodic diagnostic tests is the same procedure. In alternative embodiments, the EIS procedure can be shortened or lengthened (i.e., fewer or more ranges of frequencies checked) for the various EIS procedures depending on the need to focus on specific impedance ranges. The periodic diagnostic EIS procedures 1980 monitor impedance values to ensure that the sensor is continuing to operate at an optimal level.
  • The sensor may not be operating at an optimal level if the sensor current has dropped due to polluting species, sensor age, or a combination of polluting species and sensor age. A sensor that has aged beyond a certain length is no longer useful, but a sensor that has been hampered by polluting species can possibly be repaired. Polluting species can reduce the surface area of the electrode or the diffusion pathways of analytes and reaction byproducts, thereby causing the sensor current to drop. These polluting species are charged and gradually gather on the electrode or membrane surface under a certain voltage. Previously, polluting species would destroy the usefulness of a sensor. Now, if periodic diagnostic EIS procedures 1980 detect impedance values which indicate the presence of polluting species, remedial action can be taken. When remedial action is to be taken is described with respect to FIG. 20. Periodic diagnostic EIS procedures 1980 therefore become extremely useful because they can trigger sensor remedial action which can possibly restore the sensor current to a normal level and prolong the life of the sensor. Two possible embodiments of sensor remedial actions are described below in the descriptions of FIGS. 21A and 21B.
  • Additionally, any scheduled diagnostic EIS procedure 1980 may be suspended or rescheduled when certain events are determined imminent. Such events may include any circumstance requiring the patient to check the sensor reading, including for example when a patient measures his or her BG level using a test strip meter in order to calibrate the sensor, when a patient is alerted to a calibration error and the need to measure his or her BG level using a test strip meter a second time, or when a hyperglycemic or hypoglycemic alert has been issued but not acknowledged.
  • FIG. 20 illustrates a method of combining diagnostic EIS procedures with sensor remedial action in accordance with embodiments of the present invention. The block 2000 diagnostic procedure may be any of the periodic diagnostic EIS procedures 1980 as detailed in FIG. 19. The logic of this method begins when a diagnostic EIS procedure is performed at block 2000 in order to detect the sensor's impedance value. As noted, in specific embodiments, the EIS procedure applies a combination of a DC bias and an AC voltage of varying frequencies wherein the impedance detected by performing the EIS procedure is mapped on a Nyquist plot, and an inflection point in the Nyquist plot approximates a sum of polarization resistance and solution resistance (i.e., the real impedance value). After the block 2000 diagnostic EIS procedure detects the sensor's impedance value, the logic moves to block 2010.
  • At block 2010, the impedance value is compared to a set high and low threshold to determine if it is normal. If impedance is within the set boundaries of the high and low thresholds at block 2010, normal sensor operation is resumed at block 2020 and the logic of FIG. 20 will end until a time when another diagnostic EIS procedure is scheduled. Conversely, if impedance is determined to be abnormal (i.e., outside the set boundaries of the high and low thresholds) at block 2010, remedial action at block 2030 is triggered. An example of a high and low threshold value that would be acceptable during a sensor life would be 5.5 kOhms and 8.5 kOhms, respectively, although the values can be set higher or lower as needed.
  • The block 2030 remedial action is performed to remove any of the polluting species, which may have caused the abnormal impedance value. In preferred embodiments, the remedial action is performed by applying a reverse current, or a reverse voltage between the working electrode and the reference electrode. The specifics of the remedial action will be described in more detail with respect to FIG. 21. After the remedial action is performed at block 2030, impedance value is again tested by a diagnostic EIS procedure at block 2040. The success of the remedial action is then determined at block 2050 when the impedance value from the block 2040 diagnostic EIS procedure is compared to the set high or low threshold. As in block 2010, if impedance is within the set thresholds, it is deemed normal, and if impedance is outside the set thresholds, it is deemed abnormal.
  • If the sensor's impedance value is determined to have been restored to normal at block 2050, normal sensor operation at block 2020 will occur. If impedance is still not normal, indicating that either sensor age is the cause of the abnormal impedance or the remedial action was unsuccessful in removing the polluting species, the sensor is then terminated at block 2060. In alternative embodiments, instead of immediately terminating the sensor, the sensor may generate a sensor message initially requesting the user to wait and then perform further remedial action after a set period of time has elapsed. This alternative step may be coupled with a separate logic to determine if the impedance values are getting closer to being within the boundary of the high and low threshold after the initial remedial action is performed. For example, if no change is found in the sensor impedance values, the sensor may then decide to terminate. However, if the sensor impedance values are getting closer to the preset boundary, yet still outside the boundary after the initial remedial action, an additional remedial action could be performed. In yet another alternative embodiment, the sensor may generate a message requesting the user to calibrate the sensor by taking a finger stick meter measurement to further confirm whether the sensor is truly failing. All of the above embodiments work to prevent a user from using a faulty sensor that produces inaccurate readings.
  • FIG. 21A illustrates one embodiment of the sensor remedial action previously mentioned. In this embodiment, blockage created by polluting species is removed by reversing the voltage being applied to the sensor between the working electrode and the reference electrode. The reversed DC voltage lifts the charged, polluting species from the electrode or membrane surface, clearing diffusion pathways. With cleared pathways, the sensor's current returns to a normal level and the sensor can give accurate readings. Thus, the remedial action saves the user the time and money associated with replacing an otherwise effective sensor.
  • FIG. 21B illustrates an alternative embodiment of the sensor remedial action previously mentioned. In this embodiment, the reversed DC voltage applied between the working electrode and the reference electrode is coupled with an AC voltage. By adding the AC voltage, certain tightly absorbed species or species on the superficial layer can be removed since the AC voltage can extend its force further from the electrode and penetrate all layers of the sensor. The AC voltage can come in any number of different waveforms. Some examples of waveforms that could be used include square waves, triangular waves, sine waves, or pulses. As with the previous embodiment, once polluting species are cleared, the sensor can return to normal operation, and both sensor life and accuracy are improved.
  • While the above examples illustrate the use, primarily, of real impedance data in sensor diagnostics, embodiments of the invention are also directed to the use of other EIS-based, and substantially analyte-independent, parameters (in addition to real impedance) in sensor diagnostic procedures. For example, as mentioned previously, analysis of (substantially) glucose-independent impedance data, such as, e.g., values for 1 kHz real impedance and 1 kHz imaginary impedance, as well as Nyquist slope, provide information on the efficiency of the sensor with respect to how quickly it hydrates and is ready for data acquisition. Moreover, (substantially) glucose-independent impedance data, such as, e.g., values for 1 kHz real impedance, provides information on potential occlusion(s) that may exist on the sensor membrane surface, which occlusion(s) may temporarily block passage of glucose into the sensor and thus cause the signal to dip.
  • In addition, (substantially) glucose-independent impedance data, such as, e.g., values for higher-frequency phase angle and/or imaginary impedance at 1 kHz and higher frequencies, provides information on loss of sensor sensitivity during extended wear, which sensitivity loss may potentially be due to local oxygen deficit at the insertion site. In this regard, the underlying mechanism for oxygen deficiency-led sensitivity loss may be described as follows: when local oxygen is deficient, sensor output (i.e., Isig and SG) will be dependent on oxygen rather than glucose and, as such, the sensor will lose sensitivity to glucose. Other markers, including 0.1 Hz real impedance, the counter electrode voltage (Vcntr), and EIS-induced spikes in the Isig may also be used for the detection of oxygen deficiency-led sensitivity loss. Moreover, in a redundant sensor system, the relative differences in 1 kHz real impedance, 1 kHz imaginary impedance, and 0.1 Hz real impedance between two or more working electrodes may be used for the detection of sensitivity loss due to biofouling.
  • In accordance with embodiments of the invention, EIS-based sensor diagnostics entails consideration and analysis of EIS data relating to one or more of at least three primary factors, i.e., potential sensor failure modes: (1) signal start-up; (2) signal dip; and (3) sensitivity loss. Significantly, the discovery herein that a majority of the impedance-related parameters that are used in such diagnostic analyses and procedures can be studied at a frequency, or within a range of frequencies, where the parameter is substantially analyte-independent allows for implementation of sensor-diagnostic procedures independently of the level of the analyte in a patient's body. Thus, while EIS-based sensor diagnostics may be triggered by, e.g., large fluctuations in Isig, which is analyte-dependent, the impedance-related parameters that are used in such sensor diagnostic procedures are themselves substantially independent of the level of the analyte. As will be explored in more detail below, it has also been found that, in a majority of situations where glucose may be seen to have an effect on the magnitude (or other characteristic) of an EIS-based parameter, such effect is usually small enough—e.g., at least an order of magnitude difference between the EIS-based measurement and the glucose effect thereon—such that it can be filtered out of the measurement, e.g., via software in the IC.
  • By definition, “start-up” refers to the integrity of the sensor signal during the first few hours (e.g., t=0-6 hours) after insertion. For example, in current devices, the signal during the first 2 hours after insertion is deemed to be unreliable and, as such, the sensor glucose values are blinded to the patient/user. In situations where the sensor takes an extended amount of time to hydrate, the sensor signal is low for several hours after insertion. With the use of EIS, additional impedance information is available (by running an EIS procedure) right after the sensor has been inserted. In this regard, the total impedance equation may be used to explain the principle behind low-startup detection using 1 kHz real impedance. At relatively higher frequencies—in this case, 1 kHz and above—imaginary impedance is very small (as confirmed with in-vivo data), such that total impedance reduces to:
  • Z t ( ω ) = R s + R p 1 + ω 2 R p 2 C d 2
  • As sensor wetting is gradually completed, the double layer capacitance (Cd) increases. As a result, the total impedance will decrease because, as indicated in the equation above, total impedance is inversely proportional to Cd. This is illustrated in the form of the intercept 1600 on the real impedance axis shown, e.g., in FIG. 16B. Importantly, the 1 kHz imaginary impedance can also be used for the same purpose, as it also includes, and is inversely proportional to, a capacitance component.
  • Another marker for low startup detection is Nyquist slope, which relies solely on the relatively lower-frequency impedance which, in turn, corresponds to the Warburg impedance component of total impedance (see, e.g., FIG. 15B). FIG. 22 shows a Nyquist plot for a normally-functioning sensor, where Arrow A is indicative of the progression of time, i.e., sensor wear time, starting from t=0. Thus, EIS at the relatively-lower frequencies is performed right after sensor insertion (time t=0), which generates real and imaginary impedance data that is plotted with a first linear fit 2200 having a first (Nyquist) slope. At a time interval after t=0, a second (lower) frequency sweep is run that produces a second linear fit 2210 having a second (Nyquist) slope larger than the first Nyquist slope, and so on. As the sensor becomes more hydrated, the Nyquist slope increases, and the intercept decrease, as reflected by the lines 2200, 2210, etc. becoming steeper and moving closer to the Y-axis. In connection with low startup detection, clinical data indicates that there is typically a dramatic increase of Nyquist slope after sensor insertion and initialization, which is then stabilized to a certain level. One explanation for this is that, as the sensor is gradually wetted, the species diffusivity as well as concentration undergo dramatic change, which is reflected in Warburg impedance.
  • In FIG. 23A, the Isig 2230 for a first working electrode WE1 starts off lower than expected (at about 10 nA), and takes some time to catch up with the Isig 2240 for a second working electrode WE2. Thus, in this particular example, WE1 is designated as having a low start-up. The EIS data reflects this low start-up in two ways. First, as shown in FIG. 23A, the real impedance at 1 kHz (2235) of WE1 is much higher than the 1 kHz real impedance 2245 of WE2. Second, when compared to the Nyquist slope for WE2 (FIG. 23C), the Nyquist slope for WE1 (FIG. 23B) starts out lower, has a larger intercept 2237, and takes more time to stabilize. As will be discussed later, these two signatures—the 1 kHz real impedance and the Nyquist slope—can be used as diagnostic inputs in a fusion algorithm to decide which of the two electrodes can carry a higher weight when the fused signal is calculated. In addition, one or both of these markers may be used in a diagnostic procedure to determine whether the sensor, as a whole, is acceptable, or whether it should be terminated and replaced.
  • By definition, signal (or Isig) dips refer to instances of low sensor signal, which are mostly temporary in nature, e.g., on the order of a few hours. Such low signals may be caused, for example, by some form of biological occlusion on the sensor surface, or by pressure applied at the insertion site (e.g., while sleeping on the side). During this period, the sensor data is deemed to be unreliable; however, the signal does recover eventually. In the EIS data, this type of signal dip—as opposed to one that is caused by a glycemic change in the patient's body—is reflected in the 1 kHz real impedance data, as shown in FIG. 24.
  • Specifically, in FIG. 24, both the Isig 2250 for the first working electrode WE1 and the Isig 2260 for the second working electrode WE2 start out at about 25 nA at the far left end (i.e., at 6 pm). As time progresses, both Isigs fluctuate, which is reflective of glucose fluctuations in the vicinity of the sensor. For about the first 12 hours or so (i.e., until about 6 am), both Isigs are fairly stable, as are their respective 1 kHz real impedances 2255, 2265. However, between about 12 and 18 hours—i.e., between 6 am and noon—the Isig 2260 for WE2 starts to dip, and continues a downward trend for the next several hours, until about 9 pm. During this period, the Isig 2250 for WE1 also exhibits some dipping, but Isig 2250 is much more stable, and dips quite a bit less, than Isig 2260 for WE2. The behavior of the Isigs for WE1 and WE2 is also reflected in their respective 1 kHz real impedance data. Thus, as shown in FIG. 24, during the time period noted above, while the 1 kHz real impedance for WE1 (2255) remains fairly stable, there is a marked increase in the 1 kHz real impedance for WE2 (2265).
  • By definition, sensitivity loss refers to instances where the sensor signal (Isig) becomes low and non-responsive for an extended period of time, and is usually unrecoverable. Sensitivity loss may occur for a variety of reasons. For example, electrode poisoning drastically reduces the active surface area of the working electrode, thereby severely limiting current amplitude. Sensitivity loss may also occur due to hypoxia, or oxygen deficit, at the insertion site. In addition, sensitivity loss my occur due to certain forms of extreme surface occlusion (i.e., a more permanent form of the signal dip caused by biological or other factors) that limit the passage of both glucose and oxygen through the sensor membrane, thereby lowering the number/frequency of the chemical reactions that generate current in the electrode and, ultimately, the sensor signal (Isig). It is noted that the various causes of sensitivity loss mentioned above apply to both short-term (7-10 day wear) and long term (6 month wear) sensors.
  • In the EIS data, sensitivity loss is often preceded by an increase in the absolute value of phase (|phase|) and of the imaginary impedance (|imaginary impedance|) at the relatively higher frequency ranges (e.g., 128 Hz and above, and 1 kHz and above, respectively). FIG. 25A shows an example of a normally-functioning glucose sensor where the sensor current 2500 is responsive to glucose—i.e., Isig 2500 tracks glucose fluctuations—but all relevant impedance outputs, such as, e.g., 1 kHz real impedance 2510, 1 kHz imaginary impedance 2530, and phase for frequencies at or above about 128 Hz (2520), remain steady, as they are substantially glucose-independent.
  • Specifically, the top graph in FIG. 25A shows that, after the first few hours, the 1 kHz real impedance 2510 holds fairly steady at about 5 kOhms (and the 1 kHz imaginary impedance 2530 holds fairly steady at about −400 Ohms). In other words, at 1 kHz, the real impedance data 2510 and the imaginary impedance data 2530 are substantially glucose-independent, such that they can be used as signatures for, or independent indicators of, the health, condition, and ultimately, reliability of the specific sensor under analysis. However, as mentioned previously, different impedance-related parameters may exhibit glucose-independence at different frequency ranges, and the range, in each case, may depend on the overall sensor design, e.g., electrode type, surface area of electrode, thickness of membrane, permeability of membrane, etc.
  • Thus, in the example FIG. 25B—for a 90% short tubeless electrode design—the top graph again shows that sensor current 2501 is responsive to glucose, and that, after the first few hours, the 1 kHz real impedance 2511 holds fairly steady at about 7.5 kOhms. The bottom graph in FIG. 25B shows real impedance data for frequencies between 0.1 Hz (2518) and 1 kHz (2511). As can be seen, the real impedance data at 0.1 Hz (2518) is quite glucose-dependent. However, as indicated by reference numerals 2516, 2514, and 2512, real impedance becomes more and more glucose-independent as the frequency increases from 0.1 Hz to 1 kHz, i.e., for impedance data measured at frequencies closer to 1 kHz.
  • Returning to FIG. 25A, the middle graph shows that the phase 2520 at the relatively-higher frequencies is substantially glucose-independent. It is noted, however, that “relatively-higher frequencies” in connection with this parameter (phase) for the sensor under analysis means frequencies of 128 Hz and above. In this regard, the graph shows that the phase for all frequencies between 128 Hz and 8 kHz is stable throughout the period shown. On the other hand, as can be seen in the bottom graph of FIG. 25C, while the phase 2522 at 128 Hz (and above) is stable, the phase 2524 fluctuates—i.e., it becomes more and more glucose-dependent, and to varying degrees—at frequencies that are increasingly smaller than 128 Hz. It is noted that the electrode design for the example of FIG. 25C is the same as that used in FIG. 25B, and that the top graph in the former is identical to the top graph in the latter.
  • FIG. 26 shows an example of sensitivity loss due to oxygen deficiency at the insertion site. In this case, the insertion site becomes oxygen deprived just after day 4 (designated by dark vertical line in FIG. 26), causing the sensor current 2600 to become low and non-responsive. The 1 kHz real impedance 2610 remains stable, indicating no physical occlusion on the sensor. However, as shown by the respective downward arrows, changes in the relatively higher-frequency phase 2622 and 1 kHz imaginary impedance 2632 coincide with loss in sensitivity, indicating that this type of loss is due to an oxygen deficit at the insertion site. Specifically, FIG. 26 shows that the phase at higher frequencies (2620) and the 1 kHz imaginary impedance (2630) become more negative prior to the sensor losing sensitivity—indicated by the dark vertical line—and continue their downward trend as the sensor sensitivity loss continues. Thus, as noted above, this sensitivity loss is preceded, or predicted, by an increase in the absolute value of phase (|phase|) and of the imaginary impedance (|imaginary impedance|) at the relatively higher frequency ranges (e.g., 128 Hz and above, and 1 kHz and above, respectively).
  • The above-described signatures may be verified by in-vitro testing, an example of which is shown in FIG. 27. FIG. 27 shows the results of in-vitro testing of a sensor, where oxygen deficit at different glucose concentrations is simulated. In the top graph, the Isig fluctuates with the glucose concentration as the latter is increased from 100 mg/dl (2710) to 200 mg/dl (2720), 300 mg/dl (2730), and 400 mg/dl (2740), and then decreased back down to 200 and/dl (2750). In the bottom graph, the phase at the relatively-higher frequencies is generally stable, indicating that it is glucose-independent. However, at very low oxygen concentrations, such as, e.g., at 0.1% O2, the relatively high-frequency phase fluctuates, as indicated by the encircled areas and arrows 2760, 2770. It is noted that the magnitude and/or direction (i.e., positive or negative) of fluctuation depend on various factors. For example, the higher the ratio of glucose concentration to oxygen concentration, the higher the magnitude of the fluctuation in phase. In addition, the specific sensor design, as well as the age of the sensor (i.e., as measured by time after implant), affect such fluctuations. Thus, e.g., the older a sensor is, the more susceptible it is to perturbations.
  • FIGS. 28A-28D show another example of oxygen deficiency-led sensitivity loss with redundant working electrodes WE1 and WE2. As shown in FIG. 28A, the 1 kHz real impedance 2810 is steady, even as sensor current 2800 fluctuates and eventually becomes non-responsive. Also, as before, the change in 1 kHz imaginary impedance 2820 coincides with the sensor's loss of sensitivity. In addition, however, FIG. 28B shows real impedance data and imaginary impedance data (2830 and 2840, respectively) at 0.105 Hz. The latter, which may be more commonly referred to as “0.1 Hz data”, indicates that, whereas imaginary impedance at 0.1 Hz appears to be fairly steady, 0.1 Hz real impedance 2830 increases considerably as the sensor loses sensitivity. Moreover, as shown in FIG. 28C, with loss of sensitivity due to oxygen deficiency, Vcntr 2850 rails to 1.2 Volts.
  • In short, the diagrams illustrate the discovery that oxygen deficiency-led sensitivity loss is coupled with lower 1 kHz imaginary impedance (i.e., the latter becomes more negative), higher 0.105 Hz real impedance (i.e., the latter becomes more positive), and Vcntr rail. Moreover, the oxygen-deficiency process and Vcntr-rail are often coupled with the increase of the capacitive component in the electrochemical circuit. It is noted that, in some of the diagnostic procedures to be described later, the 0.105 Hz real impedance may not be used, as it appears that this relatively lower-frequency real impedance data may be analyte-dependent.
  • Finally, in connection with the example of FIGS. 28A-28D, it is noted that 1 kHz or higher-frequency impedance measurement typically causes EIS-induced spikes in the Isig. This is shown in FIG. 28D, where the raw Isig for WE2 is plotted against time. The drastic increase of Isig when the spike starts is a non-Faradaic process, due to double-layer capacitance charge. Thus, oxygen deficiency-led sensitivity loss may also be coupled with higher EIS-induced spikes, in addition to lower 1 kHz imaginary impedance, higher 0.105 Hz real impedance, and Vcntr rail, as discussed above.
  • FIG. 29 illustrates another example of sensitivity loss. This case may be thought of as an extreme version of the Isig dip described above in connection with FIG. 24. Here, the sensor current 2910 is observed to be low from the time of insertion, indicating that there was an issue with an insertion procedure resulting in electrode occlusion. The 1 kHz real-impedance 2920 is significantly higher, while the relatively higher-frequency phase 2930 and the 1 kHz imaginary impedance 2940 are both shifted to much more negative values, as compared to the same parameter values for the normally-functioning sensor shown in FIG. 25A. The shift in the relatively higher-frequency phase 2930 and 1 kHz imaginary impedance 2940 indicates that the sensitivity loss may be due to an oxygen deficit which, in turn, may have been caused by an occlusion on the sensor surface.
  • FIGS. 30A-30D show data for another redundant sensor, where the relative differences in 1 kHz real impedance and 1 kHz imaginary impedance, as well as 0.1 Hz real impedance, between two or more working electrodes may be used for the detection of sensitivity loss due to biofouling. In this example, WE1 exhibits more sensitivity loss than WE2, as is evident from the higher 1 kHz real impedance 3010, lower 1 kHz imaginary impedance 3020, and much higher real impedance at 0.105 kHz (3030) for WE2. In addition, however, in this example, Vcntr 3050 does not rail. Moreover, as shown in FIG. 30D, the height of the spikes in the raw Isig data does not change much as time progresses. This indicates that, for sensitivity loss due to biofouling, Vcntr-rail and the increase in spike height are correlated. In addition, the fact that the height of the spikes in the raw Isig data does not change much with time indicates that the capacitive component of the circuit does not change significantly with time, such that sensitivity loss due to biofouling is related to the resistance component of the circuit (i.e., diffusion).
  • Various of the above-described impedance-related parameters may be used, either individually or in combination, as inputs into: (1) EIS-based sensor diagnostic procedures; and/or (2) fusion algorithms for generating more reliable sensor glucose values. With regard to the former, FIG. 31 illustrates how EIS-based data—i.e., impedance-related parameters, or characteristics—may be used in a diagnostic procedure to determine, in real time, whether a sensor is behaving normally, or whether it should be replaced.
  • The diagnostic procedure illustrated in the flow diagram of FIG. 31 is based on the collection of EIS data on a periodic basis, such as, e.g., hourly, every half hour, every 10 minutes, or at any other interval—including continuously—as may be appropriate for the specific sensor under analysis. At each such interval, EIS may be run for an entire frequency spectrum (i.e., a “full sweep”), or it may be run for a selected frequency range, or even at a single frequency. Thus, for example, for an hourly data collection scheme, EIS may be performed at frequencies in the μHz to MHz range, or it may be run for a narrower range of frequencies, such as, e.g., between about 0.1 Hz and about 8 kHz, as discussed hereinabove. In embodiments of the invention, EIS data acquisition may be implemented alternatingly between a full sweep and an narrower-range spectrum, or in accordance with other schemes.
  • The temporal frequency of EIS implementation and data collection may be dictated by various factors. For example, each implementation of EIS consumes a certain amount of power, which is typically provided by the sensor's battery, i.e., the battery running the sensor electronics, including the ASIC which is described later. As such, battery capacity, as well as the remaining sensor life, may help determine the number of times EIS is run, as well as the breadth of frequencies sampled for each such run. In addition, embodiments of the invention envision situations that may require that an EIS parameter at a specific frequency (e.g., real impedance at 1 kHz) be monitored based on a first schedule (e.g., once every few seconds, or minutes), while other parameters, and/or the same parameter at other frequencies, can be monitored based on a second schedule (e.g., less frequently). In these situations, the diagnostic procedure can be tailored to the specific sensor and requirements, such that battery power may be preserved, and unnecessary and/or redundant EIS data acquisition may be avoided.
  • It is noted that, in embodiments of the invention, a diagnostic procedure, such as the one shown in FIG. 31, entails a series of separate “tests” which are implemented in order to perform real-time monitoring of the sensor. The multiple tests, or markers—also referred to as “multi markers”—are implemented because each time EIS is run (i.e., each time an EIS procedure is performed), data may be gathered about a multiplicity of impedance-based parameters, or characteristics, which can be used to detect sensor condition or quality, including, e.g., whether the sensor has failed or is failing. In performing sensor diagnostics, sometimes, there can be a diagnostic test that may indicate a failure, whereas other diagnostic(s) may indicate no failure. Therefore, the availability of multiple impedance-related parameters, and the implementation of a multi-test procedure, are advantageous, as some of the multiplicity of tests may act as validity checks against some of the other tests. Thus, real-time monitoring using a multi-marker procedure may include a certain degree of built-in redundancy.
  • With the above in mind, the logic of the diagnostic procedure shown in FIG. 31 begins at 3100, after the sensor has been inserted/implanted, and an EIS run has been made, so as to provide the EIS data as input. At 3100, using the EIS data as input, it is first determined whether the sensor is still in place. Thus, if the |Z| slope is found to be constant across the tested frequency band (or range), and/or the phase angle is about −90°, it is determined that the sensor is no longer in place, and an alert is sent, e.g., to the patient/user, indicating that sensor pullout has occurred. The specific parameters (and their respective values) described herein for detecting sensor pullout are based on the discovery that, once the sensor is out of the body and the membrane is no longer hydrated, the impedance spectrum response appears just like a capacitor.
  • If it is determined that the sensor is still in place, the logic moves to step 3110 to determine whether the sensor is properly initialized. As shown, the “Init. Check” is performed by determining: (i) whether |(Zn−Z1)/Z1|>30% at 1 kHz, where Z1 is the real impedance measured at a first time, and Zn is the measured impedance at the next interval, at discussed above; and (2) whether the phase angle change is greater than 10° at 0.1 Hz. If the answer to either one of the questions is “yes”, then the test is satisfactory, i.e., the Test 1 is not failed. Otherwise, the Test 1 is marked as a failure.
  • At step 3120, Test 2 asks whether, at a phase angle of −45°, the difference in frequency between two consecutive EIS runs (f2−f1) is greater than 10 Hz. Again, a “No” answer is marked as a fail; otherwise, Test 2 is satisfactorily met.
  • Test 3 at step 3130 is a hydration test. Here, the inquiry is whether the current impedance Zn is less than the post-initialization impedance Zpi at 1 kHz. If it is, then this test is satisfied; otherwise, Test 3 is marked as a fail. Test 4 at step 3140 is also a hydration test, but this time at a lower frequency. Thus, this test asks whether Zn is less than 300 kOhms at 0.1 Hz during post-initialization sensor operation. Again, a “No” answer indicates that the sensor has failed Test 4.
  • At step 3150, Test 5 inquires whether the low-frequency Nyquist slope is globally increasing from 0.1 Hz to 1 Hz. As discussed previously, for a normally-operating sensor, the relatively lower-frequency Nyquist slope should be increasing over time. Thus, this test is satisfied if the answer to the inquiry is “yes”; otherwise, the test is marked as failed.
  • Step 3160 is the last test for this embodiment of the diagnostic procedure. Here, the inquiry is whether real impedance is globally decreasing. Again, as was discussed previously, in a normally-operating sensor, it is expected that, as time goes by, the real impedance should be decreasing. Therefore, a “Yes” answer here would mean that the sensor is operating normally; otherwise, the sensor fails Test 6.
  • Once all 6 tests have been implemented, a decision is made at 3170 as to whether the sensor is operating normally, or whether it has failed. In this embodiment, a sensor is determined to be functioning normally (3172) if it passes at least 3 out of the 6 tests. Put another way, in order to be determined to have failed (3174), the sensor must fail at least 4 out of the 6 tests. In alternative embodiments, a different rule may be used to assess normal operation versus sensor failure. In addition, in embodiments of the invention, each of the tests may be weighted, such that the assigned weight reflects, e.g., the importance of that test, or of the specific parameter(s) queried for that test, in determining overall sensor operation (normal vs. failed). For example, one test may be weighted twice as heavily as another, but only half as heavily as a third test, etc.
  • In other alternative embodiments, a different number of tests and/or a different set of EIS-based parameters for each test may be used. FIGS. 32A and 32B show an example of a diagnostic procedure for real-time monitoring that includes 7 tests. Referring to FIG. 32A, the logic begins at 3200, after the sensor has been inserted/implanted, and an EIS procedure has been performed, so as to provide the EIS data as input. At 3200, using the EIS data as input, it is first determined whether the sensor is still in place. Thus, if the |Z| slope is found to be constant across the tested frequency band (or range), and/or the phase angle is about −90°, it is determined that the sensor is no longer in place, and an alert is sent, e.g., to the patient/user, indicating that sensor pullout has occurred. If, on the other hand, the sensor is determined to be in place, the logic moves to initiation of diagnostic checks (3202).
  • At 3205, Test 1 is similar to Test 1 of the diagnostic procedure discussed above in connection with FIG. 31, except that the instant Test 1 specifies that the later measurement Zn is taken 2 hours after the first measurement. As such, in this example, Zn=Z2hr. More specifically, Test 1 compares the real impedance 2 hours after (sensor implantation and) initialization to the pre-initialization value. Similarly, the second part of Test 1 asks whether the difference between the phase 2 hours after initialization and the pre-initialization phase is greater than 10° at 0.1 Hz. As before, if the answer to either one of the inquiries is affirmative, then it is determined that the sensor is hydrated normally and initialized, and Test 1 is satisfied; otherwise, the sensor fails this test. It should be noted that, even though the instant test inquires about impedance and phase change 2 hours after initialization, the time interval between any two consecutive EIS runs may be shorter or longer, depending on a variety of factors, including, e.g., sensor design, the level of electrode redundancy, the degree to which the diagnostic procedure includes redundant tests, battery power, etc.
  • Moving to 3210, the logic next performs a sensitivity-loss check by inquiring whether, after a 2-hour interval (n+2), the percentage change in impedance magnitude at 1 kHz, as well as that in the Isig, is greater than 30%. If the answer to both inquiries is “yes”, then it is determined that the sensor is losing sensitivity and, as such, Test 2 is determined to be failed. It is noted that, although Test 2 is illustrated herein based on a preferred percentage difference of 30%, in other embodiments, the percentage differences in the impedance magnitude at 1 kHz and in the Isig may fall within the range 10%-50% for purposes of conducting this test.
  • Test 3 (at 3220) is similar to Test 5 of the algorithm illustrated in FIG. 31. Here, as before, the question is whether the low-frequency Nyquist slope is globally increasing from 0.1 Hz to 1 Hz. If it is, then this test is passed; otherwise, the test is failed. As shown in 3220, this test is also amenable to setting a threshold, or an acceptable range, for the percent change in the low-frequency Nyquist slope, beyond which the sensor may be deemed to be failed or, at the very least, may trigger further diagnostic testing. In embodiments of the invention, such threshold value/acceptable range for the percent change in low-frequency Nyquist slope may fall within a range of about 2% to about 20%. In some preferred embodiments, the threshold value may be about 5%.
  • The logic next moves to 3230, which is another low-frequency test, this time involving the phase and the impedance magnitude. More specifically, the phase test inquires whether the phase at 0.1 Hz is continuously increasing over time. If it is, then the test is failed. As with other tests where the parameter's trending is monitored, the low-frequency phase test of Test 4 is also amenable to setting a threshold, or an acceptable range, for the percent change in the low-frequency phase, beyond which the sensor may be deemed to be failed or, at the very least, raise a concern. In embodiments of the invention, such threshold value/acceptable range for the percent change in low-frequency phase may fall within a range of about 5% to about 30%. In some preferred embodiments, the threshold value may be about 10%.
  • As noted, Test 4 also includes a low-frequency impedance magnitude test, where the inquiry is whether the impedance magnitude at 0.1 Hz is continuously increasing over time. If it is, then the test is failed. It is noted that Test 4 is considered “failed” if either the phase test or the impedance magnitude test is failed. The low-frequency impedance magnitude test of Test 4 is also amenable to setting a threshold, or an acceptable range, for the percent change in the low-frequency impedance magnitude, beyond which the sensor may be deemed to be failed or, at the very least, raise a concern. In embodiments of the invention, such threshold value/acceptable range for the percent change in low-frequency impedance magnitude may fall within a range of about 5% to about 30%. In some preferred embodiments, the threshold value may be about 10%, where the range for impedance magnitude in normal sensors is generally between about 100 KOhms and about 200 KOhms.
  • Test 5 (at 3240) is another sensitivity loss check that may be thought of as supplemental to Test 2. Here, if both the percentage change in the Isig and the percentage change in the impedance magnitude at 1 kHz are greater than 30%, then it is determined that the sensor is recovering from sensitivity loss. In other words, it is determined that the sensor had previously undergone some sensitivity loss, even if the sensitivity loss was not, for some reason, detected by Test 2. As with Test 2, although Test 5 is illustrated based on a preferred percentage difference of 30%, in other embodiments, the percentage differences in the Isig and the impedance magnitude at 1 kHz may fall within the range 10%-50% for purposes of conducting this test.
  • Moving to 3250, Test 6 provides a sensor functionality test with specific failure criteria that have been determined based on observed data and the specific sensor design. Specifically, in one embodiment, a sensor may be determined to have failed and, as such, to be unlikely to respond to glucose, if at least two out of the following three criteria are met: (1) Isig is less than 10 nA; and (2) the imaginary impedance at 1 kHz is less than −1500 Ohm; and (3) the phase at 1 kHz is less than −15°. Thus, Test 6 is determined to have been passed if any two of (1)-(3) are not met. It is noted that, in other embodiments, the Isig prong of this test may be failed if the Isig is less than about 5 nA to about 20 nA. Similarly, the second prong may be failed if the imaginary impedance at 1 kHz is less than about −1000 Ohm to about −2000 Ohms. Lastly, the phase prong may be failed if the phase at 1 kHz is less than about −10° to about −20°.
  • Lastly, step 3260 provides another sensitivity check, wherein the parameters are evaluated at low frequency. Thus, Test 7 inquires whether, at 0.1 Hz, the magnitude of the difference between the ratio of the imaginary impedance to the Isig (n+2), on the one hand, and the pervious value of the ratio, on the other, is larger than 30% of the magnitude of the previous value of the ratio. If it is, then the test is failed; otherwise, the test is passed. Here, although Test 7 is illustrated based on a preferred percentage difference of 30%, in other embodiments, the percentage difference may fall within the range 10%-50% for purposes of conducting this test.
  • Once all 7 tests have been implemented, a decision is made at 3270 as to whether the sensor is operating normally, or whether an alert should be sent out, indicating that the sensor has failed (or may be failing). As shown, in this embodiment, a sensor is determined to be functioning normally (3272) if it passes at least 4 out of the 7 tests. Put another way, in order to be determined to have failed, or to at least raise a concern (3274), the sensor must fail at least 4 out of the 7 tests. If it is determined that the sensor is “bad” (3274), an alert to that effect may be sent, e.g., to the patient/user. As noted previously, in alternative embodiments, a different rule may be used to assess normal operation versus sensor failure/concern. In addition, in embodiments of the invention, each of the tests may be weighted, such that the assigned weight reflects, e.g., the importance of that test, or of the specific parameter(s) queried for that test, in determining overall sensor operation (normal vs. failed).
  • As was noted previously, in embodiments of the invention, various of the above-described impedance-related parameters may be used, either individually or in combination, as inputs into one or more fusion algorithms for generating more reliable sensor glucose values. Specifically, it is known that, unlike a single-sensor (i.e., a single-working-electrode) system, multiple sensing electrodes provide higher-reliability glucose readouts, as a plurality of signals, obtained from two or more working electrodes, may be fused to provide a single sensor glucose value. Such signal fusion utilizes quantitative inputs provided by EIS to calculate the most reliable output sensor glucose value from the redundant working electrodes. It is noted that, while the ensuing discussion may describe various fusion algorithms in terms of a first working electrode (WE1) and a second working electrode (WE2) as the redundant electrodes, this is by way of illustration, and not limitation, as the algorithms and their underlying principles described herein are applicable to, and may be used in, redundant sensor systems having more than 2 working electrodes.
  • FIGS. 33A and 33B show top-level flowcharts for two alternative methodologies, each of which includes a fusion algorithm. Specifically, FIG. 33A is a flowchart involving a current (Isig)-based fusion algorithm, and FIG. 33B is a flowchart directed to sensor glucose (SG) fusion. As may be seen from the diagrams, the primary difference between the two methodologies is the time of calibration. Thus, FIG. 33A shows that, for Isig fusion, calibration 3590 is performed after the fusion 3540 is completed. That is, redundant Isigs from WE1 to WEn are fused into a single Isig 3589, which is then calibrated to produce a single sensor glucose value 3598. For SG fusion, on the other hand, calibration 3435 is completed for each individual Isig from WE1 to WEn to produce calibrated SG values (e.g., 3436, 3438) for each of the working electrodes. Thus, SG fusion algorithms provide for independent calibration of each of the plurality of Isigs, which may be preferred in embodiments of the invention. Once calibrated, the plurality of calibrated SG values is fused into a single SG value 3498.
  • It is important to note that each of flowcharts shown in FIGS. 33A and 33B includes a spike filtering process (3520, 3420). As was described above in the discussion relating to sensitivity loss, 1 kHz or higher-frequency impedance measurements typically cause EIS-induced spikes in the Isig. Therefore, once an EIS procedure has been performed for each of the electrodes WE1 to WEn, for both SG fusion and Isig fusion, it is preferable to first filter the Isigs 3410, 3412, etc. and 3510, 3512, etc. to obtain respective filtered Isigs 3422, 3424, etc. and 3522, 3524, etc. The filtered Isigs are then either used in Isig fusion, or first calibrated and then used in SG fusion, as detailed below. As will become apparent in the ensuing discussion, both fusion algorithms entail calculation and assignment of weights based on various factors.
  • FIG. 34 shows the details of the fusion algorithm 3440 for SG fusion. Essentially, there are four factors that need to be checked before the fusion weights are determined. First, integrity check 3450 involves determining whether each of the following parameters is within specified ranges for normal sensor operation (e.g., predetermined lower and upper thresholds): (i) Isig; (ii) 1 kHz real and imaginary impedances; (iii) 0.105 Hz real and imaginary impedances; and (iv) Nyquist slope. As shown, integrity check 3450 includes a Bound Check 3452 and a Noise Check 3456, wherein, for each of the Checks, the above-mentioned parameters are used as input parameters. It is noted that, for brevity, real and/or imaginary impedances, at one or more frequencies, appear on FIGS. 33A-35 simply as “Imp” for impedance. In addition, both real and imaginary impedances may be calculated using impedance magnitude and phase (which is also shown as an input on FIGS. 33A and 33B).
  • The output from each of the Bound Check 3452 and the Noise Check 3458 is a respective reliability index (RI) for each of the redundant working electrodes. Thus, the output from the Bound Check includes, e.g., RI_bound_We1 (3543) and RI_bound_We2 (3454). Similarly, for the Noise Check, the output includes, e.g., RI_noise_We1 (3457) and RI_noise_We2 (3458). The bound and noise reliability indices for each working electrode are calculated based on compliance with the above-mentioned ranges for normal sensor operation. Thus, if any of the parameters falls outside the specified ranges for a particular electrode, the reliability index for that particular electrode decreases.
  • It is noted that the threshold values, or ranges, for the above-mentioned parameters may depend on various factors, including the specific sensor and/or electrode design. Nevertheless, in one preferred embodiment, typical ranges for some of the above-mentioned parameters may be, e.g., as follows: Bound threshold for 1 kHz real impedance=[0.3e+4 2e+4]; Bound threshold for 1 kHz imaginary impedance=[−2e+3, 0]; Bound threshold for 0.105 Hz real impedance=[2e+4 7e+4]; Bound threshold for 0.105 Hz imaginary impedance=[−2e+5−0.25e+5]; and Bound threshold for Nyquist slope=[2 5]. Noise may be calculated, e.g., using second order central difference method where, if noise is above a certain percentage (e.g., 30%) of median value for each variable buffer, it is considered to be out of noise bound.
  • Second, sensor dips may be detected using sensor current (Isig) and 1 kHz real impedance. Thus, as shown in FIG. 34, Isig and “Imp” are used as inputs for dips detection 3460. Here, the first step is to determine whether there is any divergence between Isigs, and whether any such divergence is reflected in 1 kHz real impedance data. This may be accomplished by using mapping 3465 between the Isig similarity index (RI_sim_isig12) 3463 and the 1 kHz real impedance similarity index (RI_sim_imp12) 3464. This mapping is critical, as it helps avoid false positives in instances where a dip is not real. Where the Isig divergence is real, the algorithm will select the sensor with the higher Isig.
  • In accordance with embodiments of the invention, the divergence/convergence of two signals (e.g., two Isigs, or two 1 kHz real impedance data points) can be calculated as follows:

  • diff_va1=abs(va1−(va1+va2)/2);

  • diff_va2=abs(va2−(va1+va2)/2);

  • RI_sim=1−(diff_va1+diff_va2)/(mean(abs(va1+va2))/4)
  • where va1 and va2 are two variables, and RI_sim (similarity index) is the index to measure the convergence or divergence of the signals. In this embodiment, RI_sim must be bound between 0 and 1. Therefore, if RI_sim as calculated above is less than 0, it will be set to 0, and if it is higher than 1, it will be set to 1.
  • The mapping 3465 is performed by using ordinary linear regression (OLR). However, when OLR does not work well, a robust median slope linear regression (RMSLR) can be used. For Isig similarity index and 1 kHz real impedance index, for example, two mapping procedures are needed: (i) Map Isig similarity index to 1 kHz real impedance similarity index; and (ii) map 1 kHz real impedance similarity index to Isig similarity index. Both mapping procedures will generate two residuals: res12 and res21. Each of the dip reliability indices 3467, 3468 can then be calculated as:

  • RI_dip=1−(res12+res21)/(RI_sim_isig+RI_sim_1K_real_impedance).
  • The third factor is sensitivity loss 3470, which may be detected using 1 kHz imaginary impedance trending in, e.g., the past 8 hours. If one sensor's trending turns negative, the algorithm will rely on the other sensor. If both sensors lose sensitivity, then a simple average is taken. Trending may be calculated by using a strong low-pass filter to smooth over the 1 kHz imaginary impedance, which tends to be noisy, and by using a correlation coefficient or linear regression with respect to time during, e.g., the past 8 hours to determine whether the correlation coefficient is negative or the slope is negative. Each of the sensitivity loss reliability indices 3473, 3474 is then assigned a binary value of 1 or 0.
  • The total reliability index (RI) for each of we1, we2, . . . wen is calculated as follows:
  • RI_we 1 = RI_dip _we 1 × RI_sensitivity _loss _we 1 × RI_bound _we 1 × RI_noise _we 1 RI_we 2 = RI_dip _we 2 × RI_sensitivity _loss _we 2 × RI_bound _we 2 × RI_noise _we 2 RI_we 3 = RI_dip _we 3 × RI_sensitivity _loss _we 3 × RI_bound _we 3 × RI_noise _we 3 RI_we 4 = RI_dip _we 4 × RI_sensitivity _loss _we 4 × RI_bound _we 4 × RI_noise _we 4 RI_we n = RI_dip _we n × RI_sensitivity _loss _we n × RI_bound _we n × RI_noise _we n
  • Having calculated the respective reliability indices of the individual working electrodes, the weight for each of the electrodes may be calculated as follow:
  • weight_we 1 = RI_we 1 / ( RI_we 1 + RI_we 2 + RI_we 3 + RI_we 4 + + RI_we n ) weight_we 2 = RI_we 2 / ( RI_we 1 + RI_we 2 + RI_we 3 + RI_we 4 + + RI_we n ) weight_we 3 = RI_we 3 / ( RI_we 1 + RI_we 2 + RI_we 3 + RI_we 4 + + RI_we n ) weight_we 4 = RI_we 4 / ( RI_we 1 + RI_we 2 + RI_we 3 + RI_we 4 + + RI_we n ) weight_we n = RI_we n / ( RI_we 1 + RI_we 2 + RI_we 3 + RI_we 4 + + RI_we n )
  • Based on the above, the fused SG 3498 is then calculated as follows:

  • SG=weight_we 1 ×SG_we 1+weight_we 2 ×SG_we 2+weight_we 3 ×SG_we 3+weight_we 4 ×SG_we 4+ . . . +weight_we n ×SG_we n
  • The last factor relates to artifacts in the final sensor readout, such as may be caused by instant weight change of sensor fusion. This may be avoided by either applying a low-pass filter 3480 to smooth the RI for each electrode, or by applying a low-pass filter to the final SG. When the former is used, the filtered reliability indices—e.g., RI_We1* and RI_We2* (3482, 3484)—are used in the calculation of the weight for each electrode and, therefore, in the calculation of the fused SG 3498.
  • FIG. 35 shows the details of the fusion algorithm 3540 for Isig fusion. As can be seen, this algorithm is substantially similar to the one shown in FIG. 34 for SG fusion, with two exceptions. First, as was noted previously, for Isig fusion, calibration constitutes the final step of the process, where the single fused Isig 3589 is calibrated to generate a single sensor glucose value 3598. See also FIG. 33B. Second, whereas SG fusion uses the SG values for the plurality of electrodes to calculate the final SG value 3498, the fused Isig value 3589 is calculated using the filtered Isigs (3522, 3524, and so on) for the plurality of electrodes.
  • In one closed-loop study involving a non-diabetic population, it was found that the above-described fusion algorithms provided considerable improvements in the Mean Absolute Relative Difference (MARD) both on Day 1, when low start-up issues are most significant and, as such, may have a substantial impact on sensor accuracy and reliability, and overall (i.e., over a 7-day life of the sensor). The study evaluated data for an 88% distributed layout design with high current density (nominal) plating using three different methodologies: (1) calculation of one sensor glucose value (SG) via fusion using Medtronic Minimed's Ferrari Algorithm 1.0 (which is a SG fusion algorithm as discussed above); (2) calculation of one SG by identifying the better ISIG value using 1 kHz EIS data (through the Isig fusion algorithm discussed above); and (3) calculation of one SG by using the higher ISIG value (i.e., without using EIS). The details of the data for the study are presented below:
  • (1) SG based on Ferrari 1.0 Alg for 88% distributed layout
    with high current density (nominal) plating
    Day
    1 2 3 4 5 6 7 Total
    Mean-ARD Percentage
    040-080 19.39 17.06 22.27 17.50 37.57 11.43 19.69
    080-120 19.69 09.18 09.34 08.64 10.01 08.31 11.33 11.56
    120-240 19.01 17.46 12.44 07.97 11.75 08.82 12.15 12.92
    240-400 10.25 08.36 14.09 10.86 12.84 22.70 12.88
    Total 19.52 11.71 10.14 09.30 10.83 09.49 11.89 12.28
    Mean-Absolute Bias (sg-bg)
    040-080 14.86 11.78 15.81 11.07 29.00 07.26 14.05
    080-120 19.53 09.37 09.49 08.78 09.88 08.44 11.61 11.62
    120-240 30.04 29.73 19.34 14.45 18.25 12.66 18.89 20.60
    240-400 26.75 22.23 39.82 29.00 33.00 61.36 35.19
    Total 21.62 15.20 12.79 13.21 12.04 10.84 15.04 14.79
    Mean-Signed Bias (sg-bg)
    040-080 12.15 09.78 15.81 11.07 29.00 07.26 13.01
    080-120 −04.45 −04.92 −00.90 00.18 01.21 00.85 00.03 −01.44
    120-240 −10.18 −27.00 −16.89 −02.91 −05.40 −01.24 −11.58 −10.71
    240-400 11.25 02.23 −00.07 −27.00 −33.00 −61.36 −10.29
    Total −04.81 −09.77 −05.09 −00.23 −00.22 00.67 −04.98 −03.56
    Eval Points
    040-080 007 004 000 002 006 003 004 026
    080-120 090 064 055 055 067 056 047 434
    120-240 028 025 022 021 016 032 026 170
    240-400 000 002 004 008 003 001 002 020
    Total 125 095 081 086 092 092 079 650
  • (2) SG based on better ISIG using 1 kHz EIS for 88% distributed
    layout with high current density (nominal) plating
    Day
    1 2 3 4 5 6 7 Total
    Mean-ARD Percentage
    040-080 16.66 18.78 21.13 16.21 43.68 09.50 18.14
    080-120 16.22 11.96 08.79 10.49 09.75 08.04 10.34 11.36
    120-240 15.08 17.50 12.68 07.72 08.74 08.84 13.02 12.16
    240-400 07.66 06.42 11.10 07.52 15.95 21.13 09.84
    Total 15.96 13.70 09.92 09.95 09.96 09.40 11.31 11.83
    Mean-Absolute Bias (sg-bg)
    040-080 12.71 13.00 15.00 10.17 33.50 06.00 12.83
    080-120 15.70 12.17 08.57 10.89 09.62 08.26 10.49 11.32
    120-240 24.43 29.82 19.43 13.79 14.60 12.97 20.27 19.58
    240-400 20.00 17.00 32.50 20.00 41.00 60.00 27.29
    Total 17.72 17.20 12.56 13.55 10.95 11.21 14.12 14.20
    Mean-Signed Bias (sg-bg)
    040-080 08.71 13.00 15.00 10.17 33.50 06.00 11.67
    080-120 −04.30 −08.62 −01.11 −03.64 02.52 00.40 −01.56 −02.52
    120-240 −11.30 −29.64 −17.09 −08.74 −10.87 −07.23 −15.09 −14.05
    240-400 20.00 00.50 09.50 −17.33 −41.00 −60.00 −03.18
    Total −05.30 −12.56 −06.20 −03.63 −00.10 −02.29 −06.35 −05.21
    Eval Points
    040-080 007 004 000 001 006 002 004 024
    080-120 082 053 044 045 058 043 041 366
    120-240 030 022 023 019 015 030 022 161
    240-400 000 002 004 006 003 001 001 017
    Total 119 081 071 071 082 076 068 568