U.S. patent application number 10/292780 was filed with the patent office on 2003-05-29 for blood analyte monitoring through subcutaneous measurement.
This patent application is currently assigned to TheraSense Inc.. Invention is credited to Bonnecaze, Roger T., Freeland, Angela C..
Application Number | 20030100040 10/292780 |
Document ID | / |
Family ID | 26748059 |
Filed Date | 2003-05-29 |
United States Patent
Application |
20030100040 |
Kind Code |
A1 |
Bonnecaze, Roger T. ; et
al. |
May 29, 2003 |
Blood analyte monitoring through subcutaneous measurement
Abstract
A method for obtaining an estimate of an analyte concentration
in a first fluid from an analyte concentration in a second fluid is
disclosed. The method includes obtaining measurements of an analyte
concentration in a second fluid by using a sensing device. The
analyte concentration estimate in the first fluid is determined
from these measurements. Also disclosed is a sensing device for
obtaining the measurements of an analyte concentration in the first
fluid and a processor configured and arranged to determine the
analyte concentration in the first body fluid according to this
method. This method and device can be used, for example, to
determine blood glucose concentration from measurements of the
glucose concentration in subcutaneous tissue. These measurements
may be made using in vitro or in vivo samples. In some instances, a
subcutaneously implanted sensing device, such as electrochemical
sensor, is used to make the measurements.
Inventors: |
Bonnecaze, Roger T.;
(Austin, TX) ; Freeland, Angela C.; (Austin,
TX) |
Correspondence
Address: |
Attention of Mara E. Liepa
MERCHANT & GOULD P.C.
P.O. Box 2903
Minneapolis
MN
55402-0903
US
|
Assignee: |
TheraSense Inc.
Alameda
CA
|
Family ID: |
26748059 |
Appl. No.: |
10/292780 |
Filed: |
November 11, 2002 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
|
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10292780 |
Nov 11, 2002 |
|
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09530938 |
Jul 24, 2000 |
|
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09530938 |
Jul 24, 2000 |
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PCT/US98/25685 |
Dec 4, 1998 |
|
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60067601 |
Dec 5, 1997 |
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60067603 |
Dec 5, 1997 |
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Current U.S.
Class: |
435/14 ;
702/19 |
Current CPC
Class: |
A61B 5/14514 20130101;
A61B 5/0031 20130101; A61B 5/4839 20130101; Y10S 435/817 20130101;
A61B 5/14532 20130101; A61B 5/746 20130101; A61B 5/7203 20130101;
A61B 5/1486 20130101; A61B 5/145 20130101 |
Class at
Publication: |
435/14 ;
702/19 |
International
Class: |
C12Q 001/54; G06F
019/00; G01N 033/48; G01N 033/50 |
Claims
We claim:
1. A method for determining analyte concentration in blood,
comprising: determining a subcutaneous analyte concentration from a
subcutaneous region using a sensing device; determining an analyte
concentration in blood from the subcutaneous analyte concentration
based on mass transfer of the analyte from blood to the
subcutaneous region and on uptake of the analyte by subcutaneous
cells surrounding the sensing region.
2. The method of claim 1, wherein the sensing device comprises an
electrochemical sensor having a working electrode.
3. The method of claim 2, further comprising subcutaneously
implanting the working electrode to generate a signal related to
the subcutaneous analyte concentration.
4. The method of claim 1, wherein the analyte is glucose.
5. The method of claim 4, wherein the blood glucose concentration
is determined from the subcutaneous glucose concentration using the
relationship: 22 V S t = k m A ( B - S ) - Vk r S K m + S ( 1 )
where V is the volume of the sensor, S is the subcutaneous glucose
concentration, B is the blood glucose concentration, A is the
surface area of the region surrounding the sensor, k.sub.m is a
mass transfer coefficient, K.sub.m is a Michaelis-Menten constant,
and k.sub.r' is the reaction rate constant for uptake of glucose by
the subcutaneous tissue.
6. The method of claim 5, wherein the subcutaneous glucose
concentration is determined from the blood glucose concentration
according to the relationship: 23 S ( t ) = S ( ) - ( i - ) + t B (
) - ( i - ) . where 24 = [ 1 + k r K m B 0 + S ] ,B.sub.0 is a
normalizing constant, .theta. is the initial time and t is the
final time for a window of computation.
7. The method of claim 6, wherein the subcutaneous glucose
concentration is determined from the blood glucose concentration
using: 25 t B ( ) - ( i - ) = i = 1 N B ( t i ) - ( i - t i ) t
.times. W i where W.sub.i is a weighting factor.
8. The method of claim 6, wherein the subcutaneous glucose
concentration is determined from the blood glucose concentration by
minimizing: f[b]=.chi..sup.2[b]+.lambda..PSI.[b], where 26 2 [ b ]
= = i i + N ( t B ( ) - i ( t - ) - S ( t ) + S ( ) - i ( t - ) ) 2
. and .PSI.[b] is a smoothness function.
9. The method of claim 8, wherein .PSI.[b] is selected to provide
first-order regularization.
10. The method of claim 8, wherein .PSI.[b] is selected to provide
second-order regularization.
11. An analyte measurement system comprising: a sensing device; and
a processor coupled to the sensing device and configured and
arranged to determine an analyte level in blood, from signals
generated by the sensing device, based on mass transfer of the
analyte from blood to a subcutaneous region and on uptake of the
analyte by subcutaneous cells surrounding the subcutaneous
region.
12. The analyte measurement system of claim 11, wherein the sensing
device comprises an electrochemical sensor having a working
electrode.
13. The analyte measurement system of claim 12, wherein the working
electrode is adapted for subcutaneous implantation in an
animal.
14. The analyte measurement system of claim 11, wherein the analyte
is glucose.
15. The analyte measurement system of claim 14, wherein the
processor is configured and arranged to determine the blood glucose
concentration from a subcutaneous glucose concentration using the
relationship: 27 V S t = k m A ( B - S ) - Vk r S K m + S where V
is the volume of the sensor, S is the subcutaneous glucose
concentration, B is the blood glucose concentration, A is the
surface area of the region surrounding the sensor, k.sub.m is a
mass transfer coefficient, K.sub.m is a Michaelis-Menten constant,
and k.sub.r' is the reaction rate constant for uptake of glucose by
the subcutaneous tissue.
16. The analyte measurement system of claim 15, wherein the
processor is configured and arranged to determine the blood glucose
concentration from a subcutaneous glucose concentration using the
relationship: 28 S ( t ) = S ( ) - ( i - ) t B ( ) - ( i - ) .
where 29 = [ 1 + k r K m B 0 + S ] ,B.sub.0 is a normalizing
constant, .theta. is the initial time and t is the final time for a
window of computation.
17. The analyte measurement system of claim 16, wherein the
processor is configured and arranged to determine the blood glucose
concentration from a subcutaneous glucose concentration using the
relationship: 30 t B ( ) - ( i - ) = i = 1 N B ( t i ) - ( i - t i
) t .times. W i where W.sub.i is a weighting factor.
18. The analyte measurement system of claim 16, wherein the
processor is configured and arranged to determine the blood glucose
concentration from a subcutaneous glucose concentration by
minimizing: f[b]=.chi..sup.2[b]+.lambda..PSI.[b], where 31 2 [ b ]
= = i i + N ( t B ( ) - i ( t - ) - S ( t ) + S ( ) - i ( t - ) ) 2
. and .PSI.[b] is a smoothness function.
19. The analyte measurement system of claim 18, wherein .PSI.[b] is
selected to provide first-order regularization.
20. The analyte measurement system of claim 18, wherein .PSI.[b] is
selected to provide second-order regularization.
21. The analyte measurement system of claim 11, further comprising
a display coupled to the processor for displaying the analyte
concentration in the blood.
22. The analyte measurement system of claim 11, further comprising
an alarm coupled to the processor for alerting a user based on the
analyte concentration.
23. The analyte measurement system of claim 11, wherein the
processor is disposed in a housing adapted for placement on the
skin of an animal.
24. The analyte measurement system of claim 23, further comprising
a transmitter coupled to the electrochemical sensor and a receiver
coupled to the processor, wherein the processor and receiver are
disposed in a housing adapted for remote reception of signals from
the electrochemical sensor via the transmitter.
25. An apparatus for determining analyte concentration in blood
based on measurements of analyte concentration determined using a
sensing device, comprising: a processor configured and arranged to
determine analyte concentration in the blood from a measured
subcutaneous analyte concentration based on mass transfer of the
analyte from blood to a subcutaneous region and on uptake of the
analyte by subcutaneous cells surrounding the subcutaneous
region.
26. The apparatus of claim 25, wherein the analyte is glucose.
27. A method for determining analyte-concentration in a first body
fluid, comprising: obtaining measurements of an analyte
concentration in a second body fluid, different from the first body
fluid, from a sensing device; and determining an analyte
concentration estimate in the first fluid from the measurements by
minimizing the relation: f[b]=.chi..sup.2[b]+.lambda..PSI.[b],
wherein b is a vector representing analyte concentration in the
first body fluid, .chi..sup.2[b] is a function representing a fit
between the estimates and the measurements, .lambda. is a weighting
function, and .PSI.[b] is a function indicating smoothness of the
analyte concentration estimates in the first body fluid.
28. The method of claim 27, wherein the analyte is glucose.
29. The method of claim 28, wherein the first body fluid is blood
and the second body fluid is subcutaneous fluid.
30. The method of claim 28, wherein the model is based on the
relationship: 32 V S t = k m A ( B - S ) - Vk r S K m + S where V
is the volume of the sensor, S is the subcutaneous glucose
concentration, B is the blood glucose concentration, A is the
surface area of the region surrounding the sensor, k.sub.m is a
mass transfer coefficient, K.sub.m is a Michaelis-Menten constant,
and k.sub.r' is the reaction rate constant for uptake of glucose by
the subcutaneous tissue.
31. The method of claim 27, wherein .PSI.[b] is selected to provide
first-order regularization.
32. The method of claim 27, wherein .PSI.[b] is selected to
provides second-order regularization.
33. The method of claim 29, wherein obtaining subcutaneous glucose
measurements comprises subcutaneously implanting a working
electrode of a glucose sensor into an animal; and determining a
subcutaneous glucose concentration from a signal generated at the
working electrode.
34. An analyte measurement system, comprising: a sensing device;
and a processor, coupled to the sensing device, that is configured
and arranged to determine an analyte concentration estimate in a
first fluid from measurements of analyte concentration in a second
body fluid by minimizing the relation:
f[b]=.chi..sup.2[b]+.PSI.[b], wherein b is a vector representing
analyte concentration in the first body fluid, .chi..sup.2[b] is a
function representing a fit between the estimates and the
measurements, .lambda. is a weighting function, and .PSI.[b] is a
function indicating smoothness of the analyte concentration
estimates in the first body fluid.
35. The analyte measurement system of claim 34, wherein the analyte
is glucose.
36. The analyte measurement system of claim 35, wherein the first
body fluid is blood and the second body fluid is subcutaneous
fluid.
37. The analyte measurement system of claim 36, wherein the sensing
device comprises a working electrode adapted for subcutaneous
implantation into an animal.
Description
[0001] This application is a continuation of Ser. No. 09/530,938,
filed Jul. 24, 2000, which claims priority to PCT/US98/25685 having
an international filing date of Dec. 4, 1998, which in turn claims
priority to Serial Nos. 60/067,603 and 60/067,601, both filed Dec.
5, 1997.
[0002] The present invention is, in general, directed to devices
and methods for the monitoring of the concentration of an analyte,
such as glucose, using a subcutaneous sensor. More particularly,
the present invention relates to devices and methods for the
monitoring of an analyte using a subcutaneous electrochemical
sensor to provide information to a patient about the level of the
analyte in blood.
BACKGROUND OF THE INVENTION
[0003] The monitoring of the level of analytes, such as glucose,
lactate or oxygen, in certain individuals is vitally important to
their health. High or low levels of these analytes may have
detrimental effects. For example, the monitoring of glucose is
particularly important to individuals with diabetes, as they must
determine when insulin is needed to reduce glucose levels in their
bodies or when additional glucose is needed to raise the level of
glucose.
[0004] A variety of methods have been used to measure analyte
concentrations. For example, colorimetric, electrochemical, and
optical methods have been developed for the determination of blood
glucose concentration. Implanted electrochemical sensors may be
used to periodically or continuously monitor glucose (or other
analyte) concentration. Although sensors accurately measure the
glucose concentration when inserted directly into the bloodstream,
infection may occur at this implantation site.
[0005] A variety of sensors have been developed for implantation in
subcutaneous tissue to measure the subcutaneous glucose
concentration, which is thought to be well correlated with the
blood glucose concentration at steady-state. Subcutaneously
implanted glucose sensors, such as miniaturized electrodes "wired"
to glucose oxidase, are one technology that hold promise for
continuous monitoring of blood glucose levels by diabetic patients.
These sensors measure subcutaneous glucose concentrations as
glucose diffuses from the blood into the subcutaneous tissue and
then to the enzyme electrode surface. At this surface, the glucose
is oxidized and the reaction causes electrons to be transferred to
the electrode surface. The resulting current is proportional to the
concentration of glucose in the region of implantation.
[0006] In many cases, it is important to be able to convert a value
from a subcutaneous concentration to a blood concentration. For
example, a subcutaneous sensor may be calibrated using blood
measurements or a diagnosis or method of treatment may depend on
the knowledge of the blood analyte concentration that is obtained
using a subcutaneous sensor. However, a lag typically results
between the blood and subcutaneous glucose concentrations as the
blood glucose level increases or decreases. In addition, the
subcutaneous analyte concentrations obtained from sensor
measurements may be different from the blood analyte concentration
because of the existence of a mass transfer barrier. Thus, there is
a need to develop devices and methods that can convert subcutaneous
analyte measurements to blood analyte concentrations to ensure
accuracy, compatibility, and comparability between measurements
made by subcutaneous electrochemical sensors and those made using
other conventional blood analysis techniques.
SUMMARY OF THE INVENTION
[0007] Generally, the present invention relates to methods and
devices for determination of analyte concentration in one body
fluid using analyte concentration measurements from a second body
fluid. In particular, the present invention includes methods and
devices for the determination of blood glucose concentration using
glucose concentration measurements from subcutaneous fluids.
[0008] One embodiment of the invention is a method for obtaining an
estimate of an analyte concentration in a first fluid. First,
measurements of an analyte concentration in a second fluid are
obtained using a sensing device. An analyte concentration estimate
in the first body fluid is determined from these measurements by
minimizing the relation:
f[b]=.chi..sup.2[b]+.lambda..PSI.[b],
[0009] where b is a vector representing analyte concentration in
the first fluid, .chi..sup.2[b] is a function representing a fit
between the estimates and the measurements, .lambda. is a weighting
function and .PSI.[b] is a function indicating smoothness of the
analyte concentration estimates in the first body fluid. Another
embodiment includes a sensing device for obtaining the measurements
of analyte concentration in the first fluid and a processor
configured and arranged to determine the analyte concentration
according to this method.
[0010] This method and device can be used, for example, to
determine blood glucose concentration from measurements of the
glucose concentration in subcutaneous tissue. These measurements
may be made using in vitro or in vivo samples. In some instances, a
subcutaneously implanted sensing device such as an electrochemical
sensor, is used to make the measurements.
[0011] Another embodiment is a method of determining blood analyte
concentration including obtaining a subcutaneous analyte
concentration from a subcutaneous region using a sensing device and
determining a blood analyte concentration from the subcutaneous
analyte concentration based on a) mass transfer of the analyte from
blood to the subcutaneous region and b) uptake of the analyte by
subcutaneous cells in the subcutaneous region. Examples of analytes
include glucose, lactate, and oxygen. Yet another embodiment is an
analyte measurement device including a processor configured and
arranged to determine the analyte concentration according to this
method and an optional sensing device, such as an electrochemical
sensor, for obtaining the measurements of analyte concentration in
the first fluid. In some instances, the electrochemical sensor may
be subcutaneously implanted and the analyte measurement device may
periodically or continuously monitor glucose.
[0012] The above summary of the present invention is not intended
to describe each disclosed embodiment or every implementation of
the present invention. The Figures and the detailed description
which follow more particularly exemplify these embodiments.
BRIEF DESCRIPTION OF THE DRAWINGS
[0013] The invention may be more completely understood in
consideration of the following detailed description of various
embodiments of the invention in connection with the accompanying
drawings, in which:
[0014] FIG. 1 is a block diagram of one embodiment of a
subcutaneous analyte monitor using a subcutaneously implantable
analyte sensor, according to the invention;
[0015] FIG. 2 is a top view of one embodiment of an analyte sensor,
according to the invention;
[0016] FIG. 3A is a cross-sectional view of the analyte sensor of
FIG. 2;
[0017] FIG. 3B is a cross-sectional view of another embodiment of
an analyte sensor, according to the invention:
[0018] FIG. 4A is a cross-sectional view of a third embodiment of
an analyte sensor, according to the invention;
[0019] FIG. 4B is a cross-sectional view of a fourth embodiment of
an analyte sensor, according to the invention;
[0020] FIG. 5 is an expanded top view of a tip portion of the
analyte sensor of FIG. 2;
[0021] FIG. 6 is a cross-sectional view of a fifth embodiment of an
analyte sensor, according to the invention;
[0022] FIG. 7 is an expanded top view of a tip-portion of the
analyte sensor of FIG. 6;
[0023] FIG. 8 is an expanded bottom view of a tip-portion of the
analyte sensor of FIG. 6;
[0024] FIG. 9 is a side view of the analyte sensor of FIG. 2;
[0025] FIG. 10 is a top view of the analyte sensor of FIG. 6;
[0026] FIG. 11 is a bottom view of the analyte sensor of FIG.
6;
[0027] FIG. 12 is an expanded side view of one embodiment of a
sensor and an insertion device, according to the invention;
[0028] FIGS. 13A, 13B, 13C are cross-sectional views of three
embodiments of the insertion device of FIG. 12;
[0029] FIG. 14 is a cross-sectional view of one embodiment of a
on-skin sensor control unit, according to the invention;
[0030] FIG. 15 is a top view of a base of the on-skin sensor
control unit of FIG. 14;
[0031] FIG. 16 is a bottom view of a cover of the on-skin sensor
control unit of FIG. 14;
[0032] FIG. 17 is a perspective view of the on-skin sensor control
unit of FIG. 14 on the skin of a patient;
[0033] FIG. 18A is a block diagram of one embodiment of an on-skin
sensor control unit, according to the invention;
[0034] FIG. 18B is a block diagram of another embodiment of an
on-skin sensor control unit, according to the invention;
[0035] FIGS. 19A, 19B, 19C, and 19D are cross-sectional views of
four embodiments of conductive contacts disposed on an interior
surface of a housing of an on-skin sensor control unit, according
to the invention;
[0036] FIGS. 19E and 19F are cross-sectional views of two
embodiments of conductive contacts disposed on an exterior surface
of a housing of an on-skin sensor control unit, according to the
invention;
[0037] FIGS. 20A and 20B are schematic diagrams of two embodiments
of a current-to-voltage converter for use in an analyte monitoring
device, according to the invention;
[0038] FIG. 21 is a block diagram of one embodiment of an open loop
modulation system for use in an analyte monitoring device,
according to the invention;
[0039] FIG. 22 is a block diagram of one embodiment of a
receiver/display unit, according to the invention;
[0040] FIG. 23 is a front view of one embodiment of a
receiver/display unit;
[0041] FIG. 24 is a front view of a second embodiment of a
receiver/display unit; FIG. 25 is a block diagram of one embodiment
of a drug delivery system, according to the invention;
[0042] FIG. 26 is a perspective view of the internal structure of
an insertion gun, according to the invention;
[0043] FIG. 27A is a top view of one embodiment of an on-skin
sensor control unit, according to the invention;
[0044] FIG. 27B is a top view of one embodiment of a mounting unit
of the on-skin sensor control unit of FIG. 27A;
[0045] FIG. 28A is a top view of another embodiment of an on-skin
sensor control unit after insertion of an insertion device and a
sensor, according to the invention;
[0046] FIG. 28B is a top view of one embodiment of a mounting unit
of the on-skin sensor control unit of FIG. 28A;
[0047] FIG. 28C is a top view of one embodiment of a housing for at
least a portion of the electronics of the on-skin sensor control
unit of FIG. 28A;
[0048] FIG. 28D is a bottom view of the housing of FIG. 28C;
[0049] FIG. 28E is a top view of the on-skin sensor control unit of
FIG. 28A with a cover of the housing removed;
[0050] FIG. 29 is another embodiment of an analyte sensor;
[0051] FIG. 30 is a graph of experimental data (smooth line) from a
rat during an intravenous insulin injection and a prediction using
an inverse model with no regularization (oscillating line);
[0052] FIG. 31 is a graph of simulated blood glucose response
(solid line) and three models used to simulate subcutaneous glucose
response including a) k.sub.r=0, b) k.sub.r=1, K.sub.m=B.sub.o/3,
and c) k.sub.r=1, K.sub.m=B.sub.o;
[0053] FIG. 32 is a graph of simulated subcutaneous glucose
response with white noise (dotted line) and time-correlated noise
(solid line) at a noise level of 1%;
[0054] FIG. 33 is a graph of first- and second-order regularization
for a solution of blood glucose concentration based on simulated
subcutaneous glucose concentration;
[0055] FIG. 34 is a graph of error magnification factor versus
weighting factor for zeroeth-, first-, and second-order
regularization;
[0056] FIG. 35 is a graph of error magnification factor versus
weighting factor for zeroeth-, first-, and second-order
regularization, varying values of window size and data sampling
time;
[0057] FIG. 36 is a graph of magnification factor versus weighting
factor for k.sub.r=0 and N.DELTA.t=1.481;
[0058] FIG. 37 is a graph of magnification factor versus weighting
factor including white noise or time-correlated noise in the
simulated subcutaneous glucose concentration;
[0059] FIG. 38 is a graph of squared model error versus weight
factor;
[0060] FIG. 39 is a graph of magnification factor versus weighting
factor for a) k.sub.r=0 and b) k.sub.r=1, K.sub.m=B.sub.0;
[0061] FIG. 40 is a graph of glucose concentration illustrating a
decline in concentration of glucose after intravenous injection of
insulin;
[0062] FIG. 41 is a graph of estimated glucose concentration of a
subcutaneously implanted sensor (dotted line), an intravascularly
implanted sensor (solid line), and venous blood glucose
concentration (circles) after an i.v. bolus of insulin.
[0063] FIG. 42 is a graph of average difference (n=7) of
subcutaneous glucose estimates relative to actual blood glucose
measurements, % difference=100.times.(subcutaneous estimate-blood
measurement)/(blood measurement);
[0064] FIG. 43 is a graph of subcutaneous glucose concentration
predicted using a forward model (dotted lines) based on data from a
jugular sensor and measured subcutaneous glucose concentration
(solid line); and
[0065] FIG. 44 includes graphs for seven rats comparing blood
glucose concentration as determined by a sensor (solid line) and
predicted by an inverse model with regularization (dashed
line).
[0066] While the invention is amenable to various modifications and
alternative forms, specifics thereof have been shown by way of
example in the drawings and will be described in detail. It should
be understood, however, that the intention is not to limit the
invention to the particular embodiments described. On the contrary,
the intention is to cover all modifications, equivalents, and
alternatives falling within the spirit and scope of the invention
as defined by the appended claims.
DETAILED DESCRIPTION OF THE INVENTION
[0067] The present invention is applicable to a method and analyte
measurement systems for determining analyte concentration in one
body fluid (e.g., blood) from measured analyte concentrations in
another body fluid (e.g., subcutaneous fluid).
[0068] Suitable analyte measurement systems typically include a
sensing device and a processor. The analyte measurement system may
be configured and arranged to provide readings as required by a
user when, for-example, the user provides a sample to the device.
In other embodiments, the analyte measurement system may be
configured and arranged to be permanently or temporarily attached
to an animal (such as a human) to provide periodic or continuous
monitoring.
[0069] For example, the analyte measurement system can be an
analyte monitoring system using a subcutaneously implantable
electrochemical sensor for the in vivo determination of a blood
concentration of an analyte, such as, for example, glucose,
lactate, or oxygen. The sensor can be, for example, subcutaneously
implanted in a patient for the continuous or periodic monitoring of
the analyte. The analyte monitoring system typically includes a
subcutaneously implantable sensor and a processor coupled to the
sensor to determine the blood analyte concentration from the sensor
measurements.
[0070] A variety of suitable sensing devices are available. A
suitable sensing device is configured and arranged to provide some
signal, for example, an optical (e.g., color change, absorption,
transmission, or fluorescence) or electrical signal (e.g., a change
in current, potential, capacitance, or conductivity) that is
related to a level of the analyte in the sample. Suitable sensing
devices include electrochemical sensing devices, optical sensing
devices, and calorimetric sensing devices. A sample of a body fluid
may be provided, conveyed, or transported to the sensing device for
in vitro determination of the analyte concentration in the body
fluid. In other embodiments, the sensing device (e.g., an
electrochemical sensor) may be implanted to provide in vivo
determination of analyte concentration. In yet other embodiments,
the sensing device (e.g., an optical device) may be directed toward
the animal or a sample from the animal and the analyte
concentration determined by, for example, interaction of light with
the tissue and/or body fluid of the animal.
[0071] The determination of blood glucose concentration from
subcutaneous glucose measurements is used herein as an
illustration. It will be understood that other analytes may also be
measured. It will also be understood that the devices and methods
described herein can be applied to the determination of analyte
concentration in body fluids, other than blood, based on
measurement of analyte concentration in a second body fluid.
[0072] The following definitions are provided for terms used
herein:
[0073] A "counter electrode" refers to an electrode paired with the
working electrode, through which passes a current equal in
magnitude and opposite in sign to the current passing through the
working electrode. In the context of the invention, the term
"counter electrode" is meant to include counter electrodes which
also function as reference electrodes (i.e., a counter/reference
electrode).
[0074] An "electrochemical sensor" is a device configured to detect
the presence and/or measure the level of an analyte in a sample via
electrochemical oxidation and reduction reactions on the sensor.
These reactions are transduced to an electrical signal that can be
correlated to an amount, concentration, or level of an analyte in
the sample.
[0075] "Electrolvsis" is the electrooxidation or electroreduction
of a compound either directly at an electrode or via one or more
electron transfer agents.
[0076] A compound is "immobilized" on a surface when it is
entrapped on or chemically bound to the surface.
[0077] A "non-leachable" or "non-releasable" compound or a compound
that is "non-leachably disposed" is meant to define a compound that
is affixed on the sensor such that it does not substantially
diffuse away from the working surface of the working electrode for
the period in which the sensor is used (e.g., the period in which
the sensor is implanted in a patient or measuring a sample).
[0078] Components are "immobilized" within a sensor, for example,
when the components are covalently, ionically, or coordinatively
bound to constituents of the sensor and/or are entrapped in a
polymeric or sol-gel matrix or membrane which precludes
mobility.
[0079] An "electron transfer agent" is a compound that carries
electrons between the analyte and the working electrode, either
directly, or in cooperation with other electron transfer agents.
One example of an electron transfer agent is a redox mediator.
[0080] A "working electrode" is an electrode at which the analyte
(or a second compound whose level depends on the level of the
analyte) is electrooxidized or electroreduced with or without the
agency of an electron transfer agent.
[0081] A "working surface" is that portion of the working electrode
which is coated with or is accessible to the electron transfer
agent and configured for exposure to an analyte-containing
fluid.
[0082] A "sensing layer" is a component of the sensor which
includes constituents that facilitate the electrolysis of the
analyte. The sensing layer may include constituents such as an
electron transfer agent, a catalyst which catalyzes a reaction of
the analyte to produce a response at the electrode, or both. In
some embodiments of the sensor, the sensing layer is non-leachably
disposed in proximity to or on the working electrode.
[0083] A "non-corroding" conductive material includes non-metallic
materials, such as carbon and conductive polymers.
[0084] Sensing Devices
[0085] The methods and devices of the invention are illustrated
using electrochemical sensors. However, it will be understood that
a variety of sensing devices, including electrochemical, optical,
and colorimetric sensing devices may be used. Moreover, the methods
and devices are illustrated using implantable sensing devices,
however, it will be understood that other non-implantable sensing
devices can be used.
[0086] A variety of subcutaneously implantable sensors are
available for use. Examples of such sensors and analyte measurement
systems incorporating the sensors are described in U.S. Pat. No.
5,593,852 and U.S. patent application Ser. Nos. 09/034,372,
09/034,422, and 09/070,677, all of which are incorporated herein by
reference. An example of one sensor is illustrated in FIG. 29 and
described in detail in U.S. Pat. No. 5,593,852. This sensor
includes a metal or carbon working electrode 2 with an electrically
insulating material 4 wrapped around the electrode. A recess 6 is
provided by, for example, removing a portion of the working
electrode 2. This leaves an exposed surface 18 of the working
electrode.
[0087] A sensing layer 8 is formed over the exposed surface 18. The
sensing layer 8 may include a redox mediator and/or a redox enzyme.
In at least some embodiments, the redox mediator and/or redox
enzyme are non-leachably disposed in the sensor, as described in
U.S. Pat. No. 5,593,852. Exemplary redox mediators and redox
enzymes are described in U.S. Pat. No. 5,593,852 and U.S. patent
application Ser. Nos. 09/034,372. 09/034,422, and 09/070,677.
[0088] An optional glucose diffusion limiting layer 10, an optional
interferent eliminating layer 12, and an optional biocompatible
layer 14 can be formed in the recess 6. These layers are described
in more detail in U.S. Pat. No. 5,593,852. Another Sensor and an
Analyte Monitoring System The analyte monitoring systems of the
present invention can be utilized under a variety of conditions.
The particular configuration of a sensor and other units used in
the analyte monitoring system may depend on the use for which the
analyte monitoring system is intended and the conditions under
which the analyte monitoring system will operate. One embodiment of
the analyte monitoring system includes a sensor configured for
implantation into a patient or user. For example, implantation of
the sensor may be made in the arterial or venous systems for direct
testing of analyte levels in blood. Alternatively, a sensor may be
implanted in the interstitial tissue for determining the analyte
level in interstitial fluid. This level may be correlated and/or
converted to analyte levels in blood or other fluids. The site and
depth of implantation may affect the particular shape, components,
and configuration of the sensor. Subcutaneous implantation may be
preferred, in some cases, to limit the depth of implantation of the
sensor. Sensors may also be implanted in other regions of the body
to determine analyte levels in other fluids. Examples of suitable
sensor for use in the analyte monitoring systems of the invention
are described in U.S. patent application Ser. No. 09/034,372,
incorporated herein by reference.
[0089] One embodiment of the analyte monitoring system 40 for use
with an implantable sensor 42, and particularly for use with a
subcutaneously implantable sensor, is illustrated in block diagram
form in FIG. 1. The analyte monitoring system 40 includes, at
minimum, a sensor 42, a portion of which is configured for
implantation (e.g., subcutaneous, venous, or arterial implantation)
into a patient, and a sensor control unit 44. The sensor 42 is
coupled to the sensor control unit 44 which is typically attached
to the skin of a patient. The sensor control unit 44 operates the
sensor 42, including, for example, providing a voltage across the
electrodes of the sensor 42 and collecting signals from the sensor
42. The sensor control unit 44 may evaluate the signals from the
sensor 42 and/or transmit the signals to one or more optional
receiver/display units 46, 48 for evaluation. The sensor control
unit 44 and/or the receiver/display units 46, 48 may display or
otherwise communicate the current level of the analyte.
Furthermore, the sensor control unit 44 and/or the receiver/display
units 46, 48 may indicate to the patient, via, for example, an
audible, visual, or other sensory-stimulating alarm, when the level
of the analyte is at or near a threshold level. In some
embodiments, a electrical shock can be delivered to the patient as
a warning through one of the electrodes or the optional temperature
probe of the sensor. For example, if glucose is monitored then an
alarm may be used to alert the patient to a hypoglycemic or
hyperglycemic glucose level and/or to impending hypoglycemia or
hyperglycemia.
[0090] The Sensor
[0091] A sensor 42 includes at least one working electrode 58
formed on a substrate 50, as shown in FIG. 2. The sensor 42 may
also include at least one counter electrode 60 (or
counter/reference electrode) and/or at least one reference
electrode 62 (see FIG. 8). The counter electrode 60 and/or
reference electrode 62 may be formed on the substrate 50 or may be
separate units. For example, the counter electrode and/or reference
electrode may be formed on a second substrate which is also
implanted in the patient or, for some embodiments of the
implantable sensors, the counter electrode and/or reference
electrode may be placed on the skin of the patient with the working
electrode or electrodes being implanted into the patient. The use
of an on-the-skin counter and/or reference electrode with an
implantable working electrode is described in U.S. Pat. No.
5,593,852, incorporated herein by reference.
[0092] The working electrode or electrodes 58 are formed using
conductive traces 52 disposed on the substrate 50. The counter
electrode 60 and/or reference electrode 62, as well as other
optional portions of the sensor 42, such as a temperature probe 66
(see FIG. 8), may also be formed using conductive traces 52
disposed on the substrate 50. These conductive traces 52 may be
formed over a smooth surface of the substrate 50 or within channels
54 formed by, for example, embossing, indenting or otherwise
creating a depression in the substrate 50. A sensing layer 64 (see
FIGS. 3A and 3B) is often formed proximate to or on at least one of
the working electrodes 58 to facilitate the electrochemical
detection of the analyte and the determination of its level in the
sample fluid, particularly if the analyte can not be electrolyzed
at a desired rate and/or with a desired specificity on a bare
electrode. The sensing layer 64 may include an electron transfer
agent to transfer electrons directly or indirectly between the
analyte and the working electrode 58. The sensing layer 64 may also
contain a catalyst to catalyze a reaction of the analyte. The
components of the sensing layer may be in a fluid or gel that is
proximate to or in contact with the working electrode 58.
Alternatively, the components of the sensing layer 64 may be
disposed in a polymeric or sol-gel matrix that is proximate to or
on the working electrode 58. Preferably, the components of the
sensing layer 64 are non-leachably disposed within the sensor 42.
More preferably, the components of the sensor 42 are immobilized
within the sensor 42.
[0093] In addition to the electrodes 58, 60, 62 and the sensing
layer 64, the sensor 42 may also include a temperature probe 66
(see FIGS. 6 and 8), a mass transport limiting layer 74 (see FIG.
9), a biocompatible layer 75 (see FIG. 9), and/or other optional
components, as described below. Each of these items enhances the
functioning of and/or results from the sensor 42, as discussed
below.
[0094] The Substrate
[0095] The substrate 50 may be formed using a variety of
non-conducting materials, including, for example, polymeric or
plastic materials and ceramic materials. Suitable materials for a
particular sensor 42 may be determined, at least in part, based on
the desired use of the sensor 42 and properties of the
materials.
[0096] In some embodiments, the substrate is flexible. For example,
if the sensor 42 is configured for implantation into a patient,
then the sensor 42 may be made flexible (although rigid sensors may
also be used for implantable sensors) to reduce pain to the patient
and damage to the tissue caused by the implantation of and/or the
wearing of the sensor 42. A flexible substrate 50 often increases
the patient's comfort and allows a wider range of activities.
Suitable materials for a flexible substrate 50 include, for
example, non-conducting plastic or polymeric materials and other
non-conducting, flexible, deformable materials. Examples of useful
plastic or polymeric materials include thermoplastics such as
polycarbonates, polyesters (e.g., Mylar.TM. and polyethylene
terephthalate (PET)), polyvinyl chloride (PVC), polyurethanes,
polyethers, polyamides, polyimides, or copolymers of these
thermoplastics, such as PETG (glycol-modified polyethylene
terephthalate).
[0097] In other embodiments, the sensors 42 are made using a
relatively rigid substrate 50 to, for example, provide structural
support against bending or breaking. Examples of rigid materials
that may be used as the substrate 50 include poorly conducting
ceramics, such as aluminum oxide and silicon dioxide. One advantage
of an implantable sensor 42 having a rigid substrate is that the
sensor 42 may have a sharp point and/or a sharp edge to aid in
implantation of a sensor 42 without an additional insertion
device.
[0098] It will be appreciated that for many sensors 42 and sensor
applications, both rigid and flexible sensors will operate
adequately. The flexibility of the sensor 42 may also be controlled
and varied along a continuum by changing, for example, the
composition and/or thickness of the substrate 50.
[0099] In addition to considerations regarding flexibility, it is
often desirable that implantable sensors 42 should have a substrate
50 which is non-toxic. Preferably, the substrate 50 is approved by
one or more appropriate governmental agencies or private groups for
in vivo use.
[0100] The sensor 42 may include optional features to facilitate
insertion of an implantable sensor 42, as shown in FIG. 12. For
example, the sensor 42 may be pointed at the tip 123 to ease
insertion. In addition, the sensor 42 may include a barb 125 which
assists in anchoring the sensor 42 within the tissue of the patient
during operation of the sensor 42. However, the barb 125 is
typically small enough that little damage is caused to the
subcutaneous tissue when the sensor 42 is removed for
replacement.
[0101] Although the substrate 50 in at least some embodiments has
uniform dimensions along the entire length of the sensor 42, in
other embodiments, the substrate 50 has a distal end 67 and a
proximal end 65 with different widths 53, 55, respectively, as
illustrated in FIG. 2. In these embodiments, the distal end 67 of
the substrate 50 may have a relatively narrow width 53. For sensors
42 which are implantable into the subcutaneous tissue or another
portion of a patient's body, the narrow width 53 of the distal end
67 of the substrate 50 may facilitate the implantation of the
sensor 42. Often, the narrower the width of the sensor 42, the less
pain the patient will feel during implantation of the sensor and
afterwards.
[0102] For subcutaneously implantable sensors 42 which are designed
for continuous or periodic monitoring of the analyte during normal
activities of the patient, a distal end 67 of the sensor 42 which
is to be implanted into the patient has a width 53 of 2 mm or less,
preferably 1 mm or less, and more preferably 0.5 mm or less. If the
sensor 42 does not have regions of different widths, then the
sensor 42 will typically have an overall width of, for example. 2
mm. 1.5 mm, 1 mm. 0.5 mm. 0.25 mm, or less. However, wider or
narrower sensors may be used. In particular, wider implantable
sensors may be used for insertion into veins or arteries or when
the movement of the patient is limited, for example, when the
patient is confined in bed or in a hospital.
[0103] Returning to FIG. 2, the proximal end 65 of the sensor 42
may have a width 55 larger than the distal end 67 to facilitate the
connection between contact pads 49 of the electrodes and contacts
on a control unit. The wider the sensor 42 at this point, the
larger the contact pads 49 can be made. This may reduce the
precision needed to properly connect the sensor 42 to contacts on
the control unit (e.g., sensor control unit 44 of FIG. 1). However,
the maximum width of the sensor 42 may be constrained so that the
sensor 42 remains small for the convenience and comfort of the
patient and/or to fit the desired size of the analyte monitor. For
example, the proximal end 65 of a subcutaneously implantable sensor
42, such as the sensor 42 illustrated in FIG. 1, may have a width
55 ranging from 0.5 mm to 15 mm, preferably from 1 mm to 10 mm, and
more preferably from 3 mm to 7 mm. However, wider or narrower
sensors may be used in this and other in vivo applications.
[0104] The thickness of the substrate 50 may be determined by the
mechanical properties of the substrate material (e.g., the
strength, modulus, and/or flexibility of the material), the desired
use of the sensor 42 including stresses on the substrate 50 arising
from that use, as well as the depth of any channels or indentations
formed in the substrate 50, as discussed below. Typically, the
substrate 50 of a subcutaneously implantable sensor 42 for
continuous or periodic monitoring of the level of an analyte while
the patient engages in normal activities has a thickness of 50 to
500 .mu.m and preferably 100 to 300 .mu.m. However, thicker and
thinner substrates 50 may be used, particularly in other types of
in vivo sensors 42.
[0105] The length of the sensor 42 may have a wide range of values
depending on a variety of factors. Factors which influence the
length of an implantable sensor 42 may include the depth of
implantation into the patient and the ability of the patient to
manipulate a small flexible sensor 42 and make connections between
the sensor 42 and the sensor control unit 44. A subcutaneously
implantable sensor 42 for the analyte monitor illustrated in FIG. 1
may have a length ranging from 0.3 to 5 cm, however, longer or
shorter sensors may be used. The length of the narrow portion of
the sensor 42 (e.g., the portion which is subcutaneously inserted
into the patient), if the sensor 42 has narrow and wide portions,
is typically about 0.25 to 2 cm in length. However, longer and
shorter portions may be used. All or only a part of this narrow
portion may be subcutaneously implanted into the patient. The
lengths of other implantable sensors 42 will vary depending, at
least in part, on the portion of the patient into which the sensor
42 is to be implanted or inserted.
[0106] Conductive Traces
[0107] At least one conductive trace 52 is formed on the substrate
for use in constructing a working electrode 58. In addition, other
conductive traces 52 may be formed on the substrate 50 for use as
electrodes (e.g., additional working electrodes, as well as
counter, counter/reference, and/or reference electrodes) and other
components, such as a temperature probe. The conductive traces 52
may extend most of the distance along a length 57 of the sensor 50,
as illustrated in FIG. 2, although this is not necessary. The
placement of the conductive traces 52 may depend on the particular
configuration of the analyte monitoring system (e.g., the placement
of control unit contacts and/or the sample chamber in relation to
the sensor 42). For implantable sensors, particularly
subcutaneously implantable sensors, the conductive traces typically
extend close to the tip of the sensor 42 to minimize the amount of
the sensor that must be implanted.
[0108] The conductive traces 52 may be formed on the substrate 50
by a variety of techniques, including, for example,
photolithography, screen printing, or other impact or non-impact
printing techniques. The conductive traces 52 may also be formed by
carbonizing conductive traces 52 in an organic (e.g., polymeric or
plastic) substrate 50 using a laser. A description of some
exemplary methods for forming the sensor 42 is provided in U.S.
patent application Ser. No. 09/034,422, incorporated herein by
reference.
[0109] Another method for disposing the conductive traces 52 on the
substrate 50 includes the formation of recessed channels 54 in one
or more surfaces of the substrate 50 and the subsequent filling of
these recessed channels 54 with a conductive material 56, as shown
in FIG. 3A. The recessed channels 54 may be formed by indenting,
embossing, or otherwise creating a depression in the surface of the
substrate 50. Exemplary methods for forming channels and electrodes
in a surface of a substrate can be found in U.S. patent application
Ser. No. 09/034,422. The depth of the channels is typically related
to the thickness of the substrate 50. In one embodiment, the
channels have depths in the range of about 12.5 to 75 .mu.m (0.5 to
3 mils), and preferably about 25 to 50 .mu.m (1 to 2 mils).
[0110] The conductive traces are typically formed using a
conductive material 56 such as carbon (e.g., graphite), a
conductive polymer, a metal or alloy (e.g., gold or gold alloy), or
a metallic compound (e.g., ruthenium dioxide or titanium dioxide).
The formation of films of carbon, conductive polymer, metal, alloy,
or metallic compound are well-known and include, for example,
chemical vapor deposition (CVD), physical vapor deposition,
sputtering, reactive sputtering, printing, coating, and painting.
The conductive material 56 which fills the channels 54 is often
formed using a precursor material, such as a conductive ink or
paste. In these embodiments, the conductive material 56 is
deposited on the substrate 50 using methods such as coating,
painting, or applying the material using a spreading instrument,
such as a coating blade. Excess conductive material between the
channels 54 is then removed by, for example, running a blade along
the substrate surface.
[0111] In one embodiment, the conductive material 56 is a part of a
precursor material, such as a conductive ink, obtainable, for
example, from Ercon, Inc. (Wareham, Mass.), Metech, Inc. (Elverson,
Pa.), E. I. du Pont de Nemours and Co. (Wilmington, Del.),
Emca-Remex Products (Montgomeryville, Pa.), or MCA Services
(Melbourn, Great Britain). The conductive ink is typically applied
as a semiliquid or paste which contains particles of the carbon,
metal, alloy, or metallic compound and a solvent or dispersant.
After application of the conductive ink on the substrate 50 (e.g.,
in the channels 54), the solvent or dispersant evaporates to leave
behind a solid mass of conductive material 56.
[0112] In addition to the particles of carbon, metal, alloy, or
metallic compound, the conductive ink may also contain a binder.
The binder may optionally be cured to further bind the conductive
material 56 within the channel 54 and/or on the substrate 50.
Curing the binder increases the conductivity of the conductive
material 56. However, this is typically not necessary as the
currents carried by the conductive material 56 within the
conductive traces 52 are often relatively low (usually less than 1
.mu.A and often less than 100 nA). Typical binders include, for
example, polyurethane resins, cellulose derivatives, elastomers,
and highly fluorinated polymers. Examples of elastomers include
silicones, polymeric dienes, and acrylonitrile-butadiene-styrene
(ABS) resins. One example of a fluorinated polymer binder is
Teflon.RTM. (DuPont. Wilmington, Del.). These binders are cured
using, for example, heat or light, including ultraviolet (UV)
light. The appropriate curing method typically depends on the
particular binder which is used.
[0113] Often, when a liquid or semiliquid precursor of the
conductive material 56 (e.g., a conductive ink) is deposited in the
channel 54, the precursor fills the channel 54. However, when the
solvent or dispersant evaporates, the conductive material 56 which
remains may lose volume such that the conductive material 56 may or
may not continue to fill the channel 54. Preferred conductive
materials 56 do not pull away from the substrate 50 as they lose
volume, but rather decrease in height within the channel 54. These
conductive materials 56 typically adhere well to the substrate 50
and therefore do not pull away from the substrate 50 during
evaporation of the solvent or dispersant. Other suitable conductive
materials 56 either adhere to at least a portion of the substrate
50 and/or contain another additive, such as a binder, which adheres
the conductive material 56 to the substrate 50. Preferably, the
conductive material 56 in the channels 54 is non-leachable, and
more preferably immobilized on the substrate 50. In some
embodiments, the conductive material 56 may be formed by multiple
applications of a liquid or semiliquid precursor interspersed with
removal of the solvent or dispersant.
[0114] In another embodiment, the channels 54 are formed using a
laser. The laser carbonizes the polymer or plastic material. The
carbon formed in this process is used as the conductive material
56. Additional conductive material 56, such as a conductive carbon
ink, may be used to supplement the carbon formed by the laser.
[0115] In a further embodiment, the conductive traces 52 are formed
by pad printing techniques. For example, a film of conductive
material is formed either as a continuous film or as a coating
layer deposited on a carrier film. This film of conductive material
is brought between a print head and the substrate 50. A pattern on
the surface of the substrate 50 is made using the print head
according to a desired pattern of conductive traces 52. The
conductive material is transferred by pressure and/or heat from the
film of conductive material to the substrate 50. This technique
often produces channels (e.g., depressions caused by the print
head) in the substrate 50. Alternatively, the conductive material
is deposited on the surface of the substrate 50 without forming
substantial depressions.
[0116] In other embodiments, the conductive traces 52 are formed by
non-impact printing techniques. Such techniques include
electrophotography and magnetography. In these processes, an image
of the conductive traces 52 is electrically or magnetically formed
on a drum. A laser or LED may be used to electrically form an
image. A magnetic recording head may be used to magnetically form
an image. A toner material (e.g., a conductive material, such as a
conductive ink) is then attracted to portions of the drum according
to the image. The toner material is then applied to the substrate
by contact between the drum and the substrate. For example, the
substrate may be rolled over the drum. The toner material may then
be dried and/or a binder in the toner material may be cured to
adhere the toner material to the substrate.
[0117] Another non-impact printing technique includes ejecting
droplets of conductive material onto the substrate in a desired
pattern. Examples of this technique include ink jet printing and
piezo jet printing. An image is sent to the printer which then
ejects the conductive material (e.g., a conductive ink) according
to the pattern. The printer may provide a continuous stream of
conductive material or the printer may eject the conductive
material in discrete amounts at the desired points.
[0118] Yet another non-impact printing embodiment of forming the
conductive traces includes an ionographic process. In the this
process, a curable, liquid precursor, such as a photopolymerizable
acrylic resin (e.g., Solimer 7501 from Cubital, Bad Kreuznach,
Germany) is deposited over a surface of a substrate 50. A photomask
having a positive or negative image of the conductive traces 52 is
then used to cure the liquid precursor. Light (e.g., visible or
ultraviolet light) is directed through the photomask to cure the
liquid precursor and form a solid layer over the substrate
according to the image on the photomask. Uncured liquid precursor
is removed leaving behind channels 54 in the solid layer. These
channels 54 can then be filled with conductive material 56 to form
conductive traces 52.
[0119] Conductive traces 52 (and channels 54, if used) can be
formed with relatively narrow widths, for example, in the range of
25 to 250 .mu.m, and including widths of, for example, 250 .mu.m,
150 .mu.m, 100 .mu.m, 75 .mu.m, 50 .mu.m, 25 .mu.m or less by the
methods described above. In embodiments with two or more conductive
traces 52 on the same side of the substrate 50, the conductive
traces 52 are separated by distances sufficient to prevent
conduction between the conductive traces 52. The edge-to-edge
distance between the conductive traces is preferably in the range
of 25 to 250 .mu.m and may be, for example, 150 .mu.m, 100 .mu.m,
75 .mu.m, 50 .mu.m, or less. The density of the conductive traces
52 on the substrate 50 is preferably in the range of about 150 to
700 .mu.m/trace and may be as small as 667 .mu.m/trace or less, 333
.mu.m/trace or less, or even 167 .mu.m/trace or less.
[0120] The working electrode 58 and the counter electrode 60 (if a
separate reference electrode is used) are often made using a
conductive material 56, such as carbon. Suitable carbon conductive
inks are available from Ercon, Inc. (Wareham, Mass.), Metech, Inc.
(Elverson, Pa.), E. I. du Pont de Nemours and Co. (Wilmington,
Del.), Emca-Remex Products (Montgomeryville, Pa.), or MCA Services
(Melbourn, Great Britain). Typically, the working surface 51 of the
working electrode 58 is at least a portion of the conductive trace
52 that is in contact with the analyte-containing fluid (e.g.,
implanted in the patient).
[0121] The reference electrode 62 and/or counter/reference
electrode are typically formed using conductive material 56 that is
a suitable reference material, for example silver/silver chloride
or a non-leachable redox couple bound to a conductive material, for
example, a carbon-bound redox couple. Suitable silver/silver
chloride conductive inks are available from Ercon, Inc. (Wareham,
Mass.), Metech, Inc. (Elverson, Pa.), E. I. du Pont de Nemours and
Co. (Wilmington, Del.), Emca-Remex Products (Montgomeryville, Pa.),
or MCA Services (Melbourn, Great Britain). Silver/silver chloride
electrodes illustrate a type of reference electrode that involves
the reaction of a metal electrode with a constituent of the sample
or body fluid, in this case, Cl.sup.-.
[0122] Suitable redox couples for binding to the conductive
material of the reference electrode include, for example, redox
polymers (e.g., polymers having multiple redox centers.) It is
preferred that the reference electrode surface be non-corroding so
that an erroneous potential is not measured. Preferred conductive
materials include less corrosive metals, such as gold and
palladium. Most preferred are non-corrosive materials including
non-metallic conductors, such as carbon and conducting polymers. A
redox polymer can be adsorbed on or covalently bound to the
conductive material of the reference electrode, such as a carbon
surface of a conductive trace 52. Non-polymeric redox couples can
be similarly bound to carbon or gold surfaces.
[0123] A variety of methods may be used to immobilize a redox
polymer on an electrode surface. One method is adsorptive
immobilization. This method is particularly useful for redox
polymers with relatively high molecular weights. The molecular
weight of a polymer may be increased, for example, by
cross-linking.
[0124] Another method for immobilizing the redox polymer includes
the functionalization of the electrode surface and then the
chemical bonding, often covalently, of the redox polymer to the
functional groups on the electrode surface. One example of this
type of immobilization begins with a poly(4-vinylpyridine). The
polymer's pyridine rings are, in part, complexed with a
reducible/oxidizable species, such as [Os(bpy).sub.2Cl].sup.+/2+
where bpy is 2,2'-bipyridine. Part of the pyridine rings are
quaternized by reaction with 2-bromoethylamine. The polymer is then
crosslinked, for example, using a diepoxide, such as polyethylene
glycol diglycidyl ether.
[0125] Carbon surfaces can be modified for attachment of a redox
species or polymer, for example, by electroreduction of a diazonium
salt. As an illustration, reduction of a diazonium salt formed upon
diazotization of p-aminobenzoic acid modifies a carbon surface with
phenylcarboxylic acid functional groups. These functional groups
can then be activated by a carbodiimide, such as
1-ethyl-3-(3-dimethylaminopropyl)-carbodiimide hydrochloride. The
activated functional groups are then bound with a
amine-functionalized redox couple, such as the quaternized
osmium-containing redox polymer described above or
2-aminoethylferrocene, to form the redox couple.
[0126] Similarly, gold can be functionalized by an amine, such as
cystamine. A redox couple such as
[Os(bpy).sub.2(pyridine-4-carboxylate)C- l].sup.0/+ is activated by
1-ethyl-3-(3-dimethylaminopropyl)-carbodiimide hydrochloride to
form a reactive O-acylisourea which reacts with the gold-bound
amine to form an amide.
[0127] In one embodiment, in addition to using the conductive
traces 52 as electrodes or probe leads, two or more of the
conductive traces 52 on the substrate 50 are used to give the
patient a mild electrical shock when, for example, the analyte
level exceeds a threshold level. This shock may act as a warning or
alarm to the patient to initiate some action to restore the
appropriate level of the analyte.
[0128] The mild electrical shock is produced by applying a
potential between any two conductive traces 52 that are not
otherwise connected by a conductive path. For example, two of the
electrodes 58, 60, 62 or one electrode 58, 60, 62 and the
temperature probe 66 may be used to provide the mild shock.
Preferably, the working electrode 58 and the reference electrode 62
are not used for this purpose as this may cause some damage to the
chemical components on or proximate to the particular electrode
(e.g., the sensing layer on the working electrode or the redox
couple on the reference electrode).
[0129] The current used to produce the mild shock is typically 0.1
to 1 mA. Higher or lower currents may be used, although care should
be taken to avoid harm to the patient. The potential between the
conductive traces is typically 1 to 10 volts. However, higher or
lower voltages may be used depending, for example, on the
resistance of the conductive traces 52, the distance between the
conductive traces 52 and the desired amount of current. When the
mild shock is delivered, potentials at the working electrode 58 and
across the temperature probe 66 may be removed to prevent harm to
those components caused by unwanted conduction between the working
electrode 58 (and/or temperature probe 66, if used) and the
conductive traces 52 which provide the mild shock.
[0130] Contact Pads
[0131] Typically, each of the conductive traces 52 includes a
contact pad 49. The contact pad 49 may simply be a portion of the
conductive trace 52 that is indistinguishable from the rest of the
trace 52 except that the contact pad 49 is brought into contact
with the conductive contacts of a control unit (e.g., the sensor
control unit 44 of FIG. 1). More commonly, however, the contact pad
49 is a region of the conductive trace 52 that has a larger width
than other regions of the trace 52 to facilitate a connection with
the contacts on the control unit. By making the contact pads 49
relatively large as compared with the width of the conductive
traces 52, the need for precise registration between the contact
pads 49 and the contacts on the control unit is less critical than
with small contact pads.
[0132] The contact pads 49 are typically made using the same
material as the conductive material 56 of the conductive traces 52.
However, this is not necessary. Although metal, alloys, and
metallic compounds may be used to form the contact pads 49, in some
embodiments, it is desirable to make the contact pads 49 from a
carbon or other non-metallic material, such as a conducting
polymer. In contrast to metal or alloy contact pads, carbon and
other non-metallic contact pads are not easily corroded if the
contact pads 49 are in a wet, moist, or humid environment. Metals
and alloys may corrode under these conditions, particularly if the
contact pads 49 and contacts of the control unit are made using
different metals or alloys. However, carbon and non-metallic
contact pads 49 do not significantly corrode, even if the contacts
of the control device are metal or alloy.
[0133] One embodiment of the invention includes a sensor 42 having
contact pads 49 and a control unit 44 having conductive contacts
(not shown). During operation of the sensor 42, the contact pads 49
and conductive contacts are in contact with each other. In this
embodiment, either the contact pads 49 or the conductive contacts
are made using a non-corroding, conductive material. Such materials
include, for example, carbon and conducting polymers. Preferred
non-corroding materials include graphite and vitreous carbon. The
opposing contact pad or conductive contact is made using carbon, a
conducting polymer, a metal, such as gold, palladium, or platinum
group metal, or a metallic compound, such as ruthenium dioxide.
This configuration of contact pads and conductive contacts
typically reduces corrosion. Preferably, when the sensor is placed
in a 3 mM, and more preferably, in a 100 mM, NaCl solution, the
signal arising due to the corrosion of the contact pads and/or
conductive contacts is less than 3% of the signal generated by the
sensor when exposed to concentration of analyte in the normal
physiological range. For at least some subcutaneous glucose
sensors, the current generated by analyte in a normal physiological
range ranges from 3 to 500 nA.
[0134] Each of the electrodes 58, 60, 62, as well as the two probe
leads 68, 70 of the temperature probe 66 (described below), are
connected to contact pads 49 as shown in FIGS. 10 and 11. In one
embodiment (not shown), the contact pads 49 are on the same side of
the substrate 50 as the respective electrodes or temperature probe
leads to which the contact pads 49 are attached.
[0135] In other embodiments, the conductive traces 52 on at least
one side are connected through vias in the substrate to contact
pads 49a on the opposite surface of the substrate 50, as shown in
FIGS. 10 and 11. An advantage of this configuration is that contact
between the contacts on the control unit and each of the electrodes
58, 60, 62 and the probe leads 68,70 of the temperature probe 66
can be made from a single side of the substrate 50.
[0136] In yet other embodiments (not shown), vias through the
substrate are used to provide contact pads on both sides of the
substrate 50 for each conductive trace 52. The vias connecting the
conductive traces 52 with the contact pads 49a can be formed by
making holes through the substrate 50 at the appropriate points and
then filling the holes with conductive material 56.
[0137] Exemplary Electrode Configurations
[0138] A number of exemplary electrode configurations are described
below, however, it will be understood that other configurations may
also be used. In one embodiment, illustrated in FIG. 3A, the sensor
42 includes two working electrodes 58a, 58b and one counter
electrode 60, which also functions as a reference electrode. In
another embodiment, the sensor includes one working electrode 58a,
one counter electrode 60, and one reference electrode 62, as shown
in FIG. 3B. Each of these embodiments is illustrated with all of
the electrodes formed on the same side of the substrate 50.
[0139] Alternatively, one or more of the electrodes may be formed
on an opposing side of the substrate 50. This may be convenient if
the electrodes are formed using two different types of conductive
material 56 (e.g., carbon and silver/silver chloride). Then, at
least in some embodiments, only one type of conductive material 56
needs to be applied to each side of the substrate 50, thereby
reducing the number of steps in the manufacturing process and/or
easing the registration constraints in the process. For example, if
the working electrode 58 is formed using a carbon-based conductive
material 56 and the reference or counter/reference electrode is
formed using a silver/silver chloride conductive material 56, then
the working electrode and reference or counter/reference electrode
may be formed on opposing sides of the substrate 50 for ease of
manufacture.
[0140] In another embodiment, two working electrodes 58 and one
counter electrode 60 are formed on one side of the substrate 50 and
one reference electrode 62 and a temperature probe 66 are formed on
an opposing side of the substrate 50, as illustrated in FIG. 6. The
opposing sides of the tip of this embodiment of the sensor 42 are
illustrated in FIGS. 7 and 8.
[0141] Sensing Layer
[0142] Some analytes, such as oxygen, can be directly
electrooxidized or electroreduced on the working electrode 58.
Other analytes, such as glucose and lactate, require the presence
of at least one electron transfer agent and/or at least one
catalyst to facilitate the electrooxidation or electroreduction of
the analyte. Catalysts may also be used for those analyte, such as
oxygen, that can be directly electrooxidized or electroreduced on
the working electrode 58. For these analytes, each working
electrode 58 has a sensing layer 64 formed proximate to or on a
working surface of the working electrode 58. Typically, the sensing
layer 64 is formed near or on only a small portion of the working
electrode 58, often near a tip of the sensor 42. This limits the
amount of material needed to form the sensor 42 and places the
sensing layer 64 in the best position for contact with the
analyte-containing fluid (e.g., a body fluid, sample fluid, or
carrier fluid).
[0143] The sensing layer 64 includes one or more components
designed to facilitate the electrolysis of the analyte. The sensing
layer 64 may include, for example, a catalyst to catalyze a
reaction of the analyte and produce a response at the working
electrode 58, an electron transfer agent to indirectly or directly
transfer electrons between the analyte and the working electrode
58, or both.
[0144] The sensing layer 64 may be formed as a solid composition of
the desired components (e.g., an electron transfer agent and/or a
catalyst). These components are preferably non-leachable from the
sensor 42 and more preferably are immobilized on the sensor 42. For
example, the components may be immobilized on a working electrode
58. Alternatively, the components of the sensing layer 64 may be
immobilized within or between one or more membranes or films
disposed over the working electrode 58 or the components may be
immobilized in a polymeric or sol-gel matrix. Examples of
immobilized sensing layers are described in U.S. Pat. Nos.
5,262,035, 5,264,104, 5,264,105, 5,320,725, 5,593,852, and
5,665,222, U.S. patent application Ser. No. 08/540,789, and PCT
Patent Application No. US98/02403 entitled "Soybean Peroxidase
Electrochemical Sensor", filed on Feb. 11, 1998, Attorney Docket
No. M&G 12008.8WOI2, incorporated herein by reference.
[0145] In some embodiments, one or more of the components of the
sensing layer 64 may be solvated, dispersed, or suspended in a
fluid within the sensing layer 64, instead of forming a solid
composition. The fluid may be provided with the sensor 42 or may be
absorbed by the sensor 42 from the analyte-containing fluid.
[0146] Preferably, the components which are solvated, dispersed, or
suspended in this type of sensing layer 64 are non-leachable from
the sensing layer. Non-leachability may be accomplished, for
example, by providing barriers (e.g., the electrode, substrate,
membranes, and/or films) around the sensing layer which prevent the
leaching of the components of the sensing layer 64. One example of
such a barrier is a microporous membrane or film which allows
diffusion of the analyte into the sensing layer 64 to make contact
with the components of the sensing layer 64, but reduces or
eliminates the diffusion of the sensing layer components (e.g., a
electron transfer agent and/or a catalyst) out of the sensing layer
64.
[0147] A variety of different sensing layer configurations can be
used. In one embodiment, the sensing layer 64 is deposited on the
conductive material 56 of a working electrode 58a, as illustrated
in FIGS. 3A and 3B. The sensing layer 64 may extend beyond the
conductive material 56 of the working electrode 58a. In some cases,
the sensing layer 64 may also extend over the counter electrode 60
or reference electrode 62 without degrading the performance of the
glucose sensor. For those sensors 42 which utilize channels 54
within which the conductive material 56 is deposited, a portion of
the sensing layer 64 may be formed within the channel 54 if the
conductive material 56 does not fill the channel 54.
[0148] A sensing layer 64 in direct contact with the working
electrode 58a may contain an electron transfer agent to transfer
electrons directly or indirectly between the analyte and the
working electrode, as well as a catalyst to facilitate a reaction
of the analyte. For example, a glucose, lactate, or oxygen
electrode may be formed having a sensing layer which contains a
catalyst, such as glucose oxidase, lactate oxidase, or laccase,
respectively, and an electron transfer agent that facilitates the
electrooxidation of the glucose, lactate, or oxygen,
respectively.
[0149] In another embodiment, the sensing layer 64 is not deposited
directly on the working electrode 58a. Instead, the sensing layer
64 is spaced apart from the working electrode 58a, as illustrated
in FIG. 4A, and separated from the working electrode 58a by a
separation layer 61. The separation layer 61 typically includes one
or more membranes or films. In addition to separating the working
electrode 58a from the sensing layer 64, the separation layer 61
may also act as a mass transport limiting layer or an interferent
eliminating layer, as described below.
[0150] Typically, a sensing layer 64, which is not in direct
contact with the working electrode 58a, includes a catalyst that
facilitates a reaction of the analyte. However, this sensing layer
64 typically does not include an electron transfer agent that
transfers electrons directly from the working electrode 58a to the
analyte, as the sensing layer 64 is spaced apart from the working
electrode 58a. One example of this type of sensor is a glucose or
lactate sensor which includes an enzyme (e.g., glucose oxidase or
lactate oxidase, respectively) in the sensing layer 64. The glucose
or lactate reacts with a second compound (e.g., oxygen) in the
presence of the enzyme. The second compound is then electrooxidized
or electroreduced at the electrode. Changes in the signal at the
electrode indicate changes in the level of the second compound in
the fluid and are proportional to changes in glucose or lactate
level and, thus, correlate to the analyte level.
[0151] In another embodiment, two sensing layers 63, 64 are used,
as shown in FIG. 4B. Each of the two sensing layers 63, 64 may be
independently formed on the working electrode 58a or in proximity
to the working electrode 58a. One sensing layer 64 is typically,
although not necessarily, spaced apart from the working electrode
58a. For example, this sensing layer 64 may include a catalyst
which catalyzes a reaction of the analyte to form a product
compound. The product compound is then electrolyzed in the second
sensing layer 63 which may include an electron transfer agent to
transfer electrons between the working electrode 58a and the
product compound and/or a second catalyst to catalyze a reaction of
the product compound to generate a signal at the working electrode
58a.
[0152] For example, a glucose or lactate sensor may include a first
sensing layer 64 which is spaced apart from the working electrode
and contains an enzyme, for example, glucose oxidase or lactate
oxidase. The reaction of glucose or lactate in the presence of the
appropriate enzyme forms hydrogen peroxide. A second sensing layer
63 is provided directly on the working electrode 58a and contains a
peroxidase enzyme and an electron transfer agent to generate a
signal at the electrode in response to the hydrogen peroxide. The
level of hydrogen peroxide indicated by the sensor then correlates
to the level of glucose or lactate. Another sensor which operates
similarly can be made using a single sensing layer with both the
glucose or lactate oxidase and the peroxidase being deposited in
the single sensing layer.
[0153] Examples of such sensors are described in U.S. Pat. No.
5,593,852, U.S. patent application Ser. No. 08/540,789, and PCT
Patent Application No. US98/02403 entitled "Soybean Peroxidase
Electrochemical Sensor", filed on Feb. 11, 1998, Attorney Docket
No. M&G 12008.8WOI2, incorporated herein by reference.
[0154] In some embodiments, one or more of the working electrodes
58b do not have a corresponding sensing layer 64, as shown in FIGS.
3A and 4A, or have a sensing layer (not shown) which does not
contain one or more components (e.g., an electron transfer agent or
catalyst) needed to electrolyze the analyte. The signal generated
at this working electrode 58b typically arises from interferents
and other sources, such as ions, in the fluid, and not in response
to the analyte (because the analyte is not electrooxidized or
electroreduced). Thus, the signal at this working electrode 58b
corresponds to a background signal. The background signal can be
removed from the analyte signal obtained from other working
electrodes 58a that are associated with fully-functional sensing
layers 64 by, for example, subtracting the signal at working
electrode 58b from the signal at working electrode 58a.
[0155] Sensors having multiple working electrodes 58a may also be
used to obtain more precise results by averaging the signals or
measurements generated at these working electrodes 58a. In
addition, multiple readings at a single working electrode 58a or at
multiple working electrodes may be averaged to obtain more precise
data.
[0156] Electron Transfer Agent
[0157] In many embodiments, the sensing layer 64 contains one or
more electron transfer agents in contact with the conductive
material 56 of the working electrode 58, as shown in FIGS. 3A and
3B. In some embodiments of the invention, there is little or no
leaching of the electron transfer agent away from the working
electrode 58 during the period in which the sensor 42 is implanted
in the patient. A diffusing or leachable (i.e., releasable)
electron transfer agent often diffuses into the analyte-containing
fluid, thereby reducing the effectiveness of the electrode by
reducing the sensitivity of the sensor over time. In addition, a
diffusing or leaching electron transfer agent in an implantable
sensor 42 may also cause damage to the patient. In these
embodiments, preferably, at least 90%, more preferably, at least
95%, and, most preferably, at least 99%, of the electron transfer
agent remains disposed on the sensor after immersion in the
analyte-containing fluid for 24 hours, and, more preferably, for 72
hours. In particular, for an implantable sensor, preferably, at
least 90%, more preferably, at least 95%, and most preferably, at
least 99%, of the electron transfer agent remains disposed on the
sensor after immersion in the body fluid at 37.degree. C. for 24
hours, and, more preferably, for 72 hours.
[0158] In some embodiments of the invention, to prevent leaching,
the electron transfer agents are bound or otherwise immobilized on
the working electrode 58 or between or within one or more membranes
or films disposed over the working electrode 58. The electron
transfer agent may be immobilized on the working electrode 58
using, for example, a polymeric or sol-gel immobilization
technique. Alternatively, the electron transfer agent may be
chemically (e.g., ionically, covalently, or coordinatively) bound
to the working electrode 58, either directly or indirectly through
another molecule, such as a polymer, that is in turn bound to the
working electrode 58.
[0159] Application of the sensing layer 64 on a working electrode
58a is one method for creating a working surface for the working
electrode 58a, as shown in FIGS. 3A and 3B. The electron transfer
agent mediates the transfer of electrons to electrooxidize or
electroreduce an analyte and thereby permits a current flow between
the working electrode 58 and the counter electrode 60 via the
analyte. The mediation of the electron transfer agent facilitates
the electrochemical analysis of analytes which are not suited for
direct electrochemical reaction on an electrode.
[0160] In general, the preferred electron transfer agents are
electroreducible and electrooxidizable ions or molecules having
redox potentials that are a few hundred millivolts above or below
the redox potential of the standard calomel electrode (SCE).
Preferably, the electron transfer agents are not more reducing than
about -150 mV and not more oxidizing than about +400 mV versus
SCE.
[0161] The electron transfer agent may be organic, organometallic,
or inorganic. Examples of organic redox species are quinones and
species that in their oxidized state have quinoid structures, such
as Nile blue and indophenol. Some quinones and partially oxidized
quinhydrones react with functional groups of proteins such as the
thiol groups of cysteine, the amine groups of lysine and arginine,
and the phenolic groups of tyrosine which may render those redox
species unsuitable for some of the sensors of the present invention
because of the presence of the interfering proteins in an
analyte-containing fluid. Usually substituted quinones and
molecules with quinoid structure are less reactive with proteins
and are preferred. A preferred tetrasubstituted quinone usually has
carbon atoms in positions 1, 2, 3, and 4.
[0162] In general, electron transfer agents suitable for use in the
invention have structures or charges which prevent or
substantially, reduce the diffusional loss of the electron transfer
agent during the period of time that the sample is being analyzed.
The preferred electron transfer agents include a redox species
bound to a polymer which can in turn be immobilized on the working
electrode. The bond between the redox species and the polymer may
be covalent, coordinative, or ionic. Useful electron transfer
agents and methods for producing them are described in U.S. Pat.
Nos. 5,264,104; 5,356,786; 5,262,035; and 5,320,725, incorporated
herein by reference. Although any organic or organometallic redox
species can be bound to a polymer and used as an electron transfer
agent, the preferred redox species is a transition metal compound
or complex. The preferred transition metal compounds or complexes
include osmium, ruthenium, iron, and cobalt compounds or complexes.
The most preferred are osmium compounds and complexes. It will be
recognized that many of the redox species described below may also
be used, typically without a polymeric component, as electron
transfer agents in a carrier fluid or in a sensing layer of a
sensor where leaching of the electron transfer agent is
acceptable.
[0163] One type of non-releasable polymeric electron transfer agent
contains a redox species covalently bound in a polymeric
composition. An example of this type of mediator is
poly(vinylferrocene).
[0164] Another type of non-releasable electron transfer agent
contains an ionically-bound redox species. Typically, this type of
mediator includes a charged polymer coupled to an oppositely
charged redox species. Examples of this type of mediator include a
negatively charged polymer such as Nafion.RTM. (DuPont) coupled to
a positively charged redox species such as an osmium or ruthenium
polypyridyl cation. Another example of an ionically-bound mediator
is a positively charged Polymer such as quaternized poly(4-vinyl
pyridine) or poly(1-vinyl imidazole) coupled to a negatively
charged redox species such as ferricyanide or ferrocyanide. The
preferred ionically-bound redox species is a highly charged redox
species bound within an oppositely charged redox polymer.
[0165] In another embodiment of the invention, suitable
non-releasable electron transfer agents include a redox species
coordinatively bound to a polymer. For example, the mediator may be
formed by coordination of an osmium or cobalt 2, 2'-bipyridyl
complex to poly(1-vinyl imidazole) or poly(4-vinyl pyridine).
[0166] The preferred electron transfer agents are osmium transition
metal complexes with one or more ligands, each ligand having a
nitrogen-containing heterocycle such as 2,2'-bipyridine,
1,10-phenanthroline, or derivatives thereof Furthermore, the
preferred electron transfer agents also have one or more ligands
covalently bound in a polymer, each ligand having at least one
nitrogen-containing heterocycle, such as pyridine, imidazole, or
derivatives thereof. These preferred electron transfer agents
exchange electrons rapidly between each other and the working
electrodes 58 so that the complex can be rapidly oxidized and
reduced.
[0167] One example of a particularly useful electron transfer agent
includes (a) a polymer or copolymer having pyridine or imidazole
functional groups and (b) osmium cations complexed with two
ligands, each ligand containing 2,2'-bipyridine.
1,10-phenanthroline, or derivatives thereof, the two ligands not
necessarily being the same. Preferred derivatives of
2,2.degree.-bipyridine for complexation with the osmium cation are
4,4'-dimethyl-2,2'-bipyridine and mono-, di-, and
polyalkoxy-2,2'-bipyridines, such as
4,4'-dimethoxy-2,2'-bipyridine. Preferred derivatives of
1,10-phenanthroline for complexation with the osmium cation are
4,7-dimethyl-1,10-phenanthroline and mono, di-, and
polyalkoxy-1,10-phenanthrolines, such as
4,7-dimethoxy-1,10-phenanthrolin- e. Preferred polymers for
complexation with the osmium cation include polymers and copolymers
of poly(1-vinyl imidazole) (referred to as "PVI") and poly(4-vinyl
pyridine) (referred to as "PVP"). Suitable copolymer substituents
of poly(1-vinyl imidazole) include acrylonitrile, acrylanide, and
substituted or quaternized N-vinyl imidazole. Most preferred are
electron transfer agents with osmium complexed to a polymer or
copolymer of poly(1-vinyl imidazole).
[0168] The preferred electron transfer agents have a redox
potential ranging from -100 mV to about +150 mV versus the standard
calomel electrode (SCE). Preferably, the potential of the electron
transfer agent ranges from -100 mV to +150 mV and more preferably,
the potential ranges from -50 mV to +50 mV. The most preferred
electron transfer agents have osmium redox centers and a redox
potential ranging from +50 mV to -150 mV versus SCE.
[0169] Catalyst
[0170] The sensing layer 64 may also include a catalyst which is
capable of catalyzing a reaction of the analyte. The catalyst may
also, in some embodiments, act as an electron transfer agent. One
example of a suitable catalyst is an enzyme which catalyzes a
reaction of the analyte. For example, a catalyst, such as a glucose
oxidase, glucose dehydrogenase (e.g., pyrroloquinoline quinone
glucose dehydrogenase (PQQ)), or oligosaccharide dehydrogenase, may
be used when the analyte is glucose. A lactate oxidase or lactate
dehydrogenase may be used when the analyte is lactate. Laccase may
be used when the analyte is oxygen or when oxygen is generated or
consumed in response to a reaction of the analyte.
[0171] Preferably, the catalyst is non-leachably disposed on the
sensor, whether the catalyst is part of a solid sensing layer in
the sensor or solvated in a fluid within the sensing layer. More
preferably, the catalyst is immobilized within the sensor (e.g. on
the electrode and/or within or between a membrane or film) to
prevent unwanted leaching of the catalyst away from the working
electrode 58 and into the patient. This may be accomplished, for
example, by attaching the catalyst to a polymer, cross linking the
catalyst with another electron transfer agent (which, as described
above, can be polymeric), and/or providing one or more barrier
membranes or films with pore sizes smaller than the catalyst.
[0172] As described above, a second catalyst may also be used. This
second catalyst is often used to catalyze a reaction of a product
compound resulting from the catalyzed reaction of the analyte. The
second catalyst typically operates with an electron transfer agent
to electrolyze the product compound to generate a signal at the
working electrode. Alternatively, the second catalyst may be
provided in an interferent-eliminating layer to catalyze reactions
that remove interferents, as described below.
[0173] One embodiment of the invention is an electrochemical sensor
in which the catalyst is mixed or dispersed in the conductive
material 56 which forms the conductive trace 52 of a working
electrode 58. This may be accomplished, for example, by mixing a
catalyst, such as an enzyme, in a carbon ink and applying the
mixture into a channel 54 on the surface of the substrate 50.
Preferably, the catalyst is immobilized in the channel 53 so that
it can not leach away from the working electrode 58. This may be
accomplished, for example, by curing a binder in the carbon ink
using a curing technique appropriate-to the binder. Curing
techniques include, for example, evaporation of a solvent or
dispersant, exposure to ultraviolet light, or exposure to heat.
Typically, the mixture is applied under conditions that do not
substantially degrade the catalyst. For example, the catalyst may
be an enzyme that is heat-sensitive. The enzyme and conductive
material mixture should be applied and cured, preferably, without
sustained periods of heating. The mixture may be cured using
evaporation or UV curing techniques or by the exposure to heat that
is sufficiently short that the catalyst is not substantially
degraded.
[0174] Another consideration for in vivo analyte sensors is the
thermostability of the catalyst. Many enzymes have only limited
stability at biological temperatures. Thus, it may be necessary to
use large amounts of the catalyst and/or use a catalyst that is
thermostable at the necessary temperature (e.g., 37.degree. C. or
higher for normal body temperature). A thermostable catalyst may be
defined as a catalyst which loses less than 5% of its activity when
held at 37.degree. C. for at least one hour, preferably, at least
one day, and more preferably at least three days. One example of a
thermostable catalyst is soybean peroxidase. This particular
thermostable catalyst may be used in a glucose or lactate sensor
when combined either in the same or separate sensing layers with
glucose or lactate oxidase or dehydrogenase. A further description
of thermostable catalysts and their use in electrochemical
inventions is found in U.S. Pat. No. 5,665,222 U.S. patent
application Ser. No. 08/540,789, and PCT Application No. US98/02403
entitled "Soybean Peroxidase Electrochemical Sensor", filed on Feb.
11, 1998, Attorney Docket No. M&G 12008.8WOI2.
[0175] Electrolysis of the Analyte
[0176] To electrolyze the analyte, a potential (versus a reference
potential) is applied across the working and counter electrodes 58,
60. The minimum magnitude of the applied potential is often
dependent on the particular electron transfer agent, analyte (if
the analyte is directly electrolyzed at the electrode), or second
compound (if a second compound, such as oxygen or hydrogen
peroxide, whose level is dependent on the analyte level, is
directly electrolyzed at the electrode). The applied potential
usually equals or is more oxidizing or reducing, depending on the
desired electrochemical reaction, than the redox potential of the
electron transfer agent, analyte, or second compound, whichever is
directly electrolyzed at the electrode. The potential at the
working electrode is typically large enough to drive the
electrochemical reaction to or near completion.
[0177] The magnitude of the potential may optionally be limited to
prevent significant (as determined by the current generated in
response to the analyte) electrochemical reaction of interferents,
such as urate, ascorbate, and acetaminophen. The limitation of the
potential may be obviated if these interferents have been removed
in another way, such as by providing an interferent-limiting
barrier, as described below, or by including a working electrode
58b (see FIG. 3A) from which a background signal may be
obtained.
[0178] When a potential is applied between the working electrode 58
and the counter electrode 60, an electrical current will flow. The
current is a result of the electrolysis of the analyte or a second
compound whose level is affected by the analyte. In one embodiment,
the electrochemical reaction occurs via an electron transfer agent
and the optional catalyst. Many analytes B are oxidized (or
reduced) to products C by an electron transfer agent species A in
the presence of an appropriate catalyst (e.g., an enzyme). The
electron transfer agent A is then oxidized (or reduced) at the
electrode. Electrons are collected by (or removed from) the
electrode and the resulting current is measured. This process is
illustrated by reaction equations (1) and (2) (similar equations
may be written for the reduction of the analyte B by a redox
mediator A in the presence of a catalyst): 1
[0179] As an example, an electrochemical sensor may be based on the
reaction of a glucose molecule with two non-leachable ferricyanide
anions in the presence of glucose oxidase to produce two
non-leachable ferrocyanide anions, two hydrogen ions, and
gluconolactone. The amount of glucose present is assayed by
electrooxidizing the non-leachable ferrocyanide anions to
non-leachable ferricyanide anions and measuring the current.
[0180] In another embodiment, a second compound whose level is
affected by the analyte is electrolyzed at the working electrode.
In some cases, the analyte D and the second compound, in this case,
a reactant compound E, such as oxygen, react in the presence of the
catalyst, as shown in reaction equation (3). 2
[0181] The reactant compound E is then directly oxidized (or
reduced) at the working electrode. as shown in reaction equation
(4) electrode 3
[0182] Alternatively, the reactant compound E is indirectly
oxidized (or reduced) using an electron transfer agent H
(optionally in the presence of a catalyst), that is subsequently
reduced or oxidized at the electrode, as shown in reaction
equations (5) and (6). 4
[0183] In either case, changes in the concentration of the reactant
compound, as indicated by the signal at the working electrode,
correspond inversely to changes in the analyte (i.e., as the level
of analyte increase then the level of reactant compound and the
signal at the electrode decreases.)
[0184] In other embodiments, the relevant second compound is a
product compound F, as shown in reaction equation (3). The product
compound F is formed by the catalyzed reaction of analyte D and
then be directly electrolyzed at the electrode or indirectly
electrolyzed using an electron transfer agent and, optionally, a
catalyst. In these embodiments, the signal arising from the direct
or indirect electrolysis of the product compound F at the working
electrode corresponds directly to the level of the analyte (unless
there are other sources of the product compound). As the level of
analyte increases, the level of the product compound and signal at
the working electrode increases.
[0185] Those skilled in the art will recognize that there are many
different reactions that will achieve the same result; namely the
electrolysis of an analyte or a compound whose level depends on the
level of the analyte. Reaction equations (1) through (6) illustrate
non-limiting examples of such reactions.
[0186] Temperature Probe
[0187] A variety of optional items may be included in the sensor.
One optional item is a temperature probe 66 (FIGS. 8 and 11). The
temperature probe 66 may be made using a variety of known designs
and materials. One exemplary temperature probe 66 is formed using
two probe leads 68, 70 connected to each other through a
temperature-dependent element 72 that is formed using a material
with a temperature-dependent characteristic. An example of a
suitable temperature-dependent characteristic is the resistance of
the temperature-dependent element 72.
[0188] The two probe leads 68, 70 are typically formed using a
metal, an alloy, a semimetal, such as graphite, a degenerate or
highly doped semiconductor, or a small-band gap semiconductor.
Examples of suitable materials include gold, silver, ruthenium
oxide, titanium nitride, titanium dioxide, indium doped tin oxide,
tin doped indium oxide, or graphite. The temperature-dependent
element 72 is typically made using a fine trace (e.g., a conductive
trace that has a smaller cross-section than that of the probe leads
68, 70) of the same conductive material as the probe leads, or
another material such as a carbon ink, a carbon fiber, or platinum,
which has a temperature-dependent characteristic, such as
resistance, that provides a temperature-dependent signal when a
voltage source is attached to the two probe leads 68, 70 of the
temperature probe 66. The temperature-dependent characteristic of
the temperature-dependent element 72 may either increase or
decrease with temperature. Preferably, the temperature dependence
of the characteristic of the temperature-dependent element 72 is
approximately linear with temperature over the expected range of
biological temperatures (about 25 to 45.degree. C.), although this
is not required.
[0189] Typically, a signal (e.g., a current) having an amplitude or
other property that is a function of the temperature can be
obtained by providing a potential across the two probe leads 68, 70
of the temperature probe 66. As the temperature changes, the
temperature-dependent characteristic of the temperature-dependent
element 72 increases or decreases with a corresponding change in
the signal amplitude. The signal from the temperature probe 66
(e.g., the amount of current flowing through the probe) may be
combined with the signal obtained from the working electrode 58 by,
for example, scaling the temperature probe signal and then adding
or subtracting the scaled temperature probe signal from the signal
at the working electrode 58. In this manner, the temperature probe
66 can provide a temperature adjustment for the output from the
working electrode 58 to offset the temperature dependence of the
working electrode 58.
[0190] One embodiment of the temperature probe includes probe leads
68, 70 formed as two spaced-apart channels with a
temperature-dependent element 72 formed as a cross-channel
connecting the two spaced-apart channels, as illustrated in FIG. 8.
The two spaced-apart channels contain a conductive material, such
as a metal, alloy, semimetal, degenerate semiconductor, or metallic
compound. The cross-channel may contain the same material (provided
the cross-channel has a smaller cross-section than the two
spaced-apart channels) as the probe leads 68, 70. In other
embodiments, the material in the cross-channel is different than
the material of the probe leads 68, 70.
[0191] One exemplary method for forming this particular temperature
probe includes forming the two spaced-apart channels and then
filling them with the metallic or alloyed conductive material.
Next, the cross-channel is formed and then filled with the desired
material. The material in the cross-channel overlaps with the
conductive material in each of the two spaced-apart channels to
form an electrical connection.
[0192] For proper operation of the temperature probe 66, the
temperature-dependent element 72 of the temperature probe 66 can
not be shorted by conductive material formed between the two probe
leads 68, 70. In addition, to prevent conduction between the two
probe leads 68, 70 by ionic species within the body or sample
fluid, a covering may be provided over the temperature-dependent
element 72, and preferably over the portion of the probe leads 68,
70 that is implanted in the patient. The covering may be, for
example, a non-conducting film disposed over the
temperature-dependent element 72 and probe leads 68, 70 to prevent
the ionic conduction. Suitable non-conducting films include, for
example, Kapton.TM. polyimide films (DuPont. Wilmington, Del.).
[0193] Another method for eliminating or reducing conduction by
ionic species in the body or sample fluid is to use an ac voltage
source connected to the probe leads 68, 70. In this way, the
positive and negative ionic species are alternately attracted and
repelled during each half cycle of the ac voltage. This results in
no net attraction of the ions in the body or sample fluid to the
temperature probe 66. The maximum amplitude of the ac current
through the temperature-dependent element 72 may then be used to
correct the measurements from the working electrodes 58.
[0194] The temperature probe can be placed on the same substrate as
the electrodes. Alternatively, a temperature probe may be placed on
a separate substrate. In addition, the temperature probe may be
used by itself or in conjunction with other devices.
[0195] Another embodiment of a temperature probe utilizes the
temperature dependence of the conductivity of a solution (e.g.,
blood or interstitial fluid). Typically, the conductivity of an
electrolyte-containing solution is dependent on the temperature of
the solution, assuming that the concentration of electrolytes is
relatively constant. Blood, interstitial fluid, and other bodily
fluids are solutions with relatively constant levels of
electrolytes. Thus, a sensor 42 can include two or more conductive
traces (not shown) which are spaced apart by a known distance. A
portion of these conductive traces is exposed to the solution and
the conductivity between the exposed portions of the conductive
traces is measured using known techniques (e.g., application of a
constant or known current or potential and measurement of the
resulting potential or current, respectively, to determine the
conductivity).
[0196] A change in conductivity is related to a change in
temperature. This relation can be modeled using linear, quadratic,
exponential, or other relations. The parameters for this
relationship typically do not vary significantly between most
people. The calibration for the temperature probe can be determined
by a variety of methods, including, for example, calibration of
each sensor 42 using an independent method of determining
temperature (e.g., a thermometer, an optical or electrical
temperature detector, or the temperature probe 66, described above)
or calibrating one sensor 42 and using that calibration for all
other sensors in a batch based on uniformity in geometry.
[0197] Biocompatible Layer
[0198] An optional film layer 75 is formed over at least that
portion of the sensor 42 which is subcutaneously inserted into the
patient, as shown in FIG. 9. This optional film layer 74 may serve
one or more functions. The film layer 74 prevents the penetration
of large biomolecules into the electrodes. This is accomplished by
using a film layer 74 having a pore size that is smaller than the
biomolecules that are to be excluded. Such biomolecules may foul
the electrodes and/or the sensing layer 64 thereby reducing the
effectiveness of the sensor 42 and altering the expected signal
amplitude for a given analyte concentration. The fouling of the
working electrodes 58 may also decrease the effective life of the
sensor 42. The biocompatible layer 74 may also prevent protein
adhesion to the sensor 42, formation of blood clots, and other
undesirable interactions between the sensor 42 and body.
[0199] For example, the sensor may be completely or partially
coated on its exterior with a biocompatible coating. A preferred
biocompatible coating is a hydrogel which contains at least 20 wt.
% fluid when in equilibrium with the analyte-containing fluid.
Examples of suitable hydrogels are described in U.S. Pat. No.
5,593,852, incorporated herein by reference, and include
crosslinked polyethylene oxides, such as polyethylene oxide
tetraacrylate.
[0200] Interferent-Eliminating Layer
[0201] An interferent-eliminating layer (not shown) may be included
in the sensor 42. The interferent-eliminating layer may be
incorporated in the biocompatible layer 75 or in the mass transport
limiting layer 74 (described below) or may be a separate layer.
Interferents are molecules or other species that are electroreduced
or electrooxidized at the electrode, either directly or via an
electron transfer agent, to produce a false signal. In one
embodiment, a film or membrane prevents the penetration of one or
more interferents into the region around the working electrodes 58.
Preferably, this type of interferent-eliminating layer is much less
permeable to one or more of the interferents than to the
analyte.
[0202] The interferent-eliminating layer may include ionic
components, such as Nafion.RTM., incorporated into a polymeric
matrix to reduce the permeability of the interferent-eliminating
layer to ionic interferents having the same charge as the ionic
components. For example, negatively charged compounds or compounds
that form negative ions may be incorporated in the
interferent-eliminating layer to reduce the permeation of negative
species in the body or sample fluid.
[0203] Another example of an interferent-eliminating layer includes
a catalyst for catalyzing a reaction which removes interferents.
One example of such a catalyst is a peroxidase. Hydrogen peroxide
reacts with interferents, such as acetaminophen, urate, and
ascorbate. The hydrogen peroxide may be added to the
analyte-containing fluid or may be generated in situ, by, for
example, the reaction of glucose or lactate in the presence of
glucose oxidase or lactate oxidase, respectively. Examples of
interferent eliminating layers include a peroxidase enzyme
crosslinked (a) using gluteraldehyde as a crosslinking agent or (b)
oxidation of oligosaccharide groups in the peroxidase glycoenzyme
with NaIO.sub.4, followed by coupling of the aldehydes formed to
hydrazide groups in a polyacrylamide matrix to form hydrazones are
describe in U.S. Pat. Nos. 5,262,305 and 5,356,786, incorporated
herein by reference.
[0204] Mass Transport Limiting Layer
[0205] A mass transport limiting layer 74 may be included with the
sensor to act as a diffusion-limiting barrier to reduce the rate of
mass transport of the analyte, for example, glucose or lactate,
into the region around the working electrodes 58. By limiting the
diffusion of the analyte, the steady state concentration of the
analyte in the proximity of the working electrode 58 (which is
proportional to the concentration of the analyte in the body or
sample fluid) can be reduced. This extends the upper range of
analyte concentrations that can still be accurately measured and
may also expand the range in which the current increases
approximately linearly with the level of the analyte.
[0206] It is preferred that the permeability of the analyte through
the film layer 74 vary little or not at all with temperature, so as
to reduce or eliminate the variation of current with temperature.
For this reason, it is preferred that in the biologically relevant
temperature range from about 25.degree. C. to about 45.degree. C.
and most importantly from 30.degree. C. to 40.degree. C., neither
the size of the pores in the film nor its hydration or swelling
change excessively. Preferably, the mass transport limiting layer
is made using a film that absorbs less than 5 wt. % of fluid over
24 hours. This may reduce or obviate any need for a temperature
probe. For implantable sensors, it is preferable that the mass
transport limiting layer is made using a film that absorbs less
than 5 wt. % of fluid over 24 hours at 37.degree. C.
[0207] Particularly useful materials for the film layer 74 are
membranes that do not swell in the analyte-containing fluid that
the sensor tests. Suitable membranes include 3 to 20,000 nm
diameter pores. Membranes having 5 to 500 nm diameter pores with
well-defined, uniform pore sizes and high aspect ratios are
preferred. In one embodiment, the aspect ratio of the pores is
preferably two or greater and more preferably five or greater.
[0208] Well-defined and uniform pores can be made by track etching
a polymeric membrane using accelerated electrons, ions, or
particles emitted by radioactive nuclei. Most preferred are
anisotropic, polymeric, track etched membranes that expand less in
the direction perpendicular to the pores than in the direction of
the pores when heated. Suitable polymeric membranes included
polycarbonate membranes from Poretics (Livermore, CA, catalog
number 19401, 0.01 .mu.m pore size polycarbonate membrane) and
Coming Costar Corp. (Cambridge, Mass., Nucleopore.TM. brand
membranes with 0.015 .mu.m pore size). Other polyolefin and
polyester films may be used. It is preferred that the permeability
of the mass transport limiting membrane changes no more than 4%,
preferably, no more than 3%, and, more preferably, no more than 2%,
per .degree. C. in the range from 30.degree. C. to 40.degree. C.
when the membranes resides in the subcutaneous interstitial
fluid.
[0209] In some embodiments of the invention, the mass transport
limiting layer 74 may also limit the flow of oxygen into the sensor
42. This can improve the stability of sensors 42 that are used in
situations where variation in the partial pressure of oxygen causes
non-linearity in sensor response. In these embodiments, the mass
transport limiting layer 74 restricts oxygen transport by at least
40%, preferably at least 60%, and more preferably at least 80%,
than the membrane restricts transport of the analyte. For a given
type of polymer, films having a greater density (e.g., a density
closer to that of the crystalline polymer) are preferred.
Polyesters, such as polyethylene terephthalate, are typically less
permeable to oxygen and are, therefore, preferred over
polycarbonate membranes.
[0210] Anticlotting Agent
[0211] An implantable sensor may also, optionally, have an
anticlotting agent disposed on a portion the substrate which is
implanted into a patient. This anticlotting agent may reduce or
eliminate the clotting of blood or other body fluid around the
sensor, particularly after insertion of the sensor. Blood clots may
foul the sensor or irreproducibly reduce the amount of analyte
which diffuses into the sensor. Examples of useful anticlotting
agents include heparin and tissue plasminogen activator (TPA), as
well as other known anticlotting agents.
[0212] The anticlotting agent may be applied to at least a portion
of that part of the sensor 42 that is to be implanted. The
anticlotting agent may be applied, for example, by bath, spraying,
brushing, or dipping. The anticlotting agent is allowed to dry on
the sensor 42. The anticlotting agent may be immobilized on the
surface of the sensor or it may be allowed to diffuse away from the
sensor surface. Typically, the quantities of anticlotting agent
disposed on the sensor are far below the amounts typically used for
treatment of medical conditions involving blood clots and,
therefore, have only a limited, localized effect.
[0213] Sensor Lifetime
[0214] The sensor 42 may be designed to be a replaceable component
in an in vivo analyte monitor, and particularly in an implantable
analyte monitor. Typically, the sensor 42 is capable of operation
over a period of days. Preferably, the period of operation is at
least one day, more preferably at least three days, and most
preferably at least one week. The sensor 42 can then be removed and
replaced with a new sensor. The lifetime of the sensor 42 may be
reduced by the fouling of the electrodes or by the leaching of the
electron transfer agent or catalyst. These limitations on the
longevity of the sensor 42 can be overcome by the use of a
biocompatible layer 75 or non-leachable electron transfer agent and
catalyst, respectively, as described above.
[0215] Another primary limitation on the lifetime of the sensor 42
is the temperature stability of the catalyst. Many catalysts are
enzymes, which are very sensitive to the ambient temperature and
may degrade at temperatures of the patient's body (e.g.,
approximately 37.degree. C. for the human body). Thus, robust
enzymes should be used where available. The sensor 42 should be
replaced when a sufficient amount of the enzyme has been
deactivated to introduce an unacceptable amount of error in the
measurements.
[0216] Insertion Device
[0217] An insertion device 120 can be used to subcutaneously insert
the sensor 42 into the patient, as illustrated in FIG. 12. The
insertion device 120 is typically formed using structurally rigid
materials, such as metal or rigid plastic. Preferred materials
include stainless steel and ABS (acrylonitrile-butadiene-styrene)
plastic. In some embodiments, the insertion device 120 is pointed
and/or sharp at the tip 121 to facilitate penetration of the skin
of the patient. A sharp, thin insertion device may reduce pain felt
by the patient upon insertion of the sensor 42. In other
embodiments, the tip 121 of the insertion device 120 has other
shapes, including a blunt or flat shape. These embodiments may be
particularly useful when the insertion device 120 does not
penetrate the skin but rather serves as a structural support for
the sensor 42 as the sensor 42 is pushed into the skin.
[0218] The insertion device 120 may have a variety of
cross-sectional shapes, as shown in FIGS. 13A, 13B, and 13C. The
insertion device 120 illustrated in FIG. 13A is a flat, planar,
pointed strip of rigid material which may be attached or otherwise
coupled to the sensor 42 to ease insertion of the sensor 42 into
the skin of the patient, as well as to provide structural support
to the sensor 42 during insertion. The insertion devices 120 of
FIGS. 13B and 13C are U- or V-shaped implements that support the
sensor 42 to limit the amount that the sensor 42 may bend or bow
during insertion. The cross-sectional width 124 of the insertion
devices 120 illustrated in FIGS. 13B and 13C is typically 1 mm or
less, preferably 700 .mu.m or less, more preferably 500 .mu.m or
less, and most preferably 300 .mu.m or less. The cross-sectional
height 126 of the insertion device 120 illustrated in FIGS. 13B and
13C is typically about 1 mm or less, preferably about 700 .mu.m or
less, and more preferably about 500 .mu.m or less.
[0219] The sensor 42 itself may include optional features to
facilitate insertion. For example, the sensor 42 may be pointed at
the tip 123 to ease insertion, as illustrated in FIG. 12. In
addition, the sensor 42 may include a barb 125 which helps retain
the sensor 42 in the subcutaneous tissue of the patient. The barb
125 may also assist in anchoring the sensor 42 within the
subcutaneous tissue of the patient during operation of the sensor
42. However, the barb 125 is typically small enough that little
damage is caused to the subcutaneous tissue when the sensor 42 is
removed for replacement. The sensor 42 may also include a notch 127
that can be used in cooperation with a corresponding structure (not
shown) in the insertion device to apply pressure against the sensor
42 during insertion, but disengage as the insertion device 120 is
removed. One example of such a structure in the insertion device is
a rod (not shown) between two opposing sides of an insertion device
120 and at an appropriate height of the insertion device 120.
[0220] In operation, the sensor 42 is placed within or next to the
insertion device 120 and then a force is provided against the
insertion device 120 and/or sensor 42 to carry the sensor 42 into
the skin of the patient. In one embodiment, the force is applied to
the sensor 42 to push the sensor into the skin, while the insertion
device 120 remains stationary and provides structural support to
the sensor 42. Alternatively, the force is applied to the insertion
device 120 and optionally to the sensor 42 to push a portion of
both the sensor 4) and the insertion device 120 through the skin of
the patient and into the subcutaneous tissue. The insertion device
120 is optionally pulled out of the skin and subcutaneous tissue
with the sensor 42 remaining in the subcutaneous tissue due to
frictional forces between the sensor 42 and the patient's tissue.
If the sensor 42 includes the optional barb 125, then this
structure may also facilitate the retention of the sensor 42 within
the interstitial tissue as the barb catches in the tissue.
[0221] The force applied to the insertion device 120 and/or the
sensor 42 may be applied manually or mechanically. Preferably, the
sensor 42 is reproducibly inserted through the skin of the patient.
In one embodiment, an insertion gun is used to insert the sensor.
One example of an insertion gun 200 for inserting a sensor 42 is
shown in FIG. 26. The insertion gun 200 includes a housing 202 and
a carrier 204. The insertion device 120 is typically mounted on the
carrier 204 and the sensor 42 is pre-loaded into the insertion
device 120. The carrier 204 drives the sensor 42 and, optionally,
the insertion device 120 into the skin of the patient using, for
example, a cocked or wound spring, a burst of compressed gas, an
electromagnet repelled by a second magnet, or the like, within the
insertion gun 200. In some instances, for example, when using a
spring, the carrier 204 and insertion device may be moved, cocked,
or otherwise prepared to be directed towards the skin of the
patient.
[0222] After the sensor 42 is inserted, the insertion gun 200 may
contain a mechanism which pulls the insertion device 120 out of the
skin of the patient. Such a mechanism may use a spring,
electromagnet, or the like to remove the insertion device 120.
[0223] The insertion gun may be reusable. The insertion device 120
is often disposable to avoid the possibility of contamination.
Alternatively, the insertion device 120 may be sterilized and
reused. In addition, the insertion device 120 and/or the sensor 42
may be coated with an anticlotting agent to prevent fouling of the
sensor 42.
[0224] In one embodiment, the sensor 42 is injected between 2 to 12
mm into the interstitial tissue of the patient for subcutaneous
implantation. Preferably, the sensor is injected 3 to 9 mm, and
more preferably 5 to 7 mm, into the interstitial tissue. Other
embodiments of the invention, may include sensors implanted in
other portions of the patient, including, for example, in an
artery, vein, or organ. The depth of implantation varies depending
on the desired implantation target.
[0225] Although the sensor 42 may be inserted anywhere in the body,
it is often desirable that the insertion site be positioned so that
the on-skin sensor control unit 44 can be concealed. In addition,
it is often desirable that the insertion site be at a place on the
body with a low density of nerve endings to reduce the pain to the
patient. Examples of preferred sites for insertion of the sensor 42
and positioning of the on-skin sensor control unit 44 include the
abdomen, thigh, leg, upper arm, and shoulder.
[0226] An insertion angle is measured from the plane of the skin
(i.e., inserting the sensor perpendicular to the skin would be a
90.degree. insertion angle). Insertion angles usually range from 10
to 90.degree., typically from 15 to 60.degree., and often from 30
to 45.degree..
[0227] On-Skin Sensor Control Unit
[0228] The on-skin sensor control unit 44 is configured to be
placed on the skin of a patient. The on-skin sensor control unit 44
is optionally formed in a shape that is comfortable to the patient
and which may permit concealment, for example, under a patient's
clothing. The thigh, leg, upper arm, shoulder, or abdomen are
convenient parts of the patient's body for placement of the on-skin
sensor control unit 44 to maintain concealment. However, the
on-skin sensor control unit 44 may be positioned on other portions
of the patient's body. One embodiment of the on-skin sensor control
unit 44 has a thin, oval shape to enhance concealment, as
illustrated in FIGS. 14-16. However, other shapes and sizes may be
used.
[0229] The particular profile, as well as the height, width,
length, weight, and volume of the on-skin sensor control unit 44
may vary and depends, at least in part, on the components and
associated functions included in the on-skin sensor control unit
44, as discussed below. For example, in some embodiments, the
on-skin sensor control unit 44 has a height of 1.3 cm or less, and
preferably 0.7 cm or less. In some embodiments, the on-skin sensor
control unit 44 has a weight of 90 grams or less, preferably 45
grams or less, and more preferably 25 grams or less. In some
embodiments, the on-skin sensor control unit 44 has a volume of
about 15 cm.sup.3 or less, preferably about 10 cm.sup.-1 or less,
more preferably about 5 cm.sup.3 or less, and most preferably about
2.5 cm.sup.3 or less.
[0230] The on-skin sensor control unit 44 includes a housing 45, as
illustrated in FIGS. 14-16. The housing 45 is typically formed as a
single integral unit that rests on the skin of the patient. The
housing 45 typically contains most or all of the electronic
components, described below, of the on-skin sensor control unit 44.
The on-skin sensor control unit 44 usually includes no additional
cables or wires to other electronic components or other devices. If
the housing includes two or more parts, then those parts typically
fit together to form a single integral unit.
[0231] The housing 45 of the on-skin sensor control unit 44,
illustrated in FIGS. 14-16, may be formed using a variety of
materials, including, for example, plastic and polymeric materials,
particularly rigid thermoplastics and engineering thermoplastics.
Suitable materials include, for example, polyvinyl chloride,
polyethylene, polypropylene, polystyrene, ABS polymers, and
copolymers thereof The housing 45 of the on-skin sensor control
unit 44 may be formed using a variety of techniques including, for
example, injection molding, compression molding, casting, and other
molding methods. Hollow or recessed regions may be formed in the
housing 45 of the on-skin sensor control unit 44. The electronic
components of the on-skin sensor control unit 44, described below,
and/or other items, such as a battery or a speaker for an audible
alarm, may be placed in the hollow or recessed areas.
[0232] In some embodiments, conductive contacts 80 are provided on
the exterior of the housing 45. In other embodiments, the
conductive contacts 80 are provided on the interior of the housing
45, for example, within a hollow or recessed region.
[0233] In some embodiments, the electronic components and/or other
items are incorporated into the housing 45 of the on-skin sensor
control unit 44 as the plastic or polymeric material is molded or
otherwise formed. In other embodiments, the electronic components
and/or other items are incorporated into the housing 45 as the
molded material is cooling or after the molded material has been
reheated to make it pliable. Alternatively, the electronic
components and/or other items may be secured to the housing 45
using fasteners, such as screws, nuts and bolts, nails, staples,
rivets, and the like or adhesives, such as contact adhesives,
pressure sensitive adhesives, glues, epoxies, adhesive resins, and
the like. In some cases, the electronic components and/or other
items are not affixed to the housing 45 at all.
[0234] In some embodiments, the housing 45 of the on-skin sensor
control unit 44 is a single piece. The conductive contacts 80 may
be formed on the exterior of the housing 45 or on the interior of
the housing 45 provided there is a port 78 in the housing 45
through which the sensor 42 can be directed to access the
conductive contacts 80.
[0235] In other embodiments, the housing 45 of the on-skin sensor
control unit 44 is formed in at least two separate portions that
fit together to form the housing 45, for example, a base 74 and a
cover 76, as illustrated in FIGS. 14-16. The two or more portions
of the housing 45 may be entirely separate from each other.
Alternatively, at least some of the two or more portions of the
housing 45 may be connected together, for example, by a hinge, to
facilitate the coupling of the portions to form the housing 45 of
the on-skin sensor control unit 44.
[0236] These two or more separate portions of the housing 45 of the
on-skin sensor control unit 44 may have complementary, interlocking
structures, such as, for example, interlocking ridges or a ridge on
one component and a complementary groove on another component, so
that the two or more separate components may be easily and/or
firmly coupled together. This may be useful, particularly if the
components are taken apart and fit together occasionally, for
example, when a battery or sensor 42 is replaced. However, other
fasteners may also be used to couple the two or more components
together, including, for example, screws, nuts and bolts, nails,
staples, rivets, or the like. In addition, adhesives, both
permanent or temporary, may be used including, for example, contact
adhesives, pressure sensitive adhesives, glues, epoxies, adhesive
resins, and the like.
[0237] Typically, the housing 45 is at least water resistant to
prevent the flow of fluids into contact with the components in the
housing, including, for example, the conductive contacts 80.
Preferably, the housing is waterproof. In one embodiment, two or
more components of the housing 45, for example, the base 74 and the
cover 76, fit together tightly to form a hermetic, waterproof, or
water resistant seal so that fluids can not flow into the interior
of the on-skin sensor control unit 44. This may be useful to avoid
corrosion currents and/or degradation of items within the on-skin
sensor control unit 44, such as the conductive contacts, the
battery, or the electronic components, particularly when the
patient engages in such activities as showering, bathing, or
swimming.
[0238] Water resistant, as used herein, means that there is no
penetration of water through a water resistant seal or housing when
immersed in water at a depth of one meter at sea level. Waterproof,
as used herein, means that there is no penetration of water through
the waterproof seal or housing when immersed in water at a depth of
ten meters, and preferably fifty meters, at sea level. It is often
desirable that the electronic circuitry, power supply (e.g.,
battery), and conductive contacts of the on-skin sensor control
unit, as well as the contact pads of the sensor, are contained in a
water resistant, and preferably, a waterproof, environment.
[0239] In addition to the portions of the housing 45, such as the
base 74 and cover 76, there may be other individually-formed pieces
of the on-skin sensor control unit 44, which may be assembled
during or after manufacture. One example of an individually-formed
piece is a cover for electronic components that fits a recess in
the base 74 or cover 76. Another example is a cover for a battery
provided in the base 74 or cover 76. These individually-formed
pieces of the on-skin sensor control unit 44 may be permanently
affixed, such as, for example, a cover for electronic components,
or removably affixed, such as, for example, a removable cover for a
battery, to the base 74, cover 76, or other component of the
on-skin sensor control unit 44. Methods for affixing these
individually-formed pieces include the use of fasteners, such as
screws, nuts and bolts, staples, nails, rivets, and the like,
frictional fasteners, such as tongue and groove structures, and
adhesives, such as contact adhesives, pressure sensitive adhesives,
glues, epoxies, adhesive resins, and the like.
[0240] One embodiment of the on-skin sensor control unit 44 is a
disposable unit complete with a battery for operating the unit.
There are no portions of the unit that the patient needs to open or
remove, thereby reducing the size of the unit and simplifying its
construction. The on-skin sensor control unit 44 optionally remains
in a sleep mode prior to use to conserve the battery's power. The
on-skin sensor control unit 44 detects that it is being used and
activates itself. Detection of use may be through a number of
mechanisms. These include, for example, detection of a change in
resistance across the electrical contacts, actuation of a switch
upon mating the on-skin sensor control unit 44 with a mounting unit
77 (see FIGS. 27A and 28A). The on-skin sensor control unit 44 is
typically replaced when it no longer operates within threshold
limits, for example, if the battery or other power source does not
generate sufficient power. Often this embodiment of the on-skin
sensor control unit 44 has conductive contacts 80 on the exterior
of the housing 45. Once the sensor 42 is implanted in the patient,
the sensor control unit 44 is placed over the sensor 42 with the
conductive contacts 80 in contact with the contact pads 49 of the
sensor 42.
[0241] The on-skin sensor control unit 44 is typically attached to
the skin 75 of the patient, as illustrated in FIG. 17. The on-skin
sensor control unit 44 may be attached by a variety of techniques
including, for example, by adhering the on-skin sensor control unit
44 directly to the skin 75 of the patient with an adhesive provided
on at least a portion of the housing 45 of the on-skin sensor
control unit 44 which contacts the skin 75 or by suturing the
on-skin sensor control unit 44 to the skin 75 through suture
openings (not shown) in the sensor control unit 44.
[0242] Another method of attaching the housing 45 of the on-skin
sensor control unit 44 to the skin 75 includes using a mounting
unit, 77. The mounting unit 77 is often a part of the on-skin
sensor control unit 44. One example of a suitable mounting unit 77
is a double-sided adhesive strip, one side of which is adhered to a
surface of the skin of the patient and the other side is adhered to
the on-skin sensor control unit 44. In this embodiment, the
mounting unit 77 may have an optional opening 79 which is large
enough to allow insertion of the sensor 42 through the opening 79.
Alternatively, the sensor may be inserted through a thin adhesive
and into the skin.
[0243] A variety of adhesives may be used to adhere the on-skin
sensor control unit 44 to the skin 75 of the patient, either
directly or using the mounting unit 77, including, for example,
pressure sensitive adhesives (PSA) or contact adhesives.
Preferably, an adhesive is chosen which is not irritating to all or
a majority of patients for at least the period of time that a
particular sensor 42 is implanted in the patient. Alternatively, a
second adhesive or other skin-protecting compound may be included
with the mounting unit so that a patient, whose skin is irritated
by the adhesive on the mounting unit 77, can cover his skin with
the second adhesive or other skin-protecting compound and then
place the mounting unit 77 over the second adhesive or other
skin-protecting compound. This should substantially prevent the
irritation of the skin of the patient because the adhesive on the
mounting unit 77 is no longer in contact with the skin, but is
instead in contact with the second adhesive or other
skin-protecting compound.
[0244] When the sensor 42 is changed, the on-skin sensor control
unit 44 may be moved to a different position on the skin 75 of the
patient, for example, to avoid excessive irritation. Alternatively,
the on-skin sensor control unit 44 may remain at the same place on
the skin of the patient until it is determined that the unit 44
should be moved.
[0245] Another embodiment of a mounting unit 77 used in an on-skin
sensor control unit 44 is illustrated in FIGS. 27A and 27B. The
mounting unit 77 and a housing 45 of an on-skin sensor control unit
44 are mounted together in, for example, an interlocking manner, as
shown in FIG. 27A. The mounting unit 77 is formed, for example,
using plastic or polymer materials, including, for example,
polyvinyl chloride, polyethylene, polypropylene, polystyrene, ABS
polymers, and copolymers thereof. The mounting unit 77 may be
formed using a variety of techniques including, for example,
injection molding, compression molding, casting, and other molding
methods.
[0246] The mounting unit 77 typically includes an adhesive on a
bottom surface of the mounting unit 77 to adhere to the skin of the
patient or the mounting unit 77 is used in conjunction with, for
example, double-sided adhesive tape or the like. The mounting unit
77 typically includes an opening 79 through which the sensor 42 is
inserted, as shown in FIG. 27B. The mounting unit 77 may also
include a support structure 220 for holding the sensor 42 in place
and against the conductive contacts 80 on the on-skin sensor
control unit 42. The mounting unit 77, also, optionally, includes a
positioning structure 222, such as an extension of material from
the mounting unit 77, that corresponds to a structure (not shown),
such as an opening, on the sensor 42 to facilitate proper
positioning of the sensor 42, for example, by aligning the two
complementary structures.
[0247] In another embodiment, a coupled mounting unit 77 and
housing 45 of an on-skin sensor control unit 44 is provided on an
adhesive patch 204 with an optional over 206 to protect and/or
confine the housing 45 of the on-skin sensor control unit 44, as
illustrated in FIG. 28A. The optional cover may contain an adhesive
or other mechanism for attachment to the housing 45 and/or mounting
unit 77. The mounting unit 77 typically includes an opening 49
through which a sensor 42 is disposed, as shown in FIG. 28B. The
opening 49 may optionally be configured to allow insertion of the
sensor 42 through the opening 49 using an insertion device 120 or
insertion gun 200 (see FIG. 26). The housing 45 of the on-skin
sensor control unit 44 has a base 74 and a cover 76, as illustrated
in FIG. 28C. A bottom view of the housing 45, as shown in FIG. 28D,
illustrates ports 230 through which conductive contacts (not shown)
extend to connect with contact pads on the sensor 42. A board 232
for attachment of circuit components may optionally be provided
within the on-skin sensor control unit 44, as illustrated in FIG.
28E.
[0248] In some embodiments, the adhesive on the on-skin sensor
control unit 44 and/or on any of the embodiments of the mounting
unit 77 is water resistant or waterproof to permit activities such
as showering and/or bathing while maintaining adherence of the
on-skin sensor control unit 44 to the skin 75 of the patient and,
at least in some embodiments, preventing water from penetrating
into the sensor control unit 44. The use of a water resistant or
waterproof adhesive combined with a water resistant or waterproof
housing 45 protects the components in the sensor control unit 44
and the contact between the conductive contacts 80 and the sensor
42 from damage or corrosion. An example of a non-irritating
adhesive that repels water is Tegaderm (3M, St. Paul, Minn.).
[0249] In one embodiment, the on-skin sensor control unit 44
includes a sensor port 78 through which the sensor 42 enters the
subcutaneous tissue of the patient, as shown in FIGS. 14 to 16. The
sensor 42 may be inserted into the subcutaneous tissue of the
patient through the sensor port 78. The on-skin sensor control unit
44 may then be placed on the skin of the patient with the sensor 42
being threaded through the sensor port 78. If the housing 45 of the
sensor 42 has, for example, a base 74 and a cover 76, then the
cover 76 may be removed to allow the patient to guide the sensor 42
into the proper position for contact with the conductive contacts
80.
[0250] Alternatively, if the conductive contacts 80 are within the
housing 45 the patient may slide the sensor 42 into the housing 45
until contact is made between the contact pads 49 and the
conductive contacts 80. The sensor control unit 44 may have a
structure which obstructs the sliding of the sensor 42 further into
the housing once the sensor 42 is properly positioned with the
contact pads 49 in contact with the conductive contacts 80.
[0251] In other embodiments, the conductive contacts 80 are on the
exterior of the housing 45 (see e.g., FIGS. 27A-27B and 28A-28E).
In these embodiments, the patient guides the contacts pads 49 of
the sensor 42 into contact with the conductive contacts 80. In some
cases, a guiding structure may be provided on the housing 45 which
guides the sensor 42 into the proper position. An example of such a
structure includes a set of guiding rails extending from the
housing 45 and having the shape of the sensor 42.
[0252] In some embodiments, when the sensor 42 is inserted using an
insertion device 120 (see FIG. 12), the tip of the insertion device
120 or optional insertion gun 200 (see FIG. 26) is positioned
against the skin or the mounting unit 77 at the desired insertion
point. In some embodiments, the insertion device 120 is positioned
on the skin without any guide. In other embodiments, the insertion
device 120 or insertion gun 200 is positioned using guides (not
shown) in the mounting unit 77 or other portion of the on-skin
sensor control unit 44. In some embodiments, the guides, opening 79
in the mounting unit 77 and/or sensor port 78 in the housing 45 of
the on-skin sensor control unit 44 have a shape which is
complementary to the shape of the tip of the insertion device 120
and/or insertion gun 200 to limit the orientation of the insertion
device 120 and/or insertion gun 200 relative to the opening 79
and/or sensor port 78. The sensor can then be subcutaneously
inserted into the patient by matching the complementary shape of
the opening 79 or sensor port 78 with the insertion device 120
and/or insertion gun 200.
[0253] In some embodiments, the shapes of a) the guides, opening
79, or sensor port 78, and (b) the insertion device 120 or
insertion gun 200 are configured such that the two shapes can only
be matched in a single orientation. This aids in inserting the
sensor 42 in the same orientation each time a new sensor is
inserted into the patient. This uniformity in insertion orientation
may be required in some embodiments to ensure that the contact pads
49 on the sensor 42 are correctly aligned with appropriate
conductive contacts 80 on the on-skin sensor control unit 44. In
addition, the use of the insertion gun, as described above, may
ensure that the sensor 42 is inserted at a uniform, reproducible
depth.
[0254] The sensor 42 and the electronic components within the
on-skin sensor control unit 44 are coupled via conductive contacts
80, as shown in FIGS. 14-16. The one or more working electrodes 58,
counter electrode 60 (or counter/reference electrode), optional
reference electrode 62, and optional temperature probe 66 are
attached to individual conductive contacts 80. In the illustrated
embodiment of FIGS. 14-16, the conductive contacts 80 are provided
on the interior of the on-skin sensor control unit 44. Other
embodiments of the on-skin sensor control unit 44 have the
conductive contacts disposed on the exterior of the housing 45. The
placement of the conductive contacts 80 is such that they are in
contact with the contact pads 49 on the sensor 42 when the sensor
42 is properly positioned within the on-skin sensor control unit
44.
[0255] In the illustrated embodiment of FIGS. 14-16, the base 74
and cover 76 of the on-skin sensor control unit 44 are formed such
that, when the sensor 42 is within the on-skin sensor control unit
44 and the base 74 and cover 76 are fitted together, the sensor 42
is bent. In this manner, the contact pads 49 on the sensor 42 are
brought into contact with the conductive contacts 80 of the on-skin
sensor control unit 44. The on-skin sensor control unit 44 may
optionally contain a support structure 82 to hold, support, and/or
guide the sensor 42 into the correct position.
[0256] Non-limiting examples of suitable conductive contacts 80 are
illustrated in FIGS. 19A-19D. In one embodiment, the conductive
contacts 80 are pins 84 or the like, as illustrated in FIG. 19A,
which are brought into contact with the contact pads 49 on the
sensor 42 when the components of the on-skin sensor control unit
44, for example, the base 74 and cover 76, are fitted together. A
support 82 may be provided under the sensor 42 to promote adequate
contact between the contact pads 49 on the sensor 42 and the pins
84. The pins are typically made using a conductive material, such
as a metal or alloy, for example, copper, stainless steel, or
silver. Each pin has a distal end that extends from the on-skin
sensor control unit 44 for contacting the contact pads 49 on the
sensor 42. Each pin 84 also has a proximal end that is coupled to a
wire or other conductive strip that is, in turn, coupled to the
rest of the electronic components (e.g., the voltage source 95 and
measurement circuit 96 of FIGS. 18A and 18B) within the on-skin
sensor control unit 44. Alternatively, the pins 84 may be coupled
directly to the rest of the electronics.
[0257] In another embodiment, the conductive contacts 80 are formed
as a series of conducting regions 88 with interspersed insulating
regions 90, as illustrated in FIG. 19B. The conducting regions 88
may be as large or larger than the contact pads 49 on the sensor 42
to alleviate registration concerns. However, the insulating regions
90 should have sufficient width so that a single conductive region
88 does not overlap with two contact pads 49 as determined based on
the expected variation in the position of the sensor 42 and contact
pads 49 with respect to the conductive contacts 80. The conducting
regions 88 are formed using materials such as metals, alloys, or
conductive carbon. The insulating regions 90 may be formed using
known insulating materials including, for example, insulating
plastic or polymer materials.
[0258] In a further embodiment, a unidirectional conducting
adhesive 92 may be used between the contact pads 49 on the sensor
42 and conductive contacts 80 implanted or otherwise formed in the
on-skin sensor control unit 44, as shown in FIG. 19C.
[0259] In yet another embodiment, the conductive contacts 80 are
conductive members 94 that extend from a surface of the on-skin
sensor control unit 44 to contact the contact pads 49, as shown in
FIG. 19D. A variety of different shapes may be used for these
members, however, they should be electrically insulated from each
other. The conductive members 94 may be made using metal, alloy,
conductive carbon, or conducting plastics and polymers.
[0260] Any of the exemplary conductive contacts 80 described above
may extend from either the upper surface of the interior of the
on-skin sensor control unit 44, as illustrated in FIG. 19A-19C, or
from the lower surface of the interior of the on-skin sensor
control unit 44, as illustrated in FIG. 19D, or from both the upper
and lower surfaces of the interior of the on-skin sensor control
unit 44, particularly when the sensor 42 has contact pads 49 on
both sides of the sensor.
[0261] Conductive contacts 80 on the exterior of the housing 45 may
also have a variety of shapes as indicated in FIGS. 19E and 19F.
For example, the conductive contacts 80 may be embedded in (FIG.
19E) or extending out of (FIG. 19F) the housing 45.
[0262] The conductive contacts 80 are preferably made using a
material which will not corrode due to contact with the contact
pads 49 of the sensor 42. Corrosion may occur when two different
metals are brought in contact. Thus, if the contact pads 49 are
formed using carbon then the preferred conductive contacts 80 may
be made using any material, including metals or alloys. However, if
any of the contact pads 49 are made with a metal or alloy then the
preferred conductive contacts 80 for coupling with the metallic
contact pads are made using a non-metallic conductive material,
such as conductive carbon or a conductive polymer, or the
conductive contacts 80 and the contact pads 49 are separated by a
non-metallic material, such as a unidirectional conductive
adhesive.
[0263] In one embodiment, electrical contacts are eliminated
between the sensor 42 and the on-skin sensor control unit 44. Power
is transmitted to the sensor via inductive coupling, using, for
example, closely space antennas (e.g., facing coils) (not shown) on
the sensor and the on-skin sensor control unit. Changes in the
electrical characteristics of the sensor control unit 44 (e.g.,
current) induce a changing magnetic field in the proximity of the
antenna. The changing magnetic field induces a current in the
antenna of the sensor. The close proximity of the sensor and
on-skin sensor control unit results in reasonably efficient power
transmission. The induced current in the sensor may be used to
power potentiostats, operational amplifiers, capacitors, integrated
circuits, transmitters, and other electronic components built into
the sensor structure. Data is transmitted back to the sensor
control unit, using, for example, inductive coupling via the same
or different antennas and/or transmission of the signal via a
transmitter on the sensor. The use of inductive coupling can
eliminate electrical contacts between the sensor and the on-skin
sensor control unit. Such contacts are commonly a source of noise
and failure. Moreover, the sensor control unit may then be entirely
sealed which may increase the waterproofing of the on-skin sensor
control unit.
[0264] An exemplary on-skin sensor control unit 44 can be prepared
and used in the following manner. A mounting unit 77 having
adhesive on the bottom is applied to the skin. An insertion gun 200
(see FIG. 26) carrying the sensor 42 and the insertion device 120
is positioned against the mounting unit 77. The insertion gun 200
and mounting unit 77 are optionally designed such that there is
only one position in which the two properly mate. The insertion gun
200 is activated and a portion of the sensor 42 and optionally a
portion of the insertion device 120 are driven through the skin
into, for example, the subcutaneous tissue. The insertion gun 200
withdraws the insertion device 200, leaving the portion of the
sensor 42 inserted through the skin. The housing 45 of the on-skin
control unit 44 is then coupled to the mounting unit 77.
Optionally, the housing 45 and the mounting unit 77 are formed such
that there is only one position in which the two properly mate. The
mating of the housing 45 and the mounting unit 77 establishes
contact between the contact pads 49 (see e.g., FIG. 2) on the
sensor 42 and the conductive contacts 80 on the on-skin sensor
control unit 44. Optionally, this action activates the on-skin
sensor control unit 44 to begin operation.
[0265] On-Skin Control Unit Electronics
[0266] The on-skin sensor control unit 44 also typically includes
at least a portion of the electronic components that operate the
sensor 42 and the analyte monitoring device system 40. One
embodiment of the electronics in the on-skin control unit 44 is
illustrated as a block diagram in FIG. 18A. The electronic
components of the on-skin sensor control unit 44 typically include
a power supply 95 for operating the on-skin control unit 44 and the
sensor 42, a sensor circuit 97 for obtaining signals from and
operating the sensor 42, a measurement circuit 96 that converts
sensor signals to a desired format, and a processing circuit 109
that, at minimum, obtains signals from the sensor circuit 97 and/or
measurement circuit 96 and provides the signals to an optional
transmitter 98. In some embodiments, the processing circuit 109 may
also partially or completely evaluate the signals from the sensor
42 and convey the resulting data to the optional transmitter 98
and/or activate an optional alarm system 94 (see FIG. 18B) if the
analyte level exceeds a threshold. The processing circuit 109 often
includes digital logic circuitry.
[0267] The on-skin sensor control unit 44 may optionally contain a
transmitter 98 for transmitting the sensor signals or processed
data from the processing circuit 109 to a receiver/display unit 46,
48; a data storage unit 102 for temporarily or permanently storing
data from the processing circuit 109; a temperature probe circuit
99 for receiving signals from and operating a temperature probe 66;
a reference voltage generator 101 for providing a reference voltage
for comparison with sensor-generated signals; and/or a watchdog
circuit 103 that monitors the operation of the electronic
components in the on-skin sensor control unit 44.
[0268] Moreover, the sensor control unit 44 often includes digital
and/or analog components utilizing semiconductor devices, such as
transistors. To operate these semiconductor devices, the on-skin
control unit 44 may include other components including, for
example, a bias control generator 105 to correctly bias analog and
digital semiconductor devices, an oscillator 107 to provide a clock
signal, and a digital logic and timing component 109 to provide
timing signals and logic operations for the digital components of
the circuit.
[0269] As an example of the operation of these components, the
sensor circuit 97 and the optional temperature probe circuit 99
provide raw signals from the sensor 42 to the measurement circuit
96. The measurement circuit 96 converts the raw signals to a
desired format, using for example, a current-to-voltage converter,
current-to-frequency converter, and/or a binary counter or other
indicator that produces a signal proportional to the absolute value
of the raw signal. This may be used, for example, to convert the
raw signal to a format that can be used by digital logic circuits.
The processing circuit 109 may then, optionally, evaluate the data
and provide commands to operate the electronics.
[0270] FIG. 18B illustrates a block diagram of another exemplary
on-skin control unit 44 that also includes optional components such
as a receiver 99 to receive, for example, calibration data; a
calibration storage unit 100 to hold, for example, factory-set
calibration data, calibration data obtained via the receiver 99
and/or operational signals received, for example, from a
receiver/display unit 46, 48 or other external device; an alarm
system 104 for warning the patient; and a deactivation switch 111
to turn off the alarm system.
[0271] Functions of the analyte monitoring system 40 and the sensor
control unit 44 may be implemented using either software routines,
hardware components, or combinations thereof. The hardware
components may be implemented using a variety of technologies,
including, for example, integrated circuits or discrete electronic
components. The use of integrated circuits typically reduces the
size of the electronics, which in turn may result in a smaller
on-skin sensor control unit 44.
[0272] The electronics in the on-skin sensor control unit 44 and
the sensor 42 are operated using a power supply 95. One example of
a suitable power supply 95 is a battery, for example, a thin
circular battery, such as those used in many watches, hearing aids,
and other small electronic devices. Preferably, the battery has a
lifetime of at least 30 days, more preferably, a lifetime of at
least three months, and most preferably, a lifetime of at least one
year. The battery is often one of the largest components in the
on-skin control unit 44, so it is often desirable to minimize the
size of the battery. For example, a preferred battery's thickness
is 0.5 mm or less, preferably 0.35 mm or less, and most preferably
0.2 mm or less. Although multiple batteries may be used, it is
typically preferred to use only one battery.
[0273] The sensor circuit 97 is coupled via the conductive contacts
80 of the sensor control unit 44 to one or more sensors 42, 42'.
Each of the sensors represents, at minimum, a working electrode 58,
a counter electrode 60 (or counter/reference electrode), and an
optional reference electrode 62. When two or more sensors 42, 42'
are used, the sensors typically have individual working electrodes
58, but may share a counter electrode 60, counter/reference
electrode, and/or reference electrode 52.
[0274] The sensor circuit 97 receives signals from and operates the
sensor 42 or sensors 42, 42'. The sensor circuit 97 may obtain
signals from the sensor 42 using amperometric, coulometric,
potentiometric, voltammetric, and/or other electrochemical
techniques. The sensor circuit 97 is exemplified herein as
obtaining amperometric signals from the sensor 42, however, it will
be understood that the sensor circuit can be appropriately
configured for obtaining signals using other electrochemical
techniques. To obtain amperometric measurements, the sensor circuit
97 typically includes a potentiostat that provides a constant
potential to the sensor 42. In other embodiments, the sensor
circuit 97 includes an amperostat that supplies a constant current
to the sensor 42 and can be used to obtain coulometric or
potentiometric measurements.
[0275] The signal from the sensor 42 generally has at least one
characteristic, such as, for example, current, voltage, or
frequency, which varies with the concentration of the analyte. For
example, if the sensor circuit 97 operates using amperometry, then
the signal current varies with analyte concentration. The
measurement circuit 96 may include circuitry which converts the
information-carrying portion of the signal from one characteristic
to another. For example, the measurement circuit 96 may include a
current-to-voltage or current-to-frequency converter. The purpose
of this conversion may be to provide a signal that is, for example,
more easily transmitted, readable by digital circuits, and/or less
susceptible to noise contributions.
[0276] One example of a standard current-to-voltage converter is
provided in FIG. 20A. In this converter, the signal from the sensor
42 is provided at one input terminal 134 of an operational
amplifier 130 ("op amp") and coupled through a resistor 138 to an
output terminal 136. This particular current-to-voltage converter
131 may, however, be difficult to implement in a small CMOS chip
because resistors are often difficult to implement on an integrated
circuit. Typically, discrete resistor components are used. However,
the used of discrete components increases the space needed for the
circuitry.
[0277] An alternative current-to-voltage converter 141 is
illustrated in FIG. 20B. This converter includes an op amp 140 with
the signal from the sensor 42 provided at input terminal 144 and a
reference potential provided at input terminal 142. A capacitor 145
is placed between the input terminal 144 and the output terminal
146. In addition, switches 147a, 147b, 149a, and 149b are provided
to allow the capacitor to charge and discharge at a rate determined
by a clock (CLK) frequency. In operation, during one half cycle,
switches 147a and 147b close and switches 149a and 149b open
allowing the capacitor 145 to charge due to the attached potential
VI. During the other half cycle, switches 147a and 147b open and
switches 149a and 149b close to ground and allow the capacitor 145
to partially or fully discharge. The reactive impedance of the
capacitor 145 is analogous to the resistance of the resistor 138
(see FIG. 20A), allowing the capacitor 145 to emulate a resistor.
The value of this "resistor" depends on the capacitance of the
capacitor 145 and the clock frequency. By altering the clock
frequency, the reactive impedance ("resistance value") of the
capacitor changes. The value of the impedance ("resistance") of the
capacitor 145 may be altered by changing the clock frequency.
Switches 147a, 147b. 149a, and 149b may be implemented in a CMOS
chip using, for example, transistors.
[0278] A current-to-frequency converter may also be used in the
measurement circuit 96. One suitable current-to-frequency converter
includes charging a capacitor using the signal from the sensor 42.
When the potential across the capacitor exceeds a threshold value,
the capacitor is allowed to discharge. Thus, the larger the current
from the sensor 42, the quicker the threshold potential is
achieved. This results in a signal across the capacitor that has an
alternating characteristic, corresponding to the charging and
discharging of the capacitor, having a frequency which increases
with an increase in current from the sensor 42.
[0279] In some embodiments, the analyte monitoring system 40
includes two or more working electrodes 58 distributed over one or
more sensors 42. These working electrodes 58 may be used for
quality control purposes. For example, the output signals and/or
analyzed data derived using the two or more working electrodes 58
may be compared to determine if the signals from the working
electrodes agree within a desired level of tolerance. If the output
signals do not agree, then the patient may be alerted to replace
the sensor or sensors. In some embodiments, the patient is alerted
only if the lack of agreement between the two sensors persists for
a predetermined period of time. The comparison of the two signals
may be made for each measurement or at regular intervals.
Alternatively or additionally, the comparison may be initiated by
the patient or another person. Moreover, the signals from both
sensors may be used to generate data or one signal may be discarded
after the comparison.
[0280] Alternatively, if, for example, two working electrodes 58
have a common counter electrode 60 and the analyte concentration is
measured by amperometry, then the current at the counter electrode
60 should be twice the current at each of the working electrodes,
within a predetermined tolerance level, if the working electrodes
are operating properly. If not, then the sensor or sensors should
be replaced, as described above.
[0281] An example of using signals from only one working electrode
for quality control includes comparing consecutive readings
obtained using the single working electrode to determine if they
differ by more than a threshold level. If the difference is greater
than the threshold level for one reading or over a period of time
or for a predetermined number of readings within a period of time
then the patient is alerted to replace the sensor 42. Typically,
the consecutive readings and/or the threshold level are determined
such that all expected excursions of the sensor signal are within
the desired parameters (i.e. the sensor control unit 44 does not
consider true changes in analyte concentration to be a sensor
failure).
[0282] The sensor control unit 44 may also optionally include a
temperature probe circuit 99. The temperature probe circuit 99
provides a constant current through (or constant potential) across
the temperature probe 66. The resulting potential (or current)
varies according to the resistance of the temperature dependent
element 72.
[0283] The output from the sensor circuit 97 and optional,
temperature probe circuit is coupled into a measurement circuit 96
that obtains signals from the sensor circuit 97 and optional
temperature probe circuit 99 and, at least in some embodiments,
provides output data in a form that, for example can be read by
digital circuits. The signals from the measurement circuit 96 are
sent to the processing circuit 109, which in turn may provide data
to an optional transmitter 98. The processing circuit 109 may have
one or more of the following functions: 1) transfer the signals
from the measurement circuit 96 to the transmitter 98, 2) transfer
signals from the measurement circuit 96 to the data storage circuit
102, 3) convert the information-carrying characteristic of the
signals from one characteristic to another (when, for example, that
has not been done by the measurement circuit 96), using, for
example, a current-to-voltage converter, a current-to-frequency
converter, or a voltage-to-current converter, 4) modify the signals
from the sensor circuit 97 using calibration data and/or output
from the temperature probe circuit 99, 5) determine a level of an
analyte in the interstitial fluid, 6) determine a level of an
analyte in the bloodstream based on the sensor signals obtained
from interstitial fluid, 7) determine if the level, rate of change,
and/or acceleration in the rate of change of the analyte exceeds or
meets one or more threshold values, 8) activate an alarm if a
threshold value is met or exceeded, 9) evaluate trends in the level
of an analyte based on a series of sensor signals, 10) determine a
dose of a medication, and 11) reduce noise and/or errors, for
example, through signal averaging or comparing readings from
multiple working electrodes 58.
[0284] The processing circuit 109 may be simple and perform only
one or a small number of these functions or the processing circuit
109 may be more sophisticated and perform all or most of these
functions. The size of the on-skin sensor control unit 44 may
increase with the increasing number of functions and complexity of
those functions that the processing circuit 109 performs. Many of
these functions may not be performed by a processing circuit 109 in
the on-skin sensor control unit 44, but may be performed by another
analyzer 152 in the receiver/display units 46, 48 (see FIG.
22).
[0285] One embodiment of the measurement circuit 96 and/or
processing circuit 109 provides as output data, the current flowing
between the working electrode 58 and the counter electrode 60. The
measurement circuit 96 and/or processing circuit 109 may also
provide as output data a signal from the optional temperature probe
66 which indicates the temperature of the sensor 42. This signal
from the temperature probe 66 may be as simple as a current through
the temperature probe 66 or the processing circuit 109 may include
a device that determines a resistance of the temperature probe 66
from the signal obtained from the measurement circuit 96 for
correlation with the temperature of the sensor 42. The output data
may then be sent to a transmitter 98 that then transmits this data
to at least one receiver/display device 46,48.
[0286] Returning to the processing circuit 109, in some embodiments
processing circuit 109 is more sophisticated and is capable of
determining the analyte concentration or some measure
representative of the analyte concentration, such as a current or
voltage value. The processing circuit 109 may incorporate the
signal of the temperature probe to make a temperature correction in
the signal or analyzed data from the working electrode 58. This may
include, for example, scaling the temperature probe measurement and
adding or subtracting the scaled measurement to the signal or
analyzed data from the working electrode 58. The processing circuit
109 may also incorporate calibration data which has been received
from an external source or has been incorporated into the
processing circuit 109, both of which are described below, to
correct the signal or analyzed data from the working electrode 58.
Additionally, the processing circuit 109 may include a correction
algorithm for converting interstitial analyte level to blood
analyte level. The conversion of interstitial analyte level to
blood analyte level is described, for example, in Schmidtke, et al.
"Measurement and Modeling of the Transient Difference Between Blood
and Subcutaneous Glucose Concentrations in the Rat after Injection
of Insulin", Proc. of the Nat'l Acad. of Science, 95, 294-299
(1998) and Quinn. et al., "Kinetics of Glucose Delivery to
Subcutaneous Tissue in Rats Measured with 0.3 mm Amperometric
Microsensors", Am. J. Physiol., 269 (Endocrinol. Metab. 32),
E155-E161 (1995), incorporated herein by reference.
[0287] In some embodiments, the data from the processing circuit
109 is analyzed and directed to an alarm system 94 (see FIG. 18B)
to warn the user. In at least some of these embodiments, a
transmitter is not used as the sensor control unit performs all of
the needed functions including analyzing the data and warning the
patient.
[0288] However, in many embodiments, the data (e.g., a current
signal, a converted voltage or frequency signal, or fully or
partially analyzed data) from processing circuit 109 is transmitted
to one or more receiver/display units 46, 48 using a transmitter 98
in the on-skin sensor control unit 44. The transmitter has an
antenna 93, such as a wire or similar conductor, formed in the
housing 45. The transmitter 98 is typically designed to transmit a
signal up to about 2 meters or more, preferably up to about 5
meters or more, and more preferably up to about 10 meters or more,
when transmitting to a small receiver/display unit 46, such as a
palm-size, belt-worn receiver. The effective range is longer when
transmitting to a unit with a better antenna, such as a bedside
receiver. As described in detail below, suitable examples of
receiver/display units 46, 48 include units that can be easily worn
or carried or units that can be placed conveniently on, for
example, a nightstand when the patient is sleeping.
[0289] The transmitter 98 may send a variety of different signals
to the receiver/display units 46, 48, typically, depending on the
sophistication of the processing circuit 109. For example, the
processing circuit 109 may simply provide raw signals, for example,
currents from the working electrodes 58, without any corrections
for temperature or calibration, or the processing circuit 109 may
provide converted signals which are obtained, for example, using a
current-to-voltage converter 131 or 141 or a current-to-frequency
converter. The raw measurements or converted signals may then be
processed by an analyzer 152 (see FIG. 22) in the receiver/display
units 46, 48 to determine the level of an analyte, optionally using
temperature and calibration corrections. In another embodiment, the
processing circuit 109 corrects the raw measurements using, for
example, temperature and/or calibration information and then the
transmitter 98 sends the corrected signal, and optionally, the
temperature and/or calibration information, to the receiver/display
units 46, 48. In yet another embodiment, the processing circuit 109
calculates the analyte level in the interstitial fluid and/or in
the blood (based on the interstitial fluid level) and transmits
that information to the one or more receiver/display units 46, 48,
optionally with any of the raw data and/or calibration or
temperature information. In a further embodiment, the processing
circuit 109 calculates the analyte concentration, but the
transmitter 98 transmits only the raw measurements, converted
signals, and/or corrected signals.
[0290] One potential difficulty that may be experienced with the
on-skin sensor control unit 44 is a change in the transmission
frequency of the transmitter 98 over time. To overcome this
potential difficulty, the transmitter may include optional
circuitry that can return the frequency of the transmitter 98 to
the desired frequency or frequency band. One example of suitable
circuitry is illustrated in FIG. 21 as a block diagram of an open
loop modulation system 200. The open loop modulation system 200
includes a phase detector (PD) 210, a charge pump (CHGPMP) 212, a
loop filter (L.F.) 214, a voltage controlled oscillator (VCO) 216,
and a divide by M circuit (.div.M) 218 to form the phase-locked
loop 220.
[0291] The analyte monitoring device 40 uses an open loop
modulation system 200 for RF communication between the transmitter
98 and a receiver of, for example, the one or more receiver/display
units 46, 48. This open loop modulation system 230 is designed to
provide a high reliability RF link between a transmitter and its
associated receiver. The system employs frequency modulation (FM),
and locks the carrier center frequency using a conventional
phase-locked loop (PLL) 220. In operation, the phase-locked loop
220 is opened prior to the modulation. During the modulation the
phase-locked loop 220 remains open for as long as the center
frequency of the transmitter is within the receiver's bandwidth.
When the transmitter detects that the center frequency is going to
move outside of the receiver bandwidth, the receiver is signaled to
stand by while the center frequency is captured. Subsequent to the
capture, the transmission will resume. This cycle of capturing the
center frequency, opening the phase-locked loop 220, modulation,
and recapturing the center frequency will repeat for as many cycles
as required.
[0292] The loop control 240 detects the lock condition of the
phase-locked loop 220 and is responsible for closing and opening
the phase-locked loop 220. The totalizer 250 in conjunction with
the loop control 240, detects the status of the center frequency.
The modulation control 230 is responsible for generating the
modulating signal. A transmit amplifier 260 is provided to ensure
adequate transmit signal power. The reference frequency is
generated from a very stable signal source (not shown), and is
divided down by N through the divide by N block (.div.N) 270. Data
and control signals are received by the open loop modulation system
200 via the DATA BUS 280, and the CONTROL BUS 290.
[0293] The operation of the open loop modulation system 200 begins
with the phase-locked loop 220 in closed condition. When the lock
condition is detected by the loop control 240, the phase-locked
loop 220 is opened and the modulation control 230 begins generating
the modulating signal. The totalizer 250 monitors the VCO frequency
(divided by M), for programmed intervals. The monitored frequency
is compared to a threshold programmed in the totalizer 250. This
threshold corresponds to the 3 dB cut off frequencies of the
receiver's intermediate frequency stage. When the monitored
frequency approaches the thresholds, the loop control 240 is
notified and a stand-by code is transmitted to the receiver and the
phase-locked loop 220 is closed.
[0294] At this point the receiver is in the wait mode. The loop
control 240 in the transmitter closes the phase-locked loop 220.
Then, modulation control 230 is taken off line, the monitored value
of the totalizer 250 is reset, and the phase-locked loop 220 is
locked. When the loop control 240 detects a lock condition, the
loop control 240 opens the phase-locked loop 220, the modulation
control 230 is brought on line and the data transmission to the
receiver will resume until the center frequency of the phase-locked
loop 220 approaches the threshold values, at which point the cycle
of transmitting the stand-by code begins. The .div.N 270 and .div.M
218 block set the frequency channel of the transmitter.
[0295] Accordingly, the open loop modulation system 200 provides a
reliable low power FM data transmission for an analyte monitoring
system. The open loop modulation system 200 provides a method of
wide band frequency modulation, while the center frequency of the
carrier is kept within receiver bandwidth. The effect of parasitic
capacitors and inductors pulling the center frequency of the
transmitter is corrected by the phase-locked loop 220. Further, the
totalizer 250 and loop control 240 provide a new method of center
frequency drift detection. Finally, the open loop modulation system
200 is easily implemented in CMOS process.
[0296] The rate at which the transmitter 98 transmits data may be
the same rate at which the sensor circuit 97 obtains signals and/or
the processing circuit 109 provides data or signals to the
transmitter 98. Alternatively, the transmitter 98 may transmit data
at a slower rate. In this case, the transmitter 98 may transmit
more than one datapoint in each transmission. Alternatively, only
one datapoint may be sent with each data transmission, the
remaining data not being transmitted. Typically, data is
transmitted to the receiver/display unit 46, 48 at least every
hour, preferably, at least every fifteen minutes, more preferably,
at least every five minutes, and most preferably, at least every
one minute. However, other data transmission rates may be used. In
some embodiments, the processing circuit 109 and/or transmitter 98
are configured to process and/or transmit data at a faster rate
when a condition is indicated, for example, a low level or high
level of analyte or impending low or high level of analyte. In
these embodiments, the accelerated data transmission rate is
typically at least every five minutes and preferably at least every
minute.
[0297] In addition to a transmitter 98, an optional receiver 99 may
be included in the on-skin sensor control unit 44. In some cases,
the transmitter 98 is a transceiver, operating as both a
transmitter and a receiver. The receiver 99 may be used to receive
calibration data for the sensor 42. The calibration data may be
used by the processing circuit 109 to correct signals from the
sensor 42. This calibration data may be transmitted by the
receiver/display unit 46, 48 or from some other source such as a
control unit in a doctor's office. In addition, the optional
receiver 99 may be used to receive a signal from the
receiver/display units 46, 48, as described above, to direct the
transmitter 98, for example, to change frequencies or frequency
bands, to activate or deactivate the optional alarm system 94 (as
described below), and/or to direct the transmitter 98 to transmit
at a higher rate.
[0298] Calibration data may be obtained in a variety of ways. For
instance, the calibration data may simply be factory-determined
calibration measurements which can be input into the on-skin sensor
control unit 44 using the receiver 99 or may alternatively be
stored in a calibration data storage unit 100 within the on-skin
sensor control unit 44 itself (in which case a receiver 99 may not
be needed). The calibration data storage unit 100 may be, for
example, a readable or readable/writeable memory circuit.
[0299] Alternative or additional calibration data may be provided
based on tests performed by a doctor or some other professional or
by the patient himself. For example, it is common for diabetic
individuals to determine their own blood glucose concentration
using commercially available testing kits. The results of this test
is input into the on-skin sensor control unit 44 either directly,
if an appropriate input device (e.g., a keypad, an optical signal
receiver, or a port for connection to a keypad or computer) is
incorporated in the on-skin sensor control unit 44, or indirectly
by inputting the calibration data into the receiver/display unit
46, 48 and transmitting the calibration data to the on-skin sensor
control unit 44.
[0300] Other methods of independently determining analyte levels
may also be used to obtain calibration data. This type of
calibration data may supplant or supplement factory-determined
calibration values.
[0301] In some embodiments of the invention, calibration data may
be required at periodic intervals, for example, every eight hours,
once a day, or, once a week, to confirm that accurate analyte
levels are being reported. Calibration may also be required each
time a new sensor 42 is implanted or if the sensor exceeds a
threshold minimum or maximum value or if the rate of change in the
sensor signal exceeds a threshold value. In some cases, it may be
necessary to wait a period of time after the implantation of the
sensor 42 before calibrating to allow the sensor 42 to achieve
equilibrium. In some embodiments, the sensor 42 is calibrated only
after it has been inserted. In other embodiments, no calibration of
the sensor 42 is needed.
[0302] The on-skin sensor control unit 44 and/or a receiver/display
unit 46, 48 may include an auditory or visual indicator that
calibration data is needed, based, for example, on a predetermined
periodic time interval between calibrations or on the implantation
of a new sensor 42. The on-skin sensor control unit 44 and/or
receiver display/units 46, 48 may also include an auditory or
visual indicator to remind the patient that information, such as
analyte levels, reported by the analyte monitoring device 40, may
not be accurate because a calibration of the sensor 42 has not been
performed within the predetermined periodic time interval and/or
after implantation of a new sensor 42.
[0303] The processing circuit 109 of the on-skin sensor control
unit 44 and/or an analyzer 152 of the receiver/display unit 46, 48
may determine when calibration data is needed and if the
calibration data is acceptable. The on-skin sensor control unit 44
may optionally be configured to not allow calibration or to reject
a calibration point if, for example, 1) a temperature reading from
the temperature probe indicates a temperature that is not within a
predetermined acceptable range (e.g., 30 to 42.degree. C. or 32 to
40.degree. C.) or that is changing rapidly (for example,
0.2.degree. C./minute, 0.5 .degree. C./minute, or 0.7.degree.
C./minute or greater); 2) two or more working electrodes 58 provide
uncalibrated signals that are not within a predetermined range
(e.g., within 10% or 20%) of each other; 3) the rate of change of
the uncalibrated signal is above a threshold rate (e.g., 0.25 mg/dL
per minute or 0.5 mg/dL per minute or greater); 4) the uncalibrated
signal exceeds a threshold maximum value (e.g., 5, 10, 20, or 40
nA) or is below a threshold minimum value (e.g., 0.05, 0.2, 0.5, or
1 nA); 5) the calibrated signal exceeds a threshold maximum value
(e.g., a signal corresponding to an analyte concentration of 200
mg/dL, 250 mg/dL, or 300 mg/dL) or is below a threshold minimum
value (e.g., a signal corresponding to an analyte concentration of
50 mg/dL, 65 mg/dL, or 80 mg/dL); and/or 6) an insufficient among
of time has elapsed since implantation (e.g., 10 minutes or less,
20 minutes or less, or 30 minutes or less).
[0304] The processing circuit 109 or an analyzer 152 may also
request another calibration point if the values determined using
the sensor data before and after the latest calibration disagree by
more than a threshold amount, indicating that the calibration may
be incorrect or that the sensor characteristics have changed
radically between calibrations. This additional calibration point
may indicate the source of the difference.
[0305] The on-skin sensor control unit 44 may include an optional
data storage unit 102 which may be used to hold data (e.g.,
measurements from the sensor or processed data) from the processing
circuit 109 permanently or, more typically, temporarily. The data
storage unit 102 may hold data so that the data can be used by the
processing circuit 109 to analyze and/or predict trends in the
analyte level, including, for example, the rate and/or acceleration
of analyte level increase or decrease. The data storage unit 102
may also or alternatively be used to store data during periods in
which a receiver/display unit 46, 48 is not within range. The data
storage unit 102 may also be used to store data when the
transmission rate of the data is slower than the acquisition rate
of the data. For example, if the data acquisition rate is 10
points/min and the transmission is 2 transmissions/min, then one to
five points of data could be sent in each transmission depending on
the desired rate for processing datapoints. The data storage unit
102 typically includes a readable/writeable memory storage device
and typically also includes the hardware and/or software to write
to and/or read the memory storage device.
[0306] The on-skin sensor control unit 44 may include an optional
alarm system 104 that, based on the data from the processing
circuit 109, warns the patient of a potentially detrimental
condition of the analyte. For example, if glucose is the analyte,
than the on-skin sensor control unit 44 may include an alarm system
104 that warns the patient of conditions such as hypoglycemia,
hyperglycemia, impending hypoglycemia, and/or impending
hyperglycemia. The alarm system 104 is triggered when the data from
the processing circuit 109 reaches or exceeds a threshold value.
Examples of threshold values for blood glucose levels are about 60,
70, or 80 mg/dL for hypoglycemia; about 70, 80, or 90 mg/dL for
impending hypoglycemia; about 130, 150, 175, 200, 225, 250, or 275
mg/dL for impending hyperglycemia; and about 150, 175, 200, 225,
250, 275, or 300 mg/dL for hyperglycemia. The actual threshold
values that are designed into the alarm system 104 may correspond
to interstitial fluid glucose concentrations or electrode
measurements (e.g., current values or voltage values obtained by
conversion of current measurements) that correlate to the
above-mentioned blood glucose levels. The analyte monitor device
may be configured so that the threshold levels for these or any
other conditions may be programmable by the patient and/or a
medical professional.
[0307] A threshold value is exceeded if the datapoint has a value
that is beyond the threshold value in a direction indicating a
particular condition. For example, a datapoint which correlates to
a glucose level of 200 mg/dL exceeds a threshold value for
hyperglycemia of 180 mg/dL, because the datapoint indicates that
the patient has entered a hyperglycemic state. As another example,
a datapoint which correlates to a glucose level of 65 mg/dL exceeds
a threshold value for hypoglycemia of 70 mg/dL because the
datapoint indicates that the patient is hypoglycemic as defined by
the threshold value. However, a datapoint which correlates to a
glucose level of 75 mg/dL would not exceed the same threshold value
for hypoglycemia because the datapoint does not indicate that
particular condition as defined by the chosen threshold value.
[0308] An alarm may also be activated if the sensor readings
indicate a value that is beyond a measurement range of the sensor
42. For glucose, the physiologically relevant measurement range is
typically about 50 to 250 mg/dL, preferably about 40-300 mg/dL and
ideally 30-400 mg/dL, of glucose in the interstitial fluid.
[0309] The alarm system 104 may also, or alternatively, be
activated when the rate of change or acceleration of the rate of
change in analyte level increase or decrease reaches or exceeds a
threshold rate or acceleration. For example, in the case of a
subcutaneous glucose monitor, the alarm system might be activated
if the rate of change in glucose concentration exceeds a threshold
value which might indicate that a hyperglycemic or hypoglycemic
condition is likely to occur.
[0310] The optional alarm system 104 may be configured to activate
when a single data point meets or exceeds a particular threshold
value. Alternatively, the alarm may be activated only when a
predetermined number of datapoints spanning a predetermined amount
of time meet or exceed the threshold value. As another alternative,
the alarm may be activated only when the datapoints spanning a
predetermined amount of time have an average value which meets or
exceeds the threshold value. Each condition that can trigger an
alarm may have a different alarm activation condition. In addition,
the alarm activation condition may change depending on current
conditions (e.g., an indication of impending hyperglycemia may
alter the number of datapoints or the amount of time that is tested
to determine hyperglycemia).
[0311] The alarm system 104 may contain one or more individual
alarms. Each of the alarms may be individually activated to
indicate one or more conditions of the analyte. The alarms may be,
for example, auditory or visual. Other sensory-stimulating alarm
systems may be used including alarm systems which heat, cool,
vibrate, or produce a mild electrical shock when activated. In some
embodiments, the alarms are auditory with a different tone, note,
or volume indicating different conditions. For example, a high note
might indicate hyperglycemia and a low note might indicate
hypoglycemia. Visual alarms may use a difference in color,
brightness, or position on the on-skin sensor control device 44 to
indicate different conditions. In some embodiments, an auditory
alarm system is configured so that the volume of the alarm
increases over time until the alarm is deactivated.
[0312] In some embodiments, the alarm may be automatically
deactivated after a predetermined time period. In other
embodiments, the alarm may be configured to deactivate when the
data no longer indicate that the condition which triggered the
alarm exists. In these embodiments, the alarm may be deactivated
when a single data point indicates that the condition no longer
exists or, alternatively, the alarm may be deactivated only after a
predetermined number of datapoints or an average of datapoints
obtained over a given period of time indicate that the condition no
longer exists.
[0313] In some embodiments, the alarm may be deactivated manually
by the patient or another person in addition to or as an
alternative to automatic deactivation. In these embodiments, a
switch 101 is provided which when activated turns off the alarm.
The switch 101 may be operatively engaged (or disengaged depending
on the configuration of the switch) by, for example, operating an
actuator on the on-skin sensor control unit 44 or the
receiver/display unit 46, 48. In some cases, an actuator may be
provided on two or more units 44, 46, 48, any of which may be
actuated to deactivate the alarm. If the switch 101 and or actuator
is provided on the receiver/display unit 46, 48 then a signal may
be transmitted from the receiver/display unit 46, 48 to the
receiver 98 on the on-skin sensor control unit 44 to deactivate the
alarm.
[0314] A variety of switches 101 may be used including, for
example, a mechanical switch, a reed switch, a Hall effect switch,
a Gigantic Magnetic Ratio (GMR) switch (the resistance of the GMR
switch is magnetic field dependent) and the like. Preferably, the
actuator used to operatively engage (or disengage) the switch is
placed on the on-skin sensor control unit 44 and configured so that
no water can flow around the button and into the housing. One
example of such a button is a flexible conducting strip that is
completely covered by a flexible polymeric or plastic coating
integral to the housing. In an open position the flexible
conducting strip is bowed and bulges away from the housing. When
depressed by the patient or another person, the flexible conducting
strip is pushed directly toward a metal contact and completes the
circuit to shut off the alarm.
[0315] For a reed or GMR switch, a piece of magnetic material, such
as a permanent magnet or an electromagnet, in a flexible actuator
that is bowed or bulges away from the housing 45 and the reed or
GMR switch is used. The reed or GMR switch is activated (to
deactivate the alarm) by depressing the flexible actuator bringing
the magnetic material closer to the switch and causing an increase
in the magnetic field within the switch.
[0316] In some embodiments of the invention, the analyte monitoring
device 40 includes only an on-skin control unit 44 and a sensor 42.
In these embodiments, the processing circuit 109 of the on-skin
sensor control unit 44 is able to determine a level of the analyte
and activate an alarm system 104 if the analyte level exceeds a
threshold. The on-skin control unit 44, in these embodiments, has
an alarm system 104 and may also include a display, such as those
discussed below with respect to the receiver/display units 46, 48.
Preferably, the display is an LCD or LED display. The on-skin
control unit 44 may not have a transmitter, unless, for example, it
is desirable to transmit data, for example, to a control unit in a
doctor's office.
[0317] The on-skin sensor control unit 44 may also include a
reference voltage generator 101 to provide an absolute voltage or
current for use in comparison to voltages or currents obtained from
or used with the sensor 42. An example of a suitable reference
voltage generator is a band-gap reference voltage generator that
uses, for example, a semiconductor material with a known band-gap.
Preferably, the band-gap is temperature insensitive over the range
of temperatures that the semiconductor material will experience
during operation. Suitable semiconductor materials includes
gallium, silicon and silicates.
[0318] A bias current generator 105 may be provided to correctly
bias solid-state electronic components. An oscillator 107 may be
provided to produce a clock signal that is typically used with
digital circuitry.
[0319] The on-skin sensor control unit 44 may also include a
watchdog circuit 103 that tests the circuitry, particularly, any
digital circuitry in the control unit 44 to determine if the
circuitry is operating correctly. Non-limiting examples of watchdog
circuit operations include: a) generation of a random number by the
watchdog circuit, storage of the number in a memory location,
writing the number to a register in the watchdog circuit, and
recall of the number to compare for equality; b) checking the
output of an analog circuit to determine if the output exceeds a
predetermined dynamic range; c) checking the output of a timing
circuit for a signal at an expected pulse interval. Other examples
of functions of a watchdog circuit are known in the art. If the
watchdog circuit detects an error that watchdog circuit may
activate an alarm and/or shut down the device.
[0320] Receiver/Display Unit
[0321] One or more receiver/display units 46, 48 may be provided
with the analyte monitoring device 40 for easy access to the data
generated by the sensor 42 and may, in some embodiments, process
the signals from the on-skin sensor control unit 44 to determine
the concentration or level of analyte in the subcutaneous tissue.
Small receiver/display units 46 may be carried by the patient.
These units 46 may be palm-sized and/or may be adapted to fit on a
belt or within a bag or purse that the patient carries. One
embodiment of the small receiver/display unit 46 has the appearance
of a pager, for example, so that the user is not identified as a
person using a medical device. Such receiver/display units may
optionally have one-way or two-way paging capabilities.
[0322] Large receiver/display units 48 may also be used. These
larger units 48 may be designed to sit on a shelf or nightstand.
The large receiver/display unit 48 may be used by parents to
monitor their children while they sleep or to awaken patients
during the night. In addition, the large receiver/display unit 48
may include a lamp, clock, or radio for convenience and/or for
activation as an alarm. One or both types of receiver/display units
46, 48 may be used.
[0323] The receiver/display units 46, 48, as illustrated in block
form at FIG. 22, typically include a receiver 150 to receive data
from the on-skin sensor control unit 44, an analyzer 152 to
evaluate the data, a display 154 to provide information to the
patient, and an alarm system 156 to warn the patient when a
condition arises. The receiver/display units 46, 48 may also
optionally include a data storage device 158, a transmitter 160,
and/or an input device 162. The receiver/display units 46,48 may
also include other components (not shown), such as a power supply
(e.g., a battery and/or a power supply that can receive power from
a wall outlet), a watchdog circuit, a bias current generator, and
an oscillator. These additional components are similar to those
described above for the on-skin sensor control unit 44.
[0324] In one embodiment, a receiver/display unit 48 is a bedside
unit for use by a patient at home. The bedside unit includes a
receiver and one or more optional items, including, for example, a
clock, a lamp, an auditory alarm, a telephone connection, and a
radio. The bedside unit also has a display, preferably, with large
numbers and/or letters that can be read across a room. The unit may
be operable by plugging into an outlet and may optionally have a
battery as backup. Typically, the bedside unit has a better antenna
than a small palm-size unit, so the bedside unit's reception range
is longer.
[0325] When an alarm is indicated, the bedside unit may activate,
for example, the auditory alarm, the radio, the lamp, and/or
initiate a telephone call. The alarm may be more intense than the
alarm of a small palm-size unit to, for example, awaken or
stimulate a patient who may be asleep, lethargic, or confused.
Moreover, a loud alarm may alert a parent monitoring a diabetic
child at night.
[0326] The bedside unit may have its own data analyzer and data
storage. The data may be communicated from the on-skin sensor unit
or another receiver/display unit, such as a palm-size or small
receiver/display unit. Thus, at least one unit has all the relevant
data so that the data can be downloaded and analyzed without
significant gaps.
[0327] Optionally, the beside unit has an interface or cradle into
which a small receiver/display unit may be placed. The bedside unit
may be capable of utilizing the data storage and analysis
capabilities of the small receiver/display unit and/or receive data
from the small receiver/display unit in this position. The bedside
unit may also be capable of recharging a battery of the small
receiver/display unit.
[0328] The receiver 150 typically is formed using known receiver
and antenna circuitry and is often tuned or tunable to the
frequency or frequency band of the transmitter 98 in the on-skin
sensor control unit 44. Typically, the receiver 150 is capable of
receiving signals from a distance greater than the transmitting
distance of the transmitter 98. The small receiver/display unit 46
can typically receive a signal from an on-skin sensor control unit
44 that is up to 2 meters, preferably up to 5 meters, and more
preferably up to 10 meters or more, away. A large receiver/display
unit 48, such as a bedside unit, can typically receive a receive a
signal from an on-skin sensor control unit 44 that is up to 5
meters distant, preferably up to 10 meters distant, and more
preferably up to 20 meters distant or more.
[0329] In one embodiment, a repeater unit (not shown) is used to
boost a signal from an on-skin sensor control unit 44 so that the
signal can be received by a receiver/display unit 46, 48 that may
be distant from the on-skin sensor control unit 44. The repeater
unit is typically independent of the on-skin sensor control unit
44, but, in some cases, the repeater unit may be configured to
attach to the on-skin sensor control unit 44. Typically, the
repeater unit includes a receiver for receiving the signals from
the on-skin sensor control unit 44 and a transmitter for
transmitting the received signals. Often the transmitter of the
repeater unit is more powerful than the transmitter of the on-skin
sensor control unit, although this is not necessary. The repeater
unit may be used, for example, in a child's bedroom for
transmitting a signal from an on-skin sensor control unit on the
child to a receiver/display unit in the parent's bedroom for
monitoring the child's analyte levels. Another exemplary use is in
a hospital with a display/receiver unit at a nurse's station for
monitoring on-skin sensor control unit(s) of patients.
[0330] The presence of other devices, including other on-skin
sensor control units, may create noise or interference within the
frequency band of the transmitter 98. This may result in the
generation of false data. To overcome this potential difficulty,
the transmitter 98 may also transmit a code to indicate, for
example, the beginning of a transmission and/or to identify,
preferably using a unique identification code, the particular
on-skin sensor control unit 44 in the event that there is more than
one on-skin sensor control unit 44 or other transmission source
within range of the receiver/display unit 46, 48. The provision of
an identification code with the data may reduce the likelihood that
the receiver/display unit 46, 48 intercepts and interprets signals
from other transmission sources, as well as preventing "crosstalk"
with different on-skin sensor control units 44. The identification
code may be provided as a factory-set code stored in the sensor
control unit 44. Alternatively, the identification code may be
randomly generated by an appropriate circuit in the sensor control
unit 44 or the receiver/display unit 46, 48 (and transmitted to the
sensor control unit 44) or the identification code may be selected
by the patient and communicated to the sensor control unit 44 via a
transmitter or an input device coupled to the sensor control unit
44.
[0331] Other methods may be used to eliminate "crosstalk" and to
identify signals from the appropriate on-skin sensor control unit
44. In some embodiments, the transmitter 98 may use encryption
techniques to encrypt the datastream from the transmitter 98. The
receiver/display unit 46, 48 contains the key to decipher the
encrypted data signal. The receiver/display unit 46, 48 then
determines when false signals or "crosstalk" signals are received
by evaluation of the signal after it has been deciphered. For
example, the analyzer 152 in the one or more receiver/display units
46, 48 compares the data, such as current measurements or analyte
levels, with expected measurements (e.g., an expected range of
measurements corresponding to physiologically relevant analyte
levels). Alternatively, an analyzer in the receiver/display units
46, 48 searches for an identification code in the decrypted data
signal.
[0332] Another method to eliminate "crosstalk", which is typically
used in conjunction with the identification code or encryption
scheme, includes providing an optional mechanism in the on-skin
sensor control unit 44 for changing transmission frequency or
frequency bands upon determination that there is "crosstalk". This
mechanism for changing the transmission frequency or frequency band
may be initiated by the receiver/display unit automatically, upon
detection of the possibility of cross-talk or interference, and/or
by a patient manually. For automatic initiation, the
receiver/display unit 46, 48 transmits a signal to the optional
receiver 99 on the on-skin sensor control unit 44 to direct the
transmitter 98 of the on-skin sensor control unit 44 to change
frequency or frequency band.
[0333] Manual initiation of the change in frequency or frequency
band may be accomplished using, for example, an actuator (not
shown) on the receiver/display unit 46, 48 and/or on the on-skin
sensor control unit 44 which a patient operates to direct the
transmitter 98 to change frequency or frequency band. The operation
of a manually initiated change in transmission frequency or
frequency band may include prompting the patient to initiate the
change in frequency or frequency band by an audio or visual signal
from the receiver/display unit 46, 48 and/or on-skin sensor control
unit 44.
[0334] Returning to the receiver 150, the data received by the
receiver 150 is then sent to an analyzer 152. The analyzer 152 may
have a variety of functions, similar to the processor circuit 109
of the on-skin sensor control unit 44, including 1) modifying the
signals from the sensor 42 using calibration data and/or
measurements from the temperature probe 66, 2) determining a level
of an analyte in the interstitial fluid, 3) determining a level of
an analyte in the bloodstream based on the sensor measurements in
the interstitial fluid, 4) determining if the level, rate of
change, and/or acceleration in the rate of change of the analyte
exceeds or meets one or more threshold values, 5) activating an
alarm system 156 and/or 94 if a threshold value is met or exceeded,
6) evaluating trends in the level of an analyte based on a series
of sensor signals. 7) determine a dose of a medication, and 7)
reduce noise or error contributions (e.g., through signal averaging
or comparing readings from multiple electrodes). The analyzer 152
may be simple and perform only one or a small number of these
functions or the analyzer 152 may perform all or most of these
functions.
[0335] The output from the analyzer 152 is typically provided to a
display 154. A variety of displays 154 may be used including
cathode ray tube displays (particularly for larger units), LED
displays, or LCD displays. The display 154 may be monochromatic
(e.g., black and white) or polychromatic (i.e., having a range of
colors). The display 154 may contain symbols or other indicators
that are activated under certain conditions (e.g., a particular
symbol may become visible on the display when a condition, such as
hyperglycemia, is indicated by signals from the sensor 42). The
display 154 may also contain more complex structures, such as LCD
or LED alphanumeric structures, portions of which can be activated
to produce a letter, number, or symbol. For example, the display
154 may include region 164 to display numerically the level of the
analyte, as illustrated in FIG. 23. In one embodiment, the display
154 also provides a message to the patient to direct the patient in
an action. Such messages may include, for example, "Eat Sugar", if
the patient is hypoglycemic, or "Take Insulin", if the patient is
hyperglycemic.
[0336] One example of a receiver/display unit 46, 48 is illustrated
in FIG. 23. The display 154 of this particular receiver/display
unit 46, 48 includes a portion 164 which displays the level of the
analyte, for example, the blood glucose concentration, as
determined by the processing circuit 109 and/or the analyzer 152
using signals from the sensor 42. The display also includes various
indicators 166 which may be activated under certain conditions. For
example, the indicator 168 of a glucose monitoring device may be
activated if the patient is hyperglycemic. Other indicators may be
activated in the cases of hypoglycemia (170), impending
hyperglycemia (172), impending hypoglycemia (174), a malfunction,
an error condition, or when a calibration sample is needed (176).
In some embodiments, color coded indicators may be used.
Alternatively, the portion 164 which displays the blood glucose
concentration may also include a composite indicator 180 (see FIG.
24), portions of which may be appropriately activated to indicate
any of the conditions described above.
[0337] The display 154 may also be capable of displaying a graph
178 of the analyte level over a period of time, as illustrated in
FIG. 24. Examples of other graphs that may be useful include graphs
of the rate of change or acceleration in the rate of change of the
analyte level over time. In some embodiments, the receiver/display
unit is configured so that the patient may choose the particular
display (e.g., blood glucose concentration or graph of
concentration versus time) that the patient wishes to view. The
patient may choose the desired display mode by pushing a button or
the like, for example, on an optional input device 162.
[0338] The receiver/display units 46, 48 also typically include an
alarm system 156. The options for configuration of the alarm system
156 are similar to those for the alarm system 104 of the on-skin
sensor control unit 44. For example, if glucose is the analyte,
than the on-skin sensor control unit 44 may include an alarm system
156 that warns the patient of conditions such as hypoglycemia,
hyperglycemia, impending hypoglycemia, and/or impending
hyperglycemia. The alarm system 156 is triggered when the data from
the analyzer 152 reaches or exceeds a threshold value. The
threshold values may correspond to interstitial fluid glucose
concentrations or sensor signals (e.g., current or converted
voltage values) which correlate to the above-mentioned blood
glucose levels.
[0339] The alarm system 156 may also, or alternatively, be
activated when the rate or acceleration of an increase or decrease
in analyte level reaches or exceeds a threshold value. For example,
in the case of a subcutaneous glucose monitor, the alarm system 156
might be activated if the rate of change in glucose concentration
exceeds a threshold value which might indicate that a hyperglycemic
or hypoglycemic condition is likely to occur.
[0340] The alarm system 156 may be configured to activate when a
single data point meets or exceeds a particular threshold value.
Alternatively, the alarm may be activated only when a predetermined
number of datapoints spanning a predetermined amount of time meet
or exceed the threshold value. As another alternative, the alarm
may be activated only when the datapoints spanning a predetermined
amount of time have an average value which meets or exceeds the
threshold value. Each condition that can trigger an alarm may have
a different alarm activation condition. In addition, the alarm
activation condition may change depending on current conditions
(e.g., an indication of impending hyperglycemia may alter the
number of datapoints or the amount of time that is tested to
determine hyperglycemia).
[0341] The alarm system 156 may contain one or more individual
alarms. Each of the alarms may be individually activated to
indicate one or more conditions of the analyte. The alarms may be,
for example, auditory or visual. Other sensory-stimulating alarm
systems by be used including alarm systems 156 that direct the
on-skin sensor control unit 44 to heat, cool, vibrate, or produce a
mild electrical shock. In some embodiments, the alarms are auditory
with a different tone, note, or volume indicating different
conditions. For example, a high note might indicate hyperglycemia
and a low note might indicate hypoglycemia. Visual alarms may also
use a difference in color or brightness to indicate different
conditions. In some embodiments, an auditory alarm system might be
configured so that the volume of the alarm increases over time
until the alarm is deactivated.
[0342] In some embodiments, the alarms may be automatically
deactivated after a predetermined time period. In other
embodiments, the alarms may be configured to deactivate when the
data no longer indicate that the condition which triggered the
alarm exists. In these embodiments, the alarms may be deactivated
when a single data point indicates that the condition no longer
exists or, alternatively, the alarm may be deactivated only after a
predetermined number of datapoints or an average of datapoints
obtained over a given period of time indicate that the condition no
longer exists.
[0343] In yet other embodiments, the alarm may be deactivated
manually by the patient or another person in addition to or as an
alternative to automatic deactivation. In these embodiments, a
switch is provided which when activated turns off the alarm. The
switch may be operatively engaged (or disengaged depending on the
configuration of the switch) by, for example, pushing a button on
the receiver/display unit 46, 48. One configuration of the alarm
system 156 has automatic deactivation after a period of time for
alarms that indicate an impending condition (e.g., impending
hypoglycemia or hyperglycemia) and manual deactivation of alarms
which indicate a current condition (e.g., hypoglycemia or
hyperglycemia).
[0344] The receiver/display units 46, 48 may also include a number
of optional items. One item is a data storage unit 158. The data
storage unit 158 may be desirable to store data for use if the
analyzer 152 is configured to determine trends in the analyte
level. The data storage unit 158 may also be useful to store data
that may be downloaded to another receiver/display unit, such as a
large display unit 48. Alternatively, the data may be downloaded to
a computer or other data storage device in a patient's home, at a
doctor's office, etc. for evaluation of trends in analyte levels. A
port (not shown) may be provided on the receiver/display unit 46,
48 through which the stored data may be transferred or the data may
be transferred using an optional transmitter 160. The data storage
unit 158 may also be activated to store data when a directed by the
patient via, for example, the optional input device 162. The data
storage unit 158 may also be configured to store data upon
occurrence of a particular event, such as a hyperglycemic or
hypoglycemic episode, exercise, eating, etc. The storage unit 158
may also store event markers with the data of the particular event.
These event markers may be generated either automatically by the
display/receiver unit 46, 48 or through input by the patient.
[0345] The receiver/display unit 46, 48 may also include an
optional transmitter 160 which can be used to transmit 1)
calibration information, 2) a signal to direct the transmitter 98
of the on-skin sensor control unit 44 to change transmission
frequency or frequency bands, and/or 3) a signal to activate an
alarm system 104 on the on-skin sensor control unit 44, all of
which are described above. The transmitter 160 typically operates
in a different frequency band than the transmitter 98 of the
on-skin sensor control unit 44 to avoid cross-talk between the
transmitters 98, 160. Methods may be used to reduce cross-talk and
the reception of false signals, as described above in connection
with the transmitter 100 of the on-skin sensor control unit 44. In
some embodiments, the transmitter 160 is only used to transmit
signals to the sensor control unit 44 and has a range of less than
one foot, and preferably less than six inches. This then requires
the patient or another person to hold the receiver/display unit 46
near the sensor control unit 44 during transmission of data, for
example, during the transmission of calibration information.
Transmissions may also be performed using methods other than rf
transmission, including optical or wire transmission.
[0346] In addition, in some embodiments of the invention, the
transmitter 160 may be configured to transmit data to another
receiver/display unit 46, 48 or some other receiver. For example, a
small receiver/display unit 46 may transmit data to a large
receiver/display unit 48, as illustrated in FIG. 1. As another
example, a receiver/display unit 46, 48 may transmit data to a
computer in the patient's home or at a doctor's office. Moreover,
the transmitter 160 or a separate transmitter may direct a
transmission to another unit or to a telephone or other
communications device that alerts a doctor or other individual when
an alarm is activated and/or if, after a predetermined time period,
an activated alarm has not been deactivated, suggesting that the
patient may require assistance. In some embodiments, the
receiver/display unit is capable of one-way or two-way paging
and/or is coupled to a telephone line to send and/or receive
messages from another, such as a health professional monitoring the
patient.
[0347] Another optional component for the receiver/display unit 46,
48 is an input device 162, such as a keypad or keyboard. The input
device 162 may allow numeric or alphanumeric input. The input
device 162 may also include buttons, keys, or the like which
initiate functions of and/or provide input to the analyte
monitoring device 40. Such functions may include initiating a data
transfer, manually changing the transmission frequency or frequency
band of the transmitter 98, deactivating an alarm system 104, 156,
inputting calibration data, and/or indicating events to activate
storage of data representative of the event.
[0348] Another embodiment of the input device 162 is a touch screen
display. The touch screen display may be incorporated into the
display 154 or may be a separate display. The touch screen display
is activated when the patient touches the screen at a position
indicated by a "soft button" which corresponds to a desired
function. Touch screen displays are well known.
[0349] In addition, the analyte monitoring device 40 may include
password protection to prevent the unauthorized transmission of
data to a terminal or the unauthorized changing of settings for the
device 40. A patient may be prompted by the display 154 to input
the password using the input device 152 whenever a
password-protected function is initiated.
[0350] Another function that may be activated by the input device
162 is a deactivation mode. The deactivation mode may indicate that
the receiver/display unit 46, 48 should no longer display a portion
or all of the data. In some embodiments, activation of the
deactivation mode may even deactivate the alarm systems 104, 156.
Preferably, the patient is prompted to confirm this particular
action. During the deactivation mode, the processing circuit 109
and/or analyzer 152 may stop processing data or they may continue
to process data and not report it for display and may optionally
store the data for later retrieval.
[0351] Alternatively, a sleep mode may be entered if the input
device 162 has not been activated for a predetermined period of
time. This period of time may be adjustable by the patient or
another individual. In this sleep mode, the processing circuit 109
and/or analyzer 152 typically continue to obtain measurements and
process data, however, the display is not activated. The sleep mode
may be deactivated by actions, such as activating the input device
162. The current analyte reading or other desired information may
then be displayed.
[0352] In one embodiment, a receiver/display unit 46 initiates an
audible or visual alarm when the unit 46 has not received a
transmission from the on-skin sensor control unit within a
predetermined amount of time. The alarm typically continues until
the patient responds and/or a transmission is received. This can,
for example, remind a patient if the receiver/display unit 46 is
inadvertently left behind.
[0353] In another embodiment, the receiver/display unit 46, 48 is
integrated with a calibration unit (not shown). For example, the
receiver/display unit 46, 48 may, for example, include a
conventional blood glucose monitor. Another useful calibration
device utilizing electrochemical detection of analyte concentration
is described in U.S. patent application Ser. No. 08/795,767,
incorporated herein by reference. Other devices may be used
including those that operate using, for example, electrochemical
and colorimetric blood glucose assays, assays of interstitial or
dermal fluid, and/or non-invasive optical assays. When a
calibration of the implanted sensor is needed, the patient uses the
integrated in vitro monitor to generate a reading. The reading may
then, for example, automatically be sent by the transmitter 160 of
the receiver/display unit 46, 48 to calibrate the sensor 42.
[0354] Integration with a Drug Administration System
[0355] FIG. 25 illustrates a block diagram of a sensor-based drug
delivery system 250 according to the present invention. The system
may provide a drug to counteract the high or low level of the
analyte in response to the signals from one or more sensors 252.
Alternatively, the system monitors the drug concentration to ensure
that the drug remains within a desired therapeutic range. The drug
delivery system includes one or more (and preferably two or more)
subcutaneously implanted sensors 252, an on-skin sensor control
unit 254, a receiver/display unit 256, a data storage and
controller module 258, and a drug administration system 260. In
some cases, the receiver/display unit 256, data storage and
controller module 258, and drug administration system 260 may be
integrated in a single unit. The sensor-based drug delivery system
250 uses data form the one or more sensors 252 to provide necessary
input for a control algorithm/mechanism in the data storage and
controller module 252 to adjust the administration of drugs. As an
example, a glucose sensor could be used to control and adjust the
administration of insulin.
[0356] In FIG. 25, sensor 252 produces signals correlated to the
level of the drug or analyte in the patient. The level of the
analyte will depend on the amount of drug delivered by the drug
administration system. A processor 262 in the on-skin sensor
control unit 254, as illustrated in FIG. 25, or in the
receiver/display unit 256 determines the level of the analyte, and
possibly other information, such as the rate or acceleration of the
rate in the increase or decrease in analyte level. This information
is then transmitted to the data storage and controller module 252
using a transmitter 264 in the on-skin sensor control unit 254, as
illustrated in FIG. 25, or a non-integrated receiver/display unit
256.
[0357] If the drug delivery system 250 has two or more sensors 252,
the data storage and controller module 258 may verify that the data
from the two or more sensors 252 agrees within predetermined
parameters before accepting the data as valid. This data may then
be processed by the data storage and controller module 258,
optionally with previously obtained data, to determine a drug
administration protocol. The drug administration protocol is then
executed using the drug administration system 260, which may be an
internal or external infusion pump, syringe injector, transdermal
delivery system (e.g., a patch containing the drug placed on the
skin), or inhalation system. Alternatively, the drug storage and
controller module 258 may provide a the drug administration
protocol so that the patient or another person may provide the drug
to the patient according to the profile.
[0358] In one embodiment of the invention, the data storage and
controller module 258 is trainable. For example, the data storage
and controller module 258 may store glucose readings over a
predetermined period of time, e.g., several weeks. When an episode
of hypoglycemia or hyperglycemia is encountered, the relevant
history leading to such event may be analyzed to determine any
patterns which might improve the system's ability to predict future
episodes. Subsequent data might be compared to the known patterns
to predict hypoglycemia or hyperglycemia and deliver the drug
accordingly. In another embodiment, the analysis of trends is
performed by an external system or by the processing circuit 109 in
the on-skin sensor control unit 254 or the analyzer 152 in the
receiver/display unit 256 and the trends are incorporated in the
data storage and controller 258.
[0359] In one embodiment, the data storage and controller module
258, processing circuit 109, and/or analyzer 152 utilizes
patient-specific data from multiple episodes to predict a patient's
response to future episodes. The multiple episodes used in the
prediction are typically responses to a same or similar external or
internal stimulus. Examples of stimuli include periods of
hypoglycemia or hyperglycemia (or corresponding conditions for
analytes other than glucose), treatment of a condition, drug
delivery (e.g., insulin for glucose), food intake, exercise,
fasting, change in body temperature, elevated or lowered body
temperature (e.g., fever), and diseases, viruses, infections, and
the like. By analyzing multiple episodes, the data storage and
controller module 258, processing circuit 109, and/or analyzer 152
can predict the coarse of a future episode and provide, for
example, a drug administration protocol or administer a drug based
on this analysis. An input device (not shown) may be used by the
patient or another person to indicate when a particular episode is
occurring so that, for example, the data storage and controller
module 258, processing circuit 109, and/or analyzer 152 can tag the
data as resulting from a particular episode, for use in further
analyses.
[0360] In addition, the drug delivery system 250 may be capable of
providing on-going drug sensitivity feedback. For example, the data
from the sensor 252 obtained during the administration of the drug
by the drug administration system 260 may provide data about the
individual patient's response to the drug which can then be used to
modify the current drug administration protocol accordingly, both
immediately and in the future. An example of desirable data that
can be extracted for each patient includes the patient's
characteristic time constant for response to drug administration
(e.g., how rapidly the glucose concentration falls when a known
bolus of insulin is administered). Another example is the patient's
response to administration of various amounts of a drug (e.g., a
patient's drug sensitivity curve). The same information may be
stored by the drug storage and controller module and then used to
determine trends in the patient's drug response, which may be used
in developing subsequent drug administration protocols, thereby
personalizing the drug administration process for the needs of the
patient.
[0361] Relationship of Subcutaneous and Blood Analyte Levels
[0362] It is often useful to determine analyte concentration in one
fluid (e.g., blood) even though the measurements of analyte
concentration are performed on another fluid (e.g., subcutaneous
fluid). For example, it may be important to know blood glucose
concentration for accurate diagnosis and/or insulin injections, or
for comparison with other techniques, but it is more convenient
and/or less painful or intrusive to measure subcutaneous glucose
concentrations. Sensor measurements made using subcutaneous fluid
may be different from the desired quantity (e.g., blood glucose
concentration) because of the existence of a mass transfer barrier,
source and/or sink between compartment A, the region of measurement
(e.g., the subcutaneous tissue), and compartment B, the region of
interest (e.g., the blood). For any such problem, one needs to
develop a model that relates q.sub.A, the measured quantity, to
q.sub.B, the desired quantity, using a system of equations:
q.sub.A=q.sub.B. To solve for the desired quantity, the operator
must be inverted. If the operator happens to be noninvertable or
unstable to inversion, the use of such a model may be hindered.
[0363] One solution to this dilemma is the application of
regularization techniques that, when used in conjunction with a
model, can predict the desired quantity from the measured quantity.
These methods often permit the imposition of a smoothing
requirement that changes the operator , making it invertible. These
regularization techniques can be used to infer one function from
another measured function using a postulated relationship between
them.
[0364] With respect to subcutaneously implanted glucose sensors,
the concentration of glucose in the blood is desired, especially
for accurate dosing of insulin. There is typically a time lag
between changes in glucose concentration of the blood and the
subcutaneous tissue after, for example, an injection of insulin. To
predict this time lag and correlate the two concentrations, the
glucose transport processes that mediate the transport of glucose
from the blood to the subcutaneous tissue are investigated.
[0365] Three types of glucose transport process exist: active
transport, facilitated transport and passive transport. Active
transport processes are present in the lumen of the small intestine
and in the renal tubules, where glucose is transported against its
concentration gradient, requiring energy. Facilitated transport
processes include those in which, for example, carrier proteins,
known as a glucose transporters, or GluTs, are present at a
membrane surface to aid the diffusion of glucose across the
membrane, as in adipocytes and in the blood-brain barrier. Finally,
passive transport includes simple or Fickian diffusion which is
typically driven by a concentration gradient and needs no special
carrier proteins or energy.
[0366] Transfer of Glucose from Blood to Interstitial Fluid
[0367] A subcutaneously implanted sensor is placed in the
interstitial fluid of the subcutaneous tissue. Typically, the
important transport process are facilitated diffusion and a mass
transfer resistance to transport of glucose between the blood and
subcutaneous tissue. Thus, the relationship between the
concentrations of glucose in the blood and subcutaneous tissues can
be modeled by the mass transfer resistance from the blood to the
subcutaneous region near the sensor and by the uptake of glucose by
the surrounding subcutaneous tissue. Following a material balance,
the rate of accumulation of glucose in the sensing volume V is
given by the net rate of mass transfer of glucose into the region
less the uptake of glucose by the surrounding cells via facilitated
diffusion, which can be modeled using a reaction term. This
relationship between the concentration of glucose in the
subcutaneous tissue S and that in the blood B is given by, 1 V S t
= k m A ( B - S ) - Vk r S K m + S ( 1 )
[0368] where A is the surface area of the region surrounding the
sensor, k.sub.m is a mass transfer coefficient, K.sub.m is a
Michaelis-Menten constant, and k.sub.r' is the reaction rate
constant for uptake of glucose by the subcutaneous tissue. The mass
transfer coefficient, Michaelis-Menten constant, and reaction rate
constant for uptake of glucose by the subcutaneous tissue may be
determined experimentally for a particular animal, species, or as a
generally applicable value. Alternatively, these values may be
estimated.
[0369] The reaction rate constant may depend on the local insulin
concentration 1, as modeled, for example, by Yeh et al., Biochem.
34:523-531 (1995), incorporated by reference. However, for purposes
of this discussion, the reaction rate constant is assumed to be
constant. Appropriate changes in the equations below can be made if
the reaction rate constant is dependent on the local insulin
concentration.
[0370] Dividing equation (1) by the volume of the sensor region V
yields: 2 S t = ^ ( B - S ) - k r S K m + S ( 2 )
[0371] where {circumflex over (.beta.)}=k.sub.mA/V and corresponds
to the reciprocal of the time constant for mass transfer. It is
convenient to non-dimensionalize the equation as follows, defining
3 B _ = B B 0 , S _ = S B 0 , k r = k r
[0372] /(.beta.{circumflex over (B)}.sub.0) and {overscore
(t)}={circumflex over (.beta.)}t, where B.sub.O can be an
arbitrarily defined B.sub.0B.sub.0 blood glucose concentration
(e.g., a starting blood glucose concentration). Equation (2) then
becomes 4 S _ t _ = B _ - [ 1 + k r K m B 0 + S _ ] S _ ( 3 )
[0373] The contents in brackets can be referred to as the
pseudoconstant .beta.. When there is no reaction, .beta. is equal
to 1; when there is a reaction, .beta. is a weak function of
{overscore (S)}, but {overscore (S)} typically does not change
much. So, it can be assumed that .beta. is constant over the time
scale of the computation, letting {overscore (S)} be equal to the
value at the center of the computation window. (The computation
window is described following equation (7).) The nondimensional
variables {overscore (S)}, {overscore (B)} and {overscore (t)} will
continue to be referred to, but the overbars are removed for the
remainder of this specification. The final equation, 5 S t = B - S
, ( 4 )
[0374] determines the subcutaneous glucose concentration given the
blood glucose concentration and can be termed a forward model.
[0375] Inversion of the Forward Model
[0376] The forward model is inverted to infer the blood glucose
concentration given measured subcutaneous glucose concentrations.
Predictions made from inversions may be highly sensitive to
measurement errors and the inherent imperfections present in any
mathematical model. Thus, regularization is often useful. If no
regularization is used, the solution may be unstable and/or
unreasonable.
[0377] As part of the regularization techniques, a smoothness
condition may be imposed to minimize a function. The smoothness
condition can include a combination of model fit and required
smoothness. The minimization may result in a slightly modified set
of equations which are well-conditioned (e.g., invertible and
stable) and readily inverted and solved. Thus, rather than strictly
forcing the data to fit the model, the data is forced to be smooth
(as defined by the regularization technique) and fit the model
reasonably well. A set of equations is then derived that use the
measured value of the subcutaneous concentration of glucose to
predict the concentration of glucose in the blood.
[0378] To invert the forward model, it may be useful to rewrite the
forward model in the form of a Volterra integral equation. To do
so, both sides of equation (4) are multiplied by the function
.phi.(t)=e.sup..beta.t (5a)
[0379] to yield 6 t [ ( t ) S ] = ( t ) B . (5b)
[0380] Recall that .beta. is a pseudo-constant that is actually a
mild function of S when reaction is present. Taking the definite
integral of equation (5b) between times .theta. and t, and dividing
both sides by .phi.(t) gives, 7 S ( t ) = S ( ) - ( t - ) + t B ( )
- ( i - ) . ( 6 )
[0381] In the above equation, the variable .theta. is the initial
time and t is the final time for the present window of computation.
Integration can be done numerically using a finite difference
scheme. B(t) is computed for N times using a set of equations, or a
single matrix-vector equation, with a time t represented on each of
the N rows. As an approximation, 8 t B ( ) - ( i - ) = i = 1 N B (
t i ) - ( i - t i ) t .times. W i ( 7 )
[0382] where N is the number of discretization points in the
computation window and W is some weighting factor defined by a
choice of quadrature scheme. For example, the integral in equation
(7) is approximated by choosing weights that apply an extended
Simpson's rule.
[0383] The time window contains N times at which subcutaneous
measurements are taken, and which are separated by an interval At,
so the size of the window is (N-1).DELTA.t. The blood glucose
concentration must by computed numerically, so B is discretized,
i.e., represented by a piecewise constant over the window of
computation. An advantage of computing the blood glucose value at
the ending time of a window is that the method can be implemented
continuously, updating the blood glucose concentration as more
subcutaneous data become available. This allows the determination
of blood glucose concentration from earlier measured analyte
concentrations. This is in contrast to conventional analysis
techniques that require measurements before and after the point in
time at which the blood glucose concentration is determined.
[0384] In some instances this differential treatment may be
sufficient. However, the solution to equation (6) may be sensitive
to imperfections in the data and in the model, and its application
alone may result in oscillatory predictions of the concentration of
glucose in the blood, as shown in FIG. 30.
[0385] Regularization techniques can be used to form a better
behaved solution.
[0386] The solution of the integral equation for B(.tau.) (or
simply the vector b, which is the vector of blood glucose values,
b.sub.i at the points in the present computation window) can be
conditioned to be smooth in addition to closely satisfying equation
(6) with experimental measurements of S(t). For example, the
functional f[b].
f[b]=.chi..sup.2[b]+.lambda..PSI.[b], (8)
[0387] can be minimized over any window of data points, where 9 2 [
b ] = = i i + N ( t B ( ) - i ( t - ) - S ( t ) + S ( ) - i ( t - )
) 2 . ( 9 )
[0388] The functional .chi..sup.2 represents the fit between the
prediction of the model and the experimental data and the
functional .psi. indicates the smoothness of the prediction. The
.lambda. variable is a weight which balances the amount of
smoothing to data-matching and can be constrained to range from 0
to .infin.. The-functional .psi. may be chosen based on an a priori
belief about the quality of the output. If the output is likely to
be constant over one window of computation, a first-order
regularization, in which first derivatives are minimized over the
window of interest, can be chosen, resulting in: 10 [ b ] = t [ B '
( ) ] 2 = i i + N - 1 [ B ( t + 1 ) - B ( t ) ] 2 . ( 10 )
[0389] The last term in equation (10) is a finite difference
estimation of the integral where .DELTA..tau. is the time
difference between data points. If, instead, the solution is
thought to be linear over one window of computation, and a
second-order regularization can be imposed that will minimize
second derivatives, resulting in: 11 [ b ] = t [ B " ( ) ] 2 = i i
+ N - 2 [ - B ( t + 2 ) + 2 B ( t + 1 ) - B ( t ) ] 2 ( 11 )
[0390] Equation (8) above may be written in the following matrix
form:
f=(A.multidot.b-c).sup.2+.lambda.(b.multidot.H.multidot.b) (12)
[0391] where 12 A j = t u - i ( t - ) = { 1 2 - i ( t j - t ) t = i
or j - i ( t i - t ) t i < < j , 0 j < ( 13 )
[0392] and
c.sub..mu.=S(t.sub..mu.)-S(.theta.)e.sup..beta.,(t.sub..sup..mu..sup.-.the-
ta.). (14)
[0393] The definition for H stems from the choice of regularization
such that first or the second derivatives over the window of
computation are minimized. If a first-order regularizations is
chosen, the matrix H is given by, 13 H = [ 1 - 1 0 0 0 0 - 1 2 - 1
0 0 0 0 - 1 2 - 1 0 0 0 0 - 1 2 - 1 0 0 0 0 - 1 2 - 1 0 0 0 0 - 1 1
] ( 15 )
[0394] On the other hand, choosing a second-order regularization
gives 14 H = [ 1 - 2 1 0 0 0 0 0 - 2 5 - 4 1 0 0 0 0 1 - 4 6 - 4 1
0 0 0 0 1 - 4 6 - 4 1 0 0 0 0 1 - 4 6 - 4 1 0 0 0 0 1 - 4 6 - 4 1 0
0 0 0 1 - 4 5 - 2 0 0 0 0 0 1 - 2 1 ] ( 16 )
[0395] The matrix H that minimizes b, or implements a zeroeth-order
regularization, is given simply by the identity matrix.
[0396] Equation (12) is minimized by setting df/dB equal to 0, and,
after some algebra, the blood glucose concentrations over the
window of computation are given by.
(A.sup.T.multidot.A+.lambda.H).multidot.b=A.sup.Tc. (17)
[0397] The solution to the model was found at each time for which
measurements were acquired in the experiment. We solve for b in
equation (17) at each window of computation using known LU
decomposition and back-substitution.
[0398] In another embodiment, the formulation may include a fixed
initial condition. The functional to be minimized can be
differentiated as before and the problem solved using identical
methods. By enforcing the initial condition, the solution becomes a
bit more unstable, because the initial condition that is being
forced may not give the best fit. Other than causing more
instability, this method changes the prediction very little.
[0399] Besides assuring a relatively smooth solution for B(.tau.),
the regularization techniques may be more desirable than the
differential method for another reason. To use the differential
method, the sensor data is often smoothed before processing, which
could produce a lag in the results because backward smoothing would
be applied, since the application of real-time inversion dictates
that the future data would be unknown. By using the regularization
techniques, a relatively smooth solution can be obtained without
creating this lag.
[0400] Processor
[0401] The determination of the blood glucose concentration from
subcutaneous glucose concentration measurements can be performed by
a processor (e.g., processing circuit 109 of FIG. 18A or, 18B or
analyzer 152 of FIG. 22), with or without a storage medium, in
which the determination procedure is performed by software,
hardware, or a combination thereof. According to another
embodiment, this same determination is accomplished using discrete
or semi-programmable hardware configured, for example, using a
hardware descriptive language, such as Verilog. In yet another
embodiment, the determination may be performed using a processor
having at least one look-up table arrangement with data stored
therein to represent the complete result or partial results of the
above equations based on a given set of input data, the input data
corresponding to parameters used on the right side of the
equations.
EXAMPLES
Example 1
Oral Glucose Tolerance Test Function
[0402] In order to test the performance of the above inverse model
under realistic conditions, a test function was used that resembles
the typical human response to a substantial change in glucose
intake or utilization. A simulation of a response to an oral
glucose tolerance test (OGTT) and a simple fit of the human OGTT
data presented by Jansson et al., Am. J. Physiol., 225:E218-220
(1988), incorporated herein by reference, results in the following
non-dimensional function: 15 B ( t ) = C B 0 ( - t - - t ) + 1 , (
18 )
[0403] where .gamma.=k.sub.1/{circumflex over (.beta.)},
.zeta.=k.sub.2/{circumflex over (.beta.)}, k.sub.1=0.054
min.sup.-1, k.sub.2=0.021 min.sup.-1,
C=85.5*[k.sub.1/(k.sub.1-k.sub.2)] mg/dl, and B.sub.0=95 mg/dl.
From previous comparisons with experimental data provided in
Schmidtke et al., Proc. of the Nat'l Acad of Science, 95, 294-299
(1998), incorporated herein by reference, {circumflex over
(.beta.)} was chosen to be 0.05 min.sup.-1. If no reaction is
present, then the forward problem can be solved analytically, and
the subcutaneous glucose concentration is given by: 16 S ( t ) = C
B 0 [ - t 1 - - - t 1 - ] + 1. ( 19 )
[0404] Three cases of varying magnitudes of the reaction term were
studied, including a) k.sub.r=0, b) k.sub.r=1 and K.sub.m=B.sub.o,
and c) k.sub.r=1 and K.sub.m=B.sub.o/3. The functions S(t) for the
above three cases and their corresponding function B(t) are plotted
in FIG. 31. The shapes of the input functions S(t) are shifted down
as the effective reaction rate constant, 17 k r K m B 0 + S ,
[0405] increases.
[0406] By varying the amount of noise on the input function, the
performance of the inverse model for a wide range of .lambda.
(10.sup.-9 to 10.sup.5) was analyzed for the three hypothetical
cases. For the tests performed here, an error magnification factor,
.epsilon., as a function of .lambda., and magnitude of reaction was
computed. The error magnification factor was defined as: 18 = %
Output RMS Error % Input RMS Error where ( 20 ) % Output RMS Error
= i ( predictedB i - trueB i ) 2 i trueB i 2 .times. 100 % , and (
21 ) % Input RMS Error = i ( noisyS i - trueS i ) 2 i trueS i 2
.times. 100 % . ( 22 )
[0407] trueS are the input values that are free of generated noise
and trueB are the values that would result from the equations if
the trueS values were used as input.
[0408] The input function was modified by the addition of white
noise or time-correlated noise. White noise was produced by first
finding the average value of the input function over the test
period. Then a random gaussian distribution was generated about
that average with standard deviations of 0.5, 1, and 2% of that
average:
S.sub.w(t)=S(t)+Gauss(mean,SD) (23)
[0409] In the above equation, S.sub.w(t) is the new subcutaneous
input function with white noise superimposed, where Gauss is a
function of the mean, 19 mean = 1 n i = 1 n S ( t i ) ( 24 )
[0410] and the standard deviation, 20 SD = p n i = 1 n S ( t i ) (
25 )
[0411] Gauss is the function that generates the random Gaussian
distribution with the given average and standard deviation. When p
was set equal to 1%, an input function with 1% white noise
resulted. This distribution was added to the OGTT function in
equation (18) to produce the white noise input function.
[0412] Time-correlated noise was constructed via a simple moving
average method, where the white noise input function produced above
is averaged over a window of time that is of size m so that 21 S c
( t i + m ) = 1 m j = 0 , 1 , 2 m S w ( t i + j ) . ( 26 )
[0413] where S.sub.c(t) is the input function with superimposed
time-correlated noise. In other words, the input function with
white noise was averaged over the ith window of time to give the
new function's value at the ending time of the window. The
nondimensional time window over which the values were averaged was
{fraction (1/100)}th of the total time of the test. The input
functions with these two types of noise are illustrated in FIG. 32
with RMS errors of 1%.
Example 2
Order of Regularization and Estimation of N and .DELTA.t
[0414] A comparison of first- and second-order regularization
methods for the case of no reaction (case a) and 1% white noise on
the input is shown in FIG. 33. The data in FIG. 33 were obtained
using a window size of 10 data points (N=10), which corresponds to
{fraction (1/50)}th of the total test time, and the .DELTA.t was
0.0185. The error magnification factor versus weighting factor
curves for the above cases and for the zeroeth-order regularization
are in FIG. 34.
[0415] For the three levels of white noise superimposed on the
synthetic subcutaneous glucose measurements, and for this window
size, the first-order regularization method predicted the blood
glucose concentration better than either the zeroeth- or
second-order regularization. The lowest .epsilon. for the zeroeth-,
first-, and second-order regularizations were 63, 7, and 13
respectively. Typically, first-order regularization is the
preferred method for most problems, unless one expects a constant
profile, in which zeroeth-order regularization would be the natural
choice.
[0416] FIG. 35 shows the error magnification-factor versus the
regularization parameter for a variety of sizes of computation
windows, and also for several sampling rates, where the sampling
rate is defined as .DELTA.t.sup.-1. The computation window size and
sample rate had a strong effect on the lowest .epsilon. achievable
using the inverse method. A decrease in the sampling rate, as
expected, causes the error magnification factor to increase, so
samples should be taken as often as possible. However, increasing
the sampling rate causes the condition number of the matrix.
A.sup.T.multidot.A+.lambda.H, to increase, increasing the
error.
[0417] As the sampling time between the measurements is increased,
the error magnification factor increases. Similarly, as the window
size grows, the error factor decreases. Window size is equal to
(N-1).DELTA.t, where N is the number of data points in the window,
and .DELTA.t is the time between the data points. Keeping the
sampling rate constant, FIG. 35 shows how .epsilon. changes as the
window size increases from N=10 to N=160. When N is increased from
10 to 20, the decrease in .epsilon. is larger than when N is
increased from 40 to 80. The reason for this is that, as the size
of the window gets larger, the trailing values of measurements will
have less of an effect on the solution, since the kernal is
exponential in time (see equation 14). Finally, when N increases
from 80 to 160, there is no substantial decrease in .epsilon.. A
reason for this could be that the window size has grown so much
that the first order derivatives can no longer be minimized and
expect a good a priori estimate of the behavior of the solution.
That is, the window size is now on order of the time constant of
the mass transfer coefficient. Also, as expected, there is never an
.epsilon. below 1, because the output function will always have at
least as much error as the input function. Note that for the
improvement on .epsilon. by increasing N from 10 to 80, the
computation time expense also increases. Keeping the sampling rate
constant, an optimal window size was found for this particular
problem which was 8 times larger that the one used above.
Application of this size of computation window yielded an error
magnification factor of 1.6, and required a regularization factor
.lambda. of 3.
[0418] With the larger, optimal window size, the regularization
methods were reexamined. Both first- and second-order
regularizations give good inverses, with an error magnification
factor (.epsilon.) of about 1.5 for each case, as shown in FIG. 36.
Note that a much larger regularization parameter is required for
the second-order than for the first-order method. The
regularization parameter typically indicates the relative amount of
model matching to smoothing imposed. Either of these methods could
be used for the remainder of the analysis.
Example 3
White Noise vs. Correlated Noise
[0419] In general, the method of regularization and inversion gave
similar results for both white noise and correlated noise on the
subcutaneous data. The data with correlated noise and the data with
white noise superimposed required similar weighting factors to give
similar values of .epsilon.. The correlated noise was smoother than
the white noise, but with larger error magnification factors than
for white noise. In fact, the correlated noise causes the model to
deviate from the true function for longer sustained times, so the
larger error magnification factors are not unexpected.
[0420] FIG. 37 illustrates the .epsilon.(.lambda.) curves for white
and correlated noise in the case of 1% RMS Input Error and reaction
case a. Both functions find their minimum at similar values of
.lambda., at about 3.0 and 6.0 for white noise and correlated
noise, respectively. Overall, the curves go to infinity as .lambda.
approaches zero, which indicates that a regularization is necessary
due to the instability of the inverse problem. Also, the curves
plateau as .lambda. increases beyond order 1, which shows us that
the regularization only causes more and more damping of the
solution as .lambda. increases, causing there to be a maximum
finite difference between the prediction and the true solution
(i.e., the solution goes to a constant about the initial point of
the prediction).
[0421] The error magnification factor, .epsilon., decreased as the
input error increased for both sets of data. Tables 1 and 2 contain
the results.
1TABLE 1 White noise, first-order regularization results. Input RMS
Output RMS Reaction Case Error Error Best .lambda. .epsilon.
k.sub.r = 0 0.5 1.16 0.9 2.32 k.sub.r = 0 1.0 1.58 3.0 1.58 k.sub.r
= 0 2.0 2.17 6.0 1.09 k.sub.r = 1, K.sub.m = B.sub.0/3 0.5 1.31 0.7
2.62 k.sub.r = 1, K.sub.m = B.sub.0/3 1.0 1.65 2.0 1.65 k.sub.r =
1, K.sub.m = B.sub.0/3 2.0 2.15 5.0 1.07 k.sub.r = 1, K.sub.m =
B.sub.0 0.5 1.04 0.8 2.08 k.sub.r = 1, K.sub.m = B.sub.0 1.0 1.31
2.0 1.31 k.sub.r = 1, K.sub.m = B.sub.0 2.0 1.75 4.0 0.88
[0422]
2TABLE 2 Time-correlated noise using simple moving average method,
first-order regularization results. Input RMS Output RMS Reaction
Case Error Error Best .lambda. .epsilon. k.sub.r = 0 0.32 1.20 0.6
3.75 k.sub.r = 0 0.53 1.63 2.0 3.08 k.sub.r = 0 1.00 2.21 6.0 2.21
k.sub.r = 1, K.sub.m = B.sub.0/3 0.32 1.24 0.5 3.88 k.sub.r = 1,
K.sub.m = B.sub.0/3 0.53 1.61 1.0 3.04 k.sub.r = 1, K.sub.m =
B.sub.0/3 1.00 2.13 4.0 2.13 k.sub.r = 1, K.sub.m = B.sub.0 0.32
0.84 0.6 2.63 k.sub.r = 1, K.sub.m = B.sub.0 0.53 1.18 1.0 2.23
k.sub.r = 1, K.sub.m = B.sub.0 1.00 1.66 3.0 1.66
Example 4
Estimation of Weighting Factor
[0423] Many workers have proposed methods for estimating the best
value for the weighting or regularization factor, .lambda.,
including Beck et al., Inverse Heat Conduction, John Wiley &
Sons, New York (1985); Graham, Bell Systems Tech. J, 62:101-110
(1983); Press et al., Numerical Recipes in Fortran-2nd Ed.,
Cambridge University Press (1992); and Reinsch, Numerische
Mathematik, 10:177-183 (1967), all of which are incorporated herein
by reference. As recommended in Press et al., the weight factor
.lambda. may be roughly estimated, at first, by implementing the
equation,
.lambda.=Tr(R.sup.T.multidot.R)/Tr(H) (27a)
[0424] where
R=A/.sigma.. (27b)
[0425] and .sigma. is the standard deviation of the measurements.
This estimate of .lambda. allow for approximately equal amounts of
model matching and smoothness. Another interpretation of the
conditions is that the data are required to fit the model only
within the measurement error.
[0426] The regularization parameter may also depend on the number
of measurements available, n, in addition to the standard deviation
of those measurements. Thus, the following condition on the
residual sum of the squares, , can be applied to find an
appropriate .lambda.:
=(A.multidot.b-c).sup.T(A.multidot.b-c) (28)
[0427] and require
[n-(2n).sup.1/2].sigma..sup.2<[n+(2n).sup.1/2].sigma..sup.2.
(29)
[0428] Criterion 1 can be defined as =n.sigma..sup.2. This method
can be referred to as the discrepancy criterion.
[0429] Alternatively, .lambda. can be selected using the concept of
the minimum squared error, as described in Beck et al., Inverse
Heat Conduction, John Wiley & Sons (1985), incorporated herein
by reference. This can be called criterion 2. The parameter
.lambda. will often have an optimal value that remains
approximately constant when the integration time interval and
sampling rate is constant, so this process of determining .lambda.
may only be necessary once for a given set of parameters.
[0430] Comparison of the two results, values are obtained within
the designated range, as shown in FIG. 38. The application of
criterion 1 gives .epsilon. equal to 2.5 (.lambda.=0.5) which is
very close to the criterion 2 result of .epsilon.=2.32
(.lambda.=1.0). The largest output error that would occur by
choosing .lambda. such that is within the above bounds is 4.5%
(.epsilon.=9.0) for the case of no reaction and 0.5% RMS error in
the input in the form of white noise. Choosing equal to
n.sigma..sup.2 gives .epsilon. equal to 2.5. Thus, the criterion 2
result can often be approximated by applying criterion 1 provided
an estimate of .sigma. is available.
[0431] With regard to the method of choosing the regularization
factor above, as the window size increased, so did the
regularization factor that gave the minimum .epsilon.. The increase
in window size effectively increases the number of measurements
available to calculate a given B(t). Typically, the best .lambda.
was directly proportional to N until the window became too large to
expect a good a priori estimate of the function behavior from
first-order regularization.
Example 5
Effect of Nonlinearity
[0432] FIG. 39 illustrates the error magnification factor versus
weighting factor for reaction cases a and c when the input contains
1% white noise. For a given input RMS error, the required weighting
factor remained constant as K.sub.m increased, but the output RMS
error decreased as K.sub.m increased. Therefore, if the reaction
term is found to be important in the modeling of the lag between
the blood glucose and subcutaneous tissue glucose in humans, the
inversion will not suffer. Instead, the results are better,
relative to the input error, in the presence of a reaction term
than they are with no reaction at all. The reaction term acts as a
damping term in the forward model. In other words, the term in the
forward model that de-stabilizes the inversion is the derivative of
the subcutaneous glucose concentration with respect to time, while
the reaction term tends to stabilize the inversion.
Example 6
Preparation of Glucose Electrodes
[0433] Glucose electrodes were structurally similar to those
described in Csoregi et al., Anal. Chem., 67:1240-1244 (1995),
incorporated herein by reference. A 0.25 mm gold wire with a 0.04
mm Teflon coating. A 0.09 mm portion of the gold at the end of the
wire was removed, leaving a narrow tube of Teflon. A "wired"
glucose oxidase transduction layer was formed by depositing a
solution of 110 mg/mL of {poly[(1-vinylimidazolyl) osmium
(4,4'-dimethylbipyridine).sub.2Cl]}.sup.+/2+, 10 mg/mL glucose
oxidase (in HEPES 10 mM at pH 8.1), and 2.5 mg/mL poly(ethylene
glycol) mixed in a 78:16:6 wt. % ratio. The solution was deposited
in the Teflon tube to coat the exposed surface of the gold wire.
The electrodes were then rinsed five times and cured at 45.degree.
C. for 15 minutes. A glucose flux restricting layer was formed by
sequentially filling the 0.09 mm deep, 250 .mu.m diameter recess
and curing (at room temperature for 20 min) twice with a 1%
solution of cellulose acetate in cyclohexanone; once with a 0.5%
solution of Nafion (Aldrich, Milwaukee, Wis.) in n-propanol; and
once with a freshly prepared solution of poly (vinyl pyridine)
acetate (PVPA) (25 mg/mL in water) and polyfunctional aziridine
(PAZ) (XAMA-7, E.I.T. Inc., Lakewilie, S.C.) (30 mg/mL in water) in
a 1:2 volume ratio, this layer being cured for at least 8 hr. A
biocompatible layer was then formed of a sensitized 10 wt. %
aqueous tetraacrylated poly(ethylene oxide) solution by
photo-cross-linking (45 sec. UV exposure).
[0434] The in vitro response time of the glucose electrodes was
measured for both increasing and decreasing step changes in glucose
concentration prior to implantation. The measurements were made at
37.+-.0.5.degree. C. in a rapidly stirred, jacketed electrochemical
cell containing pH 7.4 phosphate buffered saline (PBS). The
three-electrode cell had a saturated calomel reference electrode
(SCE), a platinum counter electrode, and the glucose electrode and
was poised at 200 mV vs. SCE. Step changes increasing the glucose
concentration (90 mg/dL to 180 mg/dL) were made by injecting into
the rapidly stirred solution an aliquot of concentrated aqueous
glucose (2M). Step changes decreasing the glucose concentration
(180 mg/dL to 90 mg/dL) were made by injecting PBS into the
cell.
[0435] The intrinsic response times to increasing and decreasing
step changes in glucose concentration were 2.59.+-.1.17 min and
1.55.+-.0.79 min (n=14) respectively.
Example 7
In vivo Experiments
[0436] Male Sprague-Dawley rats, 380-520 g, were preanesthetized
with halothane (Halocarbon Laboratories, North Augusta, SC) and
anesthetized by intraperitoneal injection (0.3 ml) of a equal
volume mixture of acepromazine maleate (10 mg/ml), ketamine (100
mg/ml) and xylazine (20 mg/ml). The animals were then shaven about
the neck, the abdomen, and the area between the scapulae, then
secured on a homeothermic blanket system (Harvard Apparatus, South
Natick Mass.). First the right external jugular vein was located
and cleared of extraneous tissue. The distal side of the right vein
was tied off with 4-0 silk, and a small cut was made in the vein. A
0.0375" diameter medical-grade silastic tube was inserted into the
proximal portion of the right jugular vein and secured with 4-0
silk. A dose of 100-U/kg body wt of heparin solution was then
administered, followed by an equal volume of saline, to clear the
line. Next, the rat's skin was sutured closed. The rat was then
rolled onto its abdomen, while assuring that the line in the
jugular vein was not pulled out, and an electrode was inserted in
the subcutaneous tissue between the scapulae of the animal using a
22-gauge introducing catheter needle (PER-Q-Cath, Gesco, San
Antonio, Tex.). The animal was then returned to its back and
resecured. The left external jugular vein was then located and
cleared of extraneous tissue. Next, the distal side of the left
jugular vein was tied off and a small cut was made in the vein. A
silastic tube of 1.5 cm length was inserted into the proximal side
as a guide, and a glucose electrode was inserted inside the guide
tube. The tube and the sensor were secured with 4-0 silk, with the
electrode's insulating gold wire protruding beyond the end of the
guide tube. The incision site was then moistened and packed with
gauze. An ion-conducting gel was then applied to a skin reference
(Ag/AgCl) electrode, and the electrode was placed on the rat's
abdomen. The implanted electrodes as well as the reference
electrode were connected to a biopotentiostat (13), the output of
which was logged with a data logger.
[0437] After the output of the implanted electrodes reached a
stable baseline (0.5-1 hr), an intravenous injection of 0.5 U/kg of
regular insulin (RU-100, Eli Lilly, Indianapolis, Ind.) was
administered through the right jugular vein. Blood samples were
collected at t=-20, -10, -1, 3, 6, 9, 12, 15, 20, 25, 30, 35, 40,
45, 60, 75, 90 min after the insulin injection. The whole blood
samples were obtained from the left jugular vein and were
immediately placed in tubes containing heparin and sodium fluoride
and kept on ice until analysis. All blood samples were analyzed in
duplicate using a YSI Model 2300 glucose analyzer (YSI, Yellow
Springs, Ohio). At time t=O, the insulin dose was injected through
the infusion catheter and cleared with heparinized saline. At the
end of the experiment the rat was euthanized by sodium
pentobarbital injection i.p. or asphyxiation by CO.sub.2,
consistent with the recommendations of the panel on Euthanasia of
the American Veterinary Association. All in vivo experimentation
was approved by the University of Texas Institutional Animal Use
and Care Committee. The implanted electrodes were sufficiently
glucose selective to be calibrated by withdrawal of a single sample
of blood and assay of its glucose concentration ("one-point in vivo
calibration"). After the current output of the sensor stabilized,
20-40 min after implantation and electrical connection to the
bipotentiostat, a single sample of blood was drawn and its glucose
concentration was assayed using the YSI glucose analyzer. From this
measurement, a current to glucose concentration conversion factor
(mg/dl per nA) was calculated for the implanted electrodes. This
factor was used to obtain all glucose estimates for the remainder
of the test period.
Example 8
Data Analysis
[0438] The onset of the decline in the concentrations of venous and
subcutaneous glucose following the injection of insulin were
determined graphically using the time concentration plots and a
method used in process control to calculate time delays (14). The
tangent line at the point of inflection was drawn (see FIG. 40) and
the line, tracking the basal concentration of glucose prior to the
injection of insulin, was extended. The intersection of the two
lines defined the onset point of the decline. The onset times were
referenced to the time at which insulin was injected. The rate of
decline in glucose concentration in the period between 6 and 20
minutes after insulin injection was calculated by linear regression
analyses for the periodically sampled blood from the vein where
insulin was injected; the contralateral jugular vein, where an
electrode was implanted, and for the subcutaneous interstitial
fluid, where the second electrode was implanted. The values are
presented as means.+-.std, along with their statistical
significance, assessed when appropriate by a Student's t-test for
paired data, with p<0.05 considered as statistically
significant.
[0439] FIG. 41 shows the typical output of the subcutaneous (dotted
line) and jugular vein (solid line) electrodes during an in vivo
experiment. Following insulin injection, the average venous blood
glucose concentrations of the rats (n=7) decreased from 207.+-.67
mg/dl to 59.+-.12 mg/dl. The minimum in blood glucose concentration
was reached 36.6.+-.7.2 minutes after the injection of insulin.
Table 3 lists the average lag times between the lowest subcutaneous
sensor readings and the point of lowest glucose concentration in
the concentrations in the blood withdrawn from the vein where the
insulin was injected, and also between the lowest readings by the
sensor implanted in the contralateral jugular vein and the samples
withdrawn from the injected jugular vein.
3TABLE 3 Declining glucose characteristics. Decline rate Onset time
(mg dl.sup.-1 .multidot. t.sub.minimum glucose Lag time Location
(min) min.sup.-1) (min) (min) Blood samples 3.3 .+-. 0.5 6.8 .+-.
2.0 36.6 .+-. 7.2 -- Intravenous 5.6 .+-. 1.7 7.0 .+-. 2.5 40.3
.+-. 5.9 3.7 .+-. 4.3 sensor Subcutaneous 8.9 .+-. 2.1 3.9 .+-. 1.3
61.2 .+-. 7.5 24.5 .+-. 6.8 sensor
[0440] The onsets of the decline with respect to the time of
injection of insulin, measured in the injected jugular vein, the
contralateral jugular vein and the subcutaneous fluid are also
shown in Table 3, along with the rates of decline in the period
between 6 and 20 min after insulin injection. FIG. 42 shows the
average difference between the estimates of the subcutaneous
glucose concentrations and the actual blood glucose concentrations
as a function of time.
[0441] The nadir in subcutaneous glucose was statistically
different from the nadir in blood glucose (p<0.001) and occurred
24.5.+-.6.8 minutes later. Similarly, the onset of declining
subcutaneous glucose levels (8.9.+-.2.1 min after insulin
injection) was statistically different (p<0.001) from the onset
in blood glucose levels 0.5 min after insulin injection). The rate
of drop in glucose levels, between 6-20 minutes after insulin
injection, was slower in the subcutaneous fluid (3.9.+-.1.3 mg
dl.sup.1 min.sup.-1), than in blood (6.8.+-.2.0 mg dl.sup.-1
min.sup.-1, p=0.003).
[0442] In the contralateral jugular vein, the minimal glucose
concentration was reached 3.7 minutes after it was reached in the
injected vein (36.6 vs. 40.3 min., p=0.06). The rates of decline
during the 6 to 20 minute period were nearly identical in the two
opposite jugular veins (6.8 and 7.0 mg dl.sup.-1 min.sup.-1,
p=0.59). The onsets of the decline in glucose concentrations were
statistically different for the opposite veins (3.3 vs. 5.6 min,
p=0.01).
Example 9
Prediction of Subcutaneous Glucose Concentration
[0443] A typical plot of a prediction of the subcutaneous
concentration of glucose given the concentration of glucose in the
blood from the jugular sensor is shown in FIG. 43, where the only
fitted parameter was .beta.=0.04 min.sup.-1. The uptake term of the
model was found to be negligible for most of the data sets and was
set to zero for all sets. This finding is not surprising, because
the sensor was placed between the connective tissue and smooth
muscle tissue where the rate of glucose uptake is low compared to
the rate of uptake in adipose tissue or skeletal muscle.
[0444] The values of .beta. were determined by a least squares
minimization of the average error for each individual data set and
ranged from 0.04 to 011 min.sup.-1, except for one case, where
.beta. equaled 0.22 min.sup.-1. These results in rats show that
.beta. is relatively constant. If this proves to be true also in
humans then it may not be necessary to determine .beta. for each
patient, or for different subcutaneous placement sites in a
particular patient. Table 4 summarizes the statistics for
comparison of the prediction of the forward model with the
subcutaneous sensor data.
4TABLE 4 Average differences between the measured subcutaneous
glucose concentrations and the predicted subcutaneous glucose
concentrations. Rat Forward Model No Model 1 12.4% 23.8% 2 14.6%
23.5% 3 11.9% 16.2% 4 7.4% 24.1% 5 6.0% 16.0% 6 4.9% 14.3% 7 5.2%
9.6% Mean 8.9% 18.2% std 7.8% 14.5%
[0445] On average the forward model predicted the readings of the
subcutaneous sensor from those in blood with a difference of
8.9.+-.7.8%. If the subcutaneous concentration of glucose were
estimated to equal that measured by the jugular sensor (i.e., if
the model were not used), the average difference would have been
18.2.+-.14.5%. The values derived through the model and those
measured differed and the difference was statistically significant
(p=0.001). In the 40-min interval after injection of the insulin,
the time period that is most in need of correction, the average of
the maximal differences was decreased through the model from 30.7%
to 11.1% (Table 5, p=0.01).
5TABLE 5 Maximum differences between the measured subcutaneous
glucose concentrations and the predicted subcutaneous glucose
concentrations during the 40 minute period following insulin
injection. Rat Forward Model No Model 1 12.5% 22.9% 2 15.9% 34.8% 3
17.7% 21.6% 4 2.7% 50.4% 5 13.9% 35.3% 6 9.7% 35.8% 7 5.2% 14.2%
Mean 11.1% 30.7% Std 5.5% 12.0%
Example 10
Prediction of Blood Glucose Concentration
[0446] The value of B(.tau.) in equations (6-8) was determined as
described above. The weight factor .lambda. was first estimated by
the method described above. The initial condition of B(0)=S(0) was
then enforced within 10% to find a more exact value of .lambda.
based on the initial guess. Further refining of the value of
.lambda. had little effect on the results. Time t=0 for modeling
purposes was taken to be 20 minutes before insulin injection. Plots
of the inverse model predictions are shown in FIG. 44.
[0447] On average, the inverse model predicted in all experiments
the performance of blood glucose concentrations sensed in the
jugular veins, even when the blood and subcutaneous glucose
concentrations were dropping rapidly and from a steady state,
within 11.1.+-.10.6%, as shown in Table 6. If the subcutaneous
concentration of glucose were considered to equal that given by the
jugular sensor (i.e., if the inverse model were not used), the
average difference would have been greater 22.9.+-.14.4%
(p=0.025).
6TABLE 6 Average differences between the measured blood glucose
concentrations and the blood glucose concentrations predicted from
the subcutaneous measurements. Rat Inverse Model No Model 1 13.3%
22.7% 2 13.3% 20.5% 3 13.6% 15.9% 4 14.8% 48.9% 5 8.2% 23.2% 6 7.9%
15.7% 7 6.9% 13.7% Mean 11.1% 22.9% Std 10.6% 14.4%
[0448] Furthermore, during the 40 minute period following insulin
injection, when the dynamic difference was greatest, the maximum
difference between the blood and the subcutaneous glucose
concentrations was 84.1.+-.36.1%. By using the inverse model the
maximum difference between the computed blood glucose concentration
and the actual concentration was reduced to 29.3.+-.8.4% (Table 7,
p=0.006).
7TABLE 7 Maximum differences between the measured blood glucose
concentrations and the blood glucose concentrations predicted from
the subcutaneous measurements during the 40 minute period following
insulin injection. Rat Inverse Model No Model 1 22.8% 72.8% 2 38.8%
67.8% 3 18.0% 41.3% 4 31.5% 157.3% 5 28.5% 94.9% 6 40.9% 72.7% 7
24.8% 82.1% Mean 29.3% 84.1% Std 8.4% 36.1%
[0449] The present invention should not be considered limited to
the particular examples described above, but rather should be
understood to cover all aspects of the invention as fairly set out
in the attached claims. Various modifications, equivalent process,
as well as numerous structures to which the present invention may
be applicable will be readily apparent to those of skill in the art
to which the present invention is directed upon review of the
instant specification. The claims are intended to cover such
modifications and devices.
* * * * *