U.S. patent application number 16/416308 was filed with the patent office on 2019-10-03 for glucose sensors and methods of manufacture thereof.
The applicant listed for this patent is The University of Connecticut. Invention is credited to Fotios Papadimitrakopoulos, Santhisagar Vaddiraju.
Application Number | 20190300925 16/416308 |
Document ID | / |
Family ID | 40562358 |
Filed Date | 2019-10-03 |
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United States Patent
Application |
20190300925 |
Kind Code |
A1 |
Papadimitrakopoulos; Fotios ;
et al. |
October 3, 2019 |
GLUCOSE SENSORS AND METHODS OF MANUFACTURE THEREOF
Abstract
Disclosed herein is a device that functions as a glucose sensor.
The device has a reference electrode; a counter electrode, a
working electrode; an electrically conducting membrane; an enzyme
layer; a semi-permeable membrane; a first layer of a first hydrogel
in operative communication with the working electrode; the first
layer of the first hydrogel being operative to store oxygen;
wherein the amount of stored oxygen is proportional to the number
of freeze-thaw cycles that the hydrogel is subjected to; and a
second layer of the second hydrogel. Disclosed too is a method that
comprises using periodically biased amperometry towards
interrogation of implantable glucose sensors to improve both
sensor's sensitivity and linearity while at the same time enable
internal calibration against sensor drifts that originate from
changes in either electrode activity or membrane permeability as a
result of fouling, calcification and/or fibrosis.
Inventors: |
Papadimitrakopoulos; Fotios;
(West Hartford, CT) ; Vaddiraju; Santhisagar;
(Willimantic, CT) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
The University of Connecticut |
Farmington |
CT |
US |
|
|
Family ID: |
40562358 |
Appl. No.: |
16/416308 |
Filed: |
May 20, 2019 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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14287737 |
May 27, 2014 |
10294507 |
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16416308 |
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12256043 |
Oct 22, 2008 |
8771500 |
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14287737 |
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60999914 |
Oct 22, 2007 |
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Current U.S.
Class: |
1/1 |
Current CPC
Class: |
C01B 13/0248 20130101;
C12Q 1/006 20130101; C01B 13/0259 20130101; C01B 13/0285 20130101;
A61B 5/14532 20130101; C12Q 1/003 20130101 |
International
Class: |
C12Q 1/00 20060101
C12Q001/00; A61B 5/145 20060101 A61B005/145; C01B 13/02 20060101
C01B013/02 |
Goverment Interests
STATEMENT OF FEDERAL SUPPORT
[0002] The present invention was developed in part with funding
from the U.S. Army Research Office under Grant #W81XWH-05-1-0539.
The United States Government has certain rights in this invention.
Claims
1. A method for supplementing oxygen within a sensor, the
supplementing comprising: performing multiple freeze-thaw cycles on
a first layer of a first hydrogel; the sensor comprising: a
reference electrode; a counter electrode; a working electrode; the
working electrode being disposed in the vicinity of the reference
and counter electrode; an electrically conducting membrane; the
electrically conducting membrane being in operative communication
with the working electrode; an enzyme layer; the enzyme layer being
in operative communication with the working electrode; a
semi-permeable membrane; the semi-permeable membrane being in
operative communication with the working electrode; the first layer
of the first hydrogel in operative communication with the working
electrode; the first layer of the first hydrogel being operative to
store oxygen; wherein the amount of stored oxygen is proportional
to the number of freeze-thaw cycles that the hydrogel is subjected
to; and a second layer of a second hydrogel in operative
communication with the working electrode; the second layer of the
second hydrogel comprising tissue response modifying release
agents.
2. The method of claim 1, wherein the first hydrogel is the same as
the second hydrogel.
3. The method of claim 1, wherein the first hydrogel is different
from the second hydrogel.
4. The method of claim 1, wherein the number of freeze-thaw cycles
is about 1 to 100.
Description
CROSS-REFERENCE TO RELATED APPLICATIONS
[0001] This Divisional application claims priority to
Non-Provisional application Ser. No. 14/287,737 filed on May 27,
2014 which is a Divisional of application Ser. No. 12/256,043 filed
on Oct. 22, 2008, now U.S. Pat. No. 8,771,500, which claims
priority to provisional application 60/999,914 filed on Oct. 22,
2007, the entire contents of which are hereby incorporated by
reference.
BACKGROUND
[0003] This disclosure relates to glucose sensors and to methods of
manufacture thereof
[0004] The control of Type I Diabetes Mellitus is generally
effected by the periodic injection of insulin to maintain blood
glucose levels as close to normal as possible. The blood glucose
level is monitored by means of a device that directly measures
glucose from a blood sample. Insulin is injected in the appropriate
quantities and at the appropriate intervals to correct imbalances
in the blood glucose level. Careful control of blood glucose levels
is mandatory for preventing the onset of complications such as
retinopathy, nephropathy and neuropathy. Unfortunately in many
cases, patients neglect to perform regular glucose monitoring and
therefore suffer episodes of hyperglycemia or hypoglycemia, which
may, in turn, lead to the complications listed above or to
death.
[0005] Blood-glucose levels generally vary with activity or food
intake and insulin is therefore administered by sub-cutaneous
hypodermic injection to minimize variations in the blood glucose
levels that generally occur with activity or food intake. Small
externally worn pumps are also available to deliver insulin
percutaneously, thereby replacing the tedious use of a hypodermic
injection, but constant glucose monitoring is still an important
component of control.
[0006] Attempts to develop a closed loop system for the control of
glucose levels have led to the development of ever more
sophisticated insulin pump systems. However, an accurate long lived
implanted blood glucose level monitor that would provide the
desired signal for a closed loop insulin pump control is not yet
available. An implanted blood glucose level monitor hinges on the
accuracy of measuring glycemic levels in diabetic patients, thereby
imposing stringent requirements in the confidence level of the
continuous monitoring technology. In recent years, three kinds of
glucose sensors are being developed: non-invasive,
minimally-invasive and invasive.
[0007] Non-invasive techniques acquire spectroscopic information
through skin or from various body fluids/gases (i.e., saliva,
tears, and breath) and attempt to correlate this with glucose
concentration. Non-invasive techniques generally use explanted
sensors. Minimally-invasive sensors measure glucose concentrations
from fluids obtained from the interstitial tissue of the skin via
microdialysis, iontophoresis, laser ablation, and silicon-based
micro-needle technologies. Both non- and minimally-invasive methods
use elaborate calibration schemes and have considerable
subject-to-subject variability.
[0008] Invasive methods use implanted sensors. These are generally
advantageous in that they exhibit smaller subject-to-subject
variability. However, they are associated with a number of other
problems. In particular, inflammation associated with tissue injury
and the continuous presence of a foreign object is exacerbated by
implant size and the presence of leads or fluid-microcatheters
protruding through the skin. This constitutes the main cause of
sensor failure in vivo, along with sensor element decays due to
long-term usage.
[0009] Tissue injury-based sensor bio-instability is considered to
be a result of the in vivo environment since explanted sensors
often function normally without giving rise to any problems. It is
generally believed that inflammation initiated fibrosis,
calcification, and protein fouling are the leading causes of in
vivo sensor failure. Implantation trauma, lack of biocompatibility
of sensor materials and the physical presence of the sensor in the
tissue are responsible for such tissue responses. Negative tissue
responses (such as, biofouling, inflammation causing fibrosis and
calcification) inhibit analyte migration and hence sensor
performance; long-term sensor stability; and in vivo sensor
calibration. Fibrous encapsulation can deprive the sensor of
adequate blood and analyte supply. This can be modeled by
effectively changing the permeability constants of the membrane(s)
that surrounds the sensing element.
[0010] The D-glucose (dextrose monohydrate) specificity of
analyte-specific enzymes such as glucose oxidase (GO.sub.x), have
helped propel Clark-type electrochemical detection as a major
technological frontier in the development of implantable glucose
sensors. The most commonly used glucose sensors are Clarke-type
amperometric electrochemical sensors and are based on
GO.sub.x-catalyzed oxidation of glucose with O.sub.2, shown in
reaction (1). The principle of detection is based on the
amperometric sensing of hydrogen peroxide (H.sub.2O.sub.2), formed
by the oxidation of glucose. Under an applied potential of 0.7 V
against a silver/silver chloride (Ag/AgCl) reference electrode,
H.sub.2O.sub.2 is electrochemically oxidized according to reaction
(2), and the current produced is related to the concentration of
glucose in the system.
Glucose + O 2 .fwdarw. GO X Glucorolactone + H 2 O 2 ( 1 ) H 2 O 2
.fwdarw. 0.7 V O 2 + 2 H + + 2 e - ( 2 ) ##EQU00001##
[0011] In testing methodology, the sensor is biased continuously at
0.7 V while the change in electrochemical response is measured,
which in turn corresponds to the glucose levels. For the accurate
performance of these sensors it is desirable that the amount of
oxygen present within the sensor geometry must always be equal or
higher than that of the glucose concentration. However, the
dissolved oxygen concentration in ambient or in a biological fluid
sample is significantly lesser than that of the glucose
concentration, leading to an oxygen limiting reaction of the
GO.sub.x enzyme. This results in a saturation of the
electrochemically detected signal, making it impossible to
determine higher levels of glucose in the blood. As a result of
this saturation in the amperometric signal (defined as apparent
Michael's constant K.sub.m.sup.app) any further increase in glucose
concentration does not translate to adequate sensitivity.
[0012] This issue has been addressed by the use of diffusion
limiting outer membranes. These membranes provide a greater
impendence to the larger sized substrate (glucose) as opposed to
the smaller sized co-substrate (O.sub.2). For this, semipermeable
membranes based on NAFION.RTM., polyurethane, cellulose acetate,
epoxy resins, polyether-polyethersulfone copolymer membranes, and
layer by layer (LBL) assembled polyelectrolytes and/or multivalent
cations have been extensively investigated. However, the use of
semipermeable membranes comes at the expense of decreased
sensitivity and increased sensor response time. Furthermore, the
accumulation of exogenous reagents within these membranes (i.e.,
calcification, biofouling, or the like) leads to sensor drifts and
their eventual failure.
[0013] In another variation, an additional oxygen reservoir can be
incorporated into the outer membrane by incorporating
oxygen-absorbing zeolites. Similarly, oxygen reservoirs such as
fluorocarbon based oxygen reservoirs, mineral oils and myoglobin
can be incorporated into the glucose oxidase enzyme layer.
[0014] In another variation, second- and third-generation Clark
type biosensors employ redox mediators and direct `wiring` of
enzymes to electrodes in an attempt to minimize the effect of
O.sub.2. In the case of mediators, their toxicity and
biocompatibility along with the possibility to leach out from the
device to the surrounding tissue present a major problem. Direct
wiring of enzymes to electrodes can minimize the oxygen limitation,
although this modification adds unwanted complexities and higher
expense.
[0015] These defects have been rectified by developing a
polarographic technique for simultaneous measurement of oxygen and
glucose. However, the low sensitivity of the electrode (in the
polarographic technique) to oxygen and the involvement of oxygen in
the oxidation of other interfering species (i.e., ascorbic acid
(AA), acetaminophen (AP), uric acid (UA), and the like) render the
method unsuitable for reliable operation. Independent determination
of glucose and oxygen concentrations could in principle account for
oxygen induced sensor interferences. A number of reports have
attempted to account for these, although addition of other sensor
element adds additional complexities with respect to sensor
integration, testing and calibration.
[0016] As mentioned above, an impediment with Clark-type glucose
sensors is the fact that a number of endogenous species, such as
ascorbic acid (AA), acetaminophen (AP), uric acid (UA), and the
like), also oxidize at the same potential as H.sub.2O.sub.2 (i.e.
0.6-0.7 V), which can add error to the electrochemical signal. High
confidence sensors have to actively account for these species, and
at present not many methodologies have been developed. For example,
anionic charged membranes based on negatively charged polymers
(e.g., NAFION.RTM., polyester sulfonic acid, cellulose acetate, and
the like) have shown to exclude interferences from anionic species
like ascorbic acid, uric acid, and the like, based on the principle
of charge repulsion. However, the large response time associated
with these membranes hinders their usage. Another popular approach
to eliminate interference signals from endogenous species has been
the use of inner, ultra-thin, electropolymerized films between
working electrode and enzyme layer. These films have been shown to
exert partial screening from interference agents to first
generation analyte sensors. However, these electropolymerized films
only minimize signal from endogenous species, and eliminating such
interference has not been realized. Moreover, these membranes do
not possess long term stability, and their interference eliminating
property decreases shortly due to swelling of the polymer.
[0017] In another approach, secondary enzymes (for example
ascorbate oxidase which converts ascorbic acid to dehydroascorbate
and water) have been incorporated in the outer membrane of the
sensor to eliminate the particular species from reaching the
electrode surface and contributing to amperometric current. These
secondary enzymes do however use oxygen as a co-substrate and could
eventually deplete the sensors from the oxygen that is used for the
operation of the primary enzyme (i.e. GO.sub.x). In yet another
approach, independent determination of these interferences using
secondary working electrodes have improved sensor reliability,
although, once again, the addition of another sensor adds
additional complexities involving sensor integration, testing and
failure.
[0018] Another major problem associated with these implantable
sensors is the changes in the electrocatalytic activity of the
working electrodes as well as the in the permeability of the outer
membranes after implantation in the body. While the former is a
result of product adsorption on the surface of the working
electrode, the latter is a result of unwanted accumulation of
exogenous reagents within these membranes (i.e., calcification,
biofouling, and the like). Such factors lead to decrease in
sensitivity, drifts, and to their eventual failure. Moreover,
passivation of working electrodes and inhibition of its
electro-catalytic activity as result of continuous biasing also
leads to saturation in sensor response. To this end, higher applied
potentials, double pulsed amperometry or pulsed amperometric
detection have been the common strategies to renew the surface of
the working electrode even though such techniques are complex to be
applied for miniaturized sensors and implantable sensors with
miniaturized driving electronics. To date there is no reported
methodology to account for such in vivo induced sensor drifts and
the ability to internally calibrate the sensor against these
variations is paramount for long-term sensor operation.
SUMMARY
[0019] Disclosed herein is a device comprising a reference
electrode; a counter electrode; a working electrode; the working
electrode being disposed in the vicinity of the reference and
counter electrode; an electrically conducting membrane; the
electrically conducting membrane being in operative communication
with the working electrode; an enzyme layer; the enzyme layer being
in operative communication with the working electrode; a
semi-permeable membrane; the semi-permeable membrane being in
operative communication with the working electrode; a first layer
of a first hydrogel in operative communication with the working
electrode; the first layer of the first hydrogel being operative to
store oxygen; wherein the amount of stored oxygen is proportional
to the number of freeze-thaw cycles that the hydrogel is subjected
to; and a second layer of the second hydrogel in operative
communication with the working electrode; the second layer of the
second hydrogel comprising tissue response modifying release
agents.
[0020] Disclosed herein too is a method comprising internally
calibrating an electrochemical biosensor based on a primary
reaction when the electrochemical biosensor has reached
equilibrium; monitoring a departure from equilibrium of a secondary
electrochemical reaction; the secondary electrochemical reaction
altering a state of a working electrode of the electrochemical
biosensor; the secondary electrochemical reaction altering the
electrochemistry of the primary reaction.
[0021] Disclosed herein too is a method comprising performing
periodic biasing amperometry on a sensor, the sensor comprising a
reference electrode; a counter electrode; a working electrode; the
working electrode being disposed in the vicinity of the reference
and counter electrode; an electrically conducting membrane; the
electrically conducting membrane being in operative communication
with the working electrode; an enzyme layer; the enzyme layer being
in operative communication with the working electrode; a
semi-permeable membrane; the semi-permeable membrane being in
operative communication with the working electrode; the first layer
of the first hydrogel in operative communication with the working
electrode; the first layer of the first hydrogel being operative to
store oxygen; wherein the amount of stored oxygen is proportional
to the number of freeze-thaw cycles that the hydrogel is subjected
to; and a second layer of a second hydrogel in operative
communication with the working electrode; the second layer of the
second hydrogel comprising tissue response modifying release
agents; the periodic biasing amperometry comprising biasing the
working electrode for a short duration of time at regular intervals
at a number of testing potentials; repeating the periodic biasing
for all the testing potentials; continuing the periodic biasing
until a steady state is attained for all the testing potentials;
conducting an internal calibration of the sensor after an analyte
being measured has reached a steady state; the internal calibration
comprising a time interval where the periodic biasing is not
applied; measuring a periodic biasing amperometric signal
difference immediately before and immediately after the time
interval comprises; measuring a differential for the periodic
biasing amperometric signal difference; comparing the differential
with a calibration chart to obtain sensitivity factors; and
applying the sensitivity factors to the sensor to correct against
drifts.
[0022] Disclosed herein too is a method for supplementing oxygen
within a sensor, the supplementing comprising performing multiple
freeze-thaw cycles on a first layer of a first hydrogel; the sensor
comprising a reference electrode; a counter electrode; a working
electrode; the working electrode being disposed in the vicinity of
the reference and counter electrode; an electrically conducting
membrane; the electrically conducting membrane being in operative
communication with the working electrode; an enzyme layer; the
enzyme layer being in operative communication with the working
electrode; a semi-permeable membrane; the semi-permeable membrane
being in operative communication with the working electrode; the
first layer of the first hydrogel in operative communication with
the working electrode; the first layer of the first hydrogel being
operative to store oxygen; wherein the amount of stored oxygen is
proportional to the number of freeze-thaw cycles that the hydrogel
is subjected to; and a second layer of a second hydrogel in
operative communication with the working electrode; the second
layer of the second hydrogel comprising tissue response modifying
release agents.
BRIEF DESCRIPTION OF THE FIGURES
[0023] FIG. 1 is a schematic representation of a modified
amperometric glucose sensor, along with various chemical,
electrochemical and diffusion processes associated with its
operation. The glucose oxidase (GO.sub.x) layer is coated with a
semi-permeable membrane to reduce the amount of glucose entering
the sensor. The hydrogel coating shows embedded microspheres at
different stages of degradation and release of tissue response
modifying (TRM) agents;
[0024] FIG. 2 is a graphical amperometric response for an
electrochemical sensor as function of voltage-biasing duration.
This response exhibits two operational regimes;(A) non-equilibrated
regime and (B) equilibrated regime. Upon cessation of voltage
biasing along with exposure to redox-active agents (i.e.,
H.sub.2O.sub.2, O.sub.2, H.sub.2O, and the like) the amperometric
response of the sensor will start shifting upwards along the
indicated arrow. The departure from the equilibrium will depend on
the cessation duration of the voltage-bias and concentration of
redox-active agent(s);
[0025] FIG. 3 is a graphical response depicting the applied
bias/measuring time sequence employed in the periodically-biased
amperometric testing of glucose sensors;
[0026] FIG. 4 is a graphical calibration sequence for assessing
changes in the permeability of semi-permeable membranes for
prolonged in vivo sensor use. The top panel illustrates the
periodic bias sequence for performing this calibration. The bottom
panel depicts the expected amperometric response, with the
permeability constants inversely proportional to the difference
between responses S.sub.15 and S.sub.16 for a given t.sub.cal
interval;
[0027] FIG. 5 is a graphical calibration routine for assessing
changes in the electrocatalytic activity of the sensor electrodes.
Following an internal calibration sequence similar to that of the
FIG. 4, a feeding event is conducted. Upon glucose-level
equilibration following the feeding event, a calibration sequence
similar to FIG. 4 is commenced. Knowing that the permeability of
semi-permeable membranes does not change substantially, the change
in the assay reading is proportional to the electrocatalytic
activity of the sensor electrodes;
[0028] FIG. 6 is a cyclic voltammogram of an electrochemical sensor
containing a composite of sensing elements. The sequential
interrogation of this sensor at various biasing potentials (shown
by broken lines), where different analytes contribute to the
amperometric signal to different extents, provide the means to
accessing the individual concentrations of various analytes;
[0029] FIG. 7 is a graphical schematic of the applied
bias/measuring time sequence employed in multi-analyte detection,
using periodically-biased amperometric testing at various
potentials;
[0030] FIG. 8 is an ultraviolet-visible-near infrared (UV-Vis-NIR)
absorption spectra of a poly(ortho-phenylene diamine) (PPD) film as
a function of the applied biasing voltage;
[0031] FIG. 9 is a UV-Vis absorption spectra of poly phenylene
diamine (PPD) film, first biased for 150 sec at 0.2 V followed by
removal of the biasing voltage and exposure to various
concentration of H.sub.2O.sub.2;
[0032] FIG. 10 is a graphical response vs. glucose concentration of
a Pt/PPD/GO.sub.x/(HAs--Fe.sup.3+).sub.5 glucose working electrode
vs. a Ag/AgCl reference electrode, when tested in continuous and
periodically biased amperometry. The periodically biased
amperometry is carried out at the initial stages of Regime A in
FIG. 3;
[0033] FIG. 11 is a graph depicting a 1-second periodic-biased
amperometric response as a function of t.sub.wait for a
Pt/PPD/GO.sub.x/(HAs--Fe.sup.3+).sub.5 working electrode operated
at 0.7 V vs. a Ag/AgCl reference electrode on a constant glucose
concentration of 2 mM;
[0034] FIG. 12 is a graph depicting a continuous biased
amperometric response of a Pt/PPD/GO.sub.x/(HAs--Fe.sup.3+).sub.5
working electrode, biased at 0.7 V vs. a Ag/AgCl reference
electrode, in the presence and absence of top PVA layer that has
been subjected to three freeze-thaw cycles;
[0035] FIG. 13 is graph depicting a variation of PVA-stored O.sub.2
content as a function of the number of freeze thaw cycles;
[0036] FIG. 14 is a cyclic voltammogram of a Pt/PPD+SWNT working
electrode versus a Ag/AgCl reference electrode in PBS buffer
solution that has been saturated with either air (i.e. O.sub.2)
(dashed line) or N.sub.2 (solid line). For purpose of clarity, the
inset shows a blowup of the results in the 0.1 to -0.2 Volt
region;
[0037] FIG. 15 is graph depicting a continuously biased
amperometric response of a Pt/PPD+SWNT working electrode, biased at
-0.1 V versus a Ag/AgCl reference for three successive cycles of
addition and removal of O.sub.2; and
[0038] FIG. 16 is a schematic representation of the process used to
modify a surface of a working electrode with a composite containing
a network of SWNTs intercalated with a layer of an
electropolymerized conducting polymer.
DETAILED DESCRIPTION
[0039] Disclosed herein is an implantable glucose sensor
(hereinafter sensor or biosensor). The sensor comprises a working
electrode in operative communication with an electrically
conducting membrane, an enzyme layer, a semi-permeable membrane, a
first layer of a first hydrogel and a second layer of the second
hydrogel.
[0040] In one embodiment, the working electrode comprises a metal
upon which is disposed an electrically conducting polymer and an
enzyme specific to an analyte of interest (hereinafter the
"analyte"), a layer-by-layer film to fine-tune permeability to the
analyte, a poly(vinyl alcohol) hydrogel layer to store and provide
additional oxygen to the sensor, and a biocompatible coating that
are also capable of releasing a variety of drugs. The biosensor is
advantageous over other comparative biosensors in that it (a)
exhibits high linearity; (b) exhibits high sensitivity; (c) takes
into account the contribution of exogenous interfering species; and
(d) provides internal calibration routines to take into account
sensor drifts based on in vivo induced effects that change the
permeability of semi-permeable membrane. It also accounts for
gradual decay of electrode activity.
[0041] Disclosed too is a method that comprises using periodically
biased amperometry towards interrogation of implantable glucose
sensors to improve both sensor's sensitivity and linearity while at
the same time enable internal calibration against sensor drifts
that originate from changes in either electrode activity or
membrane permeability as a result of fouling, calcification and/or
fibrosis. This method involves the application of a biasing voltage
to the working electrode with respect to the neighboring reference
electrode for a short duration of time, at controlled intervals.
This reduces sensor stressing and enhances long-term stability
while at the same time provides better power management and signal
to noise ratio. Variations in bias duration and time intervals
allow us to modulate the electro-catalytic activity of the working
electrode, herein termed as "action". This action is afforded by
varying the redox state of the working electrode through the
application of specific bias and time duration. The redox state of
the working electrode is however reversely affected by the amount
of time and concentration of H.sub.2O.sub.2 that is adjacent to the
electrode, which constitutes a "counter-action" to bias. As it
turns out, the concentration of H.sub.2O.sub.2 concentration is
related to both sub-cutaneous (s.c.) tissue concentration of
glucose and the permeability coefficient of semi-permeable
membranes adjacent to the electrodes. At constant glucose
concentration, the competition of "action" and "counter-action"
provides us with the ability to decipher and quantify sensor drifts
originating from changes in the permeability of semi-permeable
membranes. In a similar manner, the same membrane permeability and
different glucose concentration also enables the determination of
electrode activity. The combination of these two routines provides
the means to internally re-calibrate the implantable sensor against
drifts and avoid frequent external calibrations. By varying bias
voltage, periodically-biased amperometry together with the
aforementioned "action" and "counter-action" from various
electroactive analytes (i.e. oxygen, uric acid, acetaminophen,
ascorbic acid) can also be utilized to enable their simultaneous
detection along with glucose.
[0042] FIG. 1 depicts an exemplary configuration of the biosensor
100, which comprises a working electrode 102 in operative
communication with an electrically conducting membrane 106, an
enzyme layer 110, a semi-permeable membrane 114, a first layer of a
first hydrogel 118 and a second layer of the second hydrogel 122.
As can be seen in the FIG. 1, the second layer of the second
hydrogel composite 122 contacts tissue 126 in a living being.
Opposed to the working electrode is a reference electrode 202.
[0043] The working electrode 102 generally comprises a metal. In an
exemplary embodiment, the metal is an inert metal. Examples of the
metal are platinum, gold, palladium, or the like, or a combination
comprising at least one of the foregoing metals. Alternatively, the
working electrode can comprise carbon. In an exemplary embodiment,
the working electrode 102 comprises platinum. The working electrode
102 is opposedly disposed next to reference electrode 202.
[0044] The working electrode 102 has an area of about 0.1 square
millimeters (mm.sup.2) to about 100 mm.sup.2. In a preferred
embodiment, the working electrode 102 has a thickness of about 0.2
mm.sup.2 to about 0.3 mm.sup.2. Alternatively, the area of the
working can be smaller than 0.1 mm.sup.2.
[0045] As noted above, the working electrode 102 is in operative
communication with an electrically conducting membrane 106. It is
desirable for the electrically conducting membrane 106 to prevent
the diffusion of a number of endogenous species like ascorbic acid,
uric acid and acetaminophen. In one embodiment, the working
electrode 102 is in physical communication with the electrically
conducting membrane 106. In an exemplary embodiment, the
electrically conducting membrane 106 is disposed upon and in
intimate contact with the working electrode 102.
[0046] The electrically conducting membrane 106 undergoes redox
changes depending on the time and duration of the applied voltage
as well as the concentration of various soluble redox species that
are in its immediate vicinity. In one embodiment, the electrically
conducting membrane 106 is an electrically conducting nanocomposite
that affords sensitivity to more than one analyte at a various
testing potentials.
[0047] The electrically conducting membrane can comprise
intrinsically electrically conducting polymers and copolymers or
polymers that are made electrically conducting by virtue of being
filled with a percolating network of electrically conducting
particles.
[0048] Intrinsically electrically conducting polymers are
polypyrrole, polyaniline, polythiophene, polyacetylene,
polyphenylene diamine, poly(3,4-ethylenedioxythiophene)
poly(styrenesulfonate), sulfonated poly aniline, sulfonated
polypyrrole, poly(ethylene dioxythiophene),
poly(ethylenedioxypyrrole), poly(p-phenylene vinylene),
polycarbazole, substituted polycarbazole, polyindole, or the like,
or a combination comprising at least one of the foregoing
intrinsically electrically conducting polymers.
[0049] The intrinsically conducting polymer can be copolymerized
with other insulating organic polymers. Examples of organic
polymers that can be copolymerized with the intrinsically
conducting polymer are polyacetals, polyacrylics, polycarbonates
polystyrenes, polyesters, polyamides, polyamideimides,
polyarylates, polyacrylates, polymethylmethacrylates,
polyarylsulfones, polyethersulfones, polyphenylene sulfides,
polyvinyl chlorides, polysulfones, polyimides, polyetherimides,
polytetrafluoroethylenes, polyetherketones, polyether etherketones,
polyether ketone ketones, polybenzoxazoles, polyoxadiazoles,
polybenzothiazinophenothiazines, polybenzothiazoles,
polypyrazinoquinoxalines, polypyromellitimides, polyquinoxalines,
polybenzimidazoles, polyoxindoles, polyoxoisoindolines,
polydioxoisoindolines, polytriazines, polypyridazines,
polypiperazines, polypyridines, polypiperidines, polytriazoles,
polypyrazoles, polypyrrolidines, polycarboranes,
polyoxabicyclononanes, polydibenzofurans, polyphthalides,
polyacetals, polyanhydrides, polyvinyl ethers, polyvinyl
thioethers, polyvinyl alcohols, polyvinyl ketones, polyvinyl
halides, polyvinyl nitriles, polyvinyl esters, polysulfonates,
polysulfides, polythioesters, polysulfones, polysulfonamides,
polyureas, polyphosphazenes, polysilazanes, polysiloxane,
polyolefins, or the like, or a combination comprising at least one
of the foregoing organic polymers.
[0050] As noted above, the electrically conducting membrane 106 can
comprise an electrically insulating organic polymer that is filled
with electrically conducting filler. Examples of electrically
conducting fillers are carbon nanotubes, carbon black, carbon
nanoparticles, nanorods, intrinsically electrically conducting
polymer powders, metal powders, electrically conducting ceramic
powders, or the like, or a combination comprising at least one of
the foregoing electrically conducting fillers. Other fillers that
can be used in the electrically conducting membrane 106 are
nano-sized inorganic compounds, nano-sized inorganic (e.g.,
TiO.sub.2), Au, Ag, Rd, Pd, or Pt nanoparticles, SnO.sub.2
nanoparticles, SnO.sub.2 nanorods, SiO.sub.x nanoparticles, or the
like, or a combination comprising at least one of the foregoing
nanoparticles.
[0051] In one embodiment, the electrically conducting membrane 106
can comprise conducting polymers that are copolymers of
(3,4-dihydroxy-L-phenylalanine), hydroxyquinones, ferrocene and
ferrocene derivatives, ferricyanide,
tetrathiafulvalene-tetracyanoquinodimethane, osmium salts,
phenothiazine, phenoxazine, porporphorins, flavins, pyroloquinoline
quinines, or the like, or a combination comprising at least one of
the foregoing copolymers.
[0052] In another embodiment, the electrically conducting membrane
106 can comprise redox enzymes; the redox enzymes being horseradish
peroxidase, myoglobin, glucose dehydrogenase, or the like, or a
combination comprising at least one of the foregoing redox
compounds.
[0053] In one embodiment, the electrically conducting membrane 106
can comprise redox enzymes in an amount of about 1 to about 99
weight percent (wt %), specifically about 2 to about 95 wt %, and
more specifically about 5 to about 80 wt %, based on the total
weight of the electrically conducting membrane.
[0054] The electrically conducting membrane 106 can be spin coated,
crosslinked, inkjet printed and patterned on top of the working
electrode 102. In one embodiment, the inkjet printed nanocomposite
is crosslinked. The crosslinking can be attained by inkjet printing
crosslinking agents or the crosslinking can be conducted by
immersing the device into crosslinking agents.
[0055] In an exemplary embodiment, the electrically conducting
membrane 106 can be constructed by electropolymerizing a thin layer
of ortho-phenylene diamine (OPD) to yield poly(ortho-phenylene
diamine) (PPD). In another exemplary embodiment, the electrically
conducting membrane 106 can be manufactured by electropolymerizing
a thin layer of PPD in the presence of electrically conducting
nanotubes and/or nanorods. The nanotubes, nanowires and/or the
nanorods are embedded in the thin layer of PPD. Examples of
nanotubes are multiwall carbon nanotubes (MWNTs), single wall
carbon nanotubes (SWNTs), or a combination comprising at least one
of the foregoing carbon nanotubes. Examples of nanorods are
aluminum nanorods, copper nanorods, or the like, or a combination
comprising at least one of the foregoing nanorods.
[0056] In one embodiment, the electrically conducting membrane 106
can comprise nanoparticles or nanotubes in an amount of about 1 to
about 99 weight percent (wt %), specifically about 2 to about 95 wt
%, and more specifically about 5 to about 80 wt %, based on the
total weight of the electrically conducting membrane.
[0057] In yet another exemplary embodiment, the electrically
conducting membrane 106 can be realized by first assembling a
plurality of shortened single-walled carbon nanotubes and
subsequently electropolymerizing around it a thin layer of PPD.
[0058] The electrically conducting membrane 106 has a thickness of
about 5 to about 100 nanometers. In a preferred embodiment, the
electrically conducting membrane 106 has a thickness of about 10 to
about 20 nanometers.
[0059] The electrically conducting membrane 106 is in operative
communication with an enzyme layer 110. In one embodiment, the
enzyme layer 110 comprises glucose oxidase, lactate oxidase, poly
vinyl alcohol (PVA), bovine serum albumin, or the like, or a
combination comprising at least one of the foregoing materials. In
another embodiment, the enzyme layer 110 is crosslinked with
glutaraldehyde. The enzyme layer 110 may comprise a conductive
polymer if desired. In an exemplary embodiment, the enzyme layer
110 is a glucose oxidase (GO.sub.x) enzyme layer 110.
[0060] In one embodiment, the electrically conducting membrane 106
is in physical communication with a glucoseoxidase (GO.sub.x)
enzyme layer 110. In another embodiment, the glucoseoxidase
(GO.sub.x) enzyme layer 110 contacts a surface of the electrically
conducting membrane 106 that is opposed to the surface in contact
with the working electrode 102. The glucoseoxidase (GO.sub.x)
enzyme layer 110 is immobilized on the electrically conducting
membrane and is hence referred to as the immobilized glucoseoxidase
(GO.sub.x) enzyme layer 110.
[0061] Within the immobilized glucoseoxidase (GO.sub.x) enzyme
layer 110, glucose reacts with oxygen (O.sub.2) to produce hydrogen
peroxide in accordance with reaction (1) detailed above. The
generated hydrogen peroxide is anodically (with respect to the
reference electrode 202) detected at the working electrode 102.
[0062] In one embodiment, the enzyme layer 110 is
electropolymerized on top of the electrically conducting membrane
106. In another embodiment, the electrically conducting membrane
106 and the enzyme layer 110 are electropolymerized
concurrently.
[0063] The immobilized glucoseoxidase (GO.sub.x) enzyme layer 110
has a thickness of about 1 nanometer to about 1,000 micrometers. In
a preferred embodiment, the immobilized glucoseoxidase (GO.sub.x)
enzyme layer 110 has a thickness of about 10 nanometers to about
100 micrometers.
[0064] In order to regulate the amount of glucose with respect to
oxygen and ensure better sensor linearity, the immobilized
glucoseoxidase (GO.sub.x) enzyme layer 110 is in operative
communication with a semi-permeable membrane 114. In one
embodiment, the immobilized glucoseoxidase (GO.sub.x) enzyme layer
110 is in physical communication with a semi-permeable membrane
114. In another embodiment, the immobilized glucoseoxidase
(GO.sub.x) enzyme layer 110 is disposed upon and in intimate
contact with a semi-permeable membrane 114. The immobilized
glucoseoxidase (GO.sub.x) enzyme layer 110 is disposed upon a
surface of the semi-permeable membrane that is opposed to the
surface that contacts the electrically conducting membrane 106.
[0065] The semi-permeable membrane 114 comprises alternating layers
of positive and negative polyion species (i.e., polymers, oligomers
and/or multi-valent cations) stacked in a layer-by-layer (LBL)
fashion. Variations in the number of LBL-deposited bi-layers have
been shown to regulate the inward diffusion of glucose and outward
diffusion of hydrogen peroxide.
[0066] The semipermeable membrane 114 can comprise a plurality of
alternating layers of a poly acid and metal ions. The alternating
layers are also termed multilayers. The poly acid can be a
polymeric acid or a non-polymeric inorganic acid. In one
embodiment, the poly acid is humic acid while the metal ions are
Fe.sup.3+.
[0067] The semipermeable membrane can also comprise polystyrene
sulfonate, polydimethyl diallyl ammonium chloride,
polyethyleneamine, hyaluronic acid, polyaspartic acid, polylysine,
chitosan, collagen, or the like, or a combination comprising at
least one of the foregoing materials. The semipermeable membrane is
manufactured through layer-by-layer assembly. In one embodiment,
the semipermeable membrane is patterned on top of the working
electrode. In another embodiment, the semipermeable membrane is ink
jet printed in layer-by-layer fashion with intermediate washing
steps.
[0068] The semipermeable membrane comprises 1 multilayer to 1,000
multilayers, specifically about 3 to about 100 multilayers, and
more specifically about 5 to about 10 multilayers.
[0069] The semi-permeable membrane 114 has a thickness of about 2
to about 1000 nanometers. In a preferred embodiment, the
semi-permeable membrane 114 has a thickness of about 2 to about 100
nanometers.
[0070] In order to immobilize and locally deliver various tissue
response modifying (TRM) agents that control and suppress
inflammation of the surrounding tissue, while at the same time
permitting passage of glucose and O.sub.2, a hydrogel coating can
be incorporated on the surface of the sensor that contacts the
surface of the tissue 106. In one embodiment, the hydrogel coating
comprises a first layer of a first hydrogel 118 and a second layer
of the second hydrogel 122.
[0071] As can be seen in the FIG. 1, the first layer of hydrogel
118 is in operative communication with the semi-permeable membrane
114, while the second layer of hydrogel 122 is in operative
communication with the first layer of hydrogel 118. In another
embodiment, the first layer of hydrogel 118 is in physical
communication with the semi-permeable membrane 114, while the
second layer of hydrogel 122 is in physical communication with the
first layer of hydrogel 118.
[0072] In an exemplary embodiment, the first layer of hydrogel 118
is disposed upon and in intimate contact with a surface of the
semi-permeable membrane 114 that is opposed to the surface in
contact with the immobilized glucose oxidase (GO.sub.x) enzyme
layer 110. The first layer of hydrogel 118 generally comprises a
water-soluble polymer that can absorb oxygen. It is desirable for
the first layer of hydrogel 118 to be crosslinked. In one
embodiment, the first layer of hydrogel 118 is crosslinked by
freeze-thaw pumping. The number of freeze-thaw pumping cycles can
be varied.
[0073] In one embodiment, the number of freeze-thaw pumping cycles
can be varied from about 1 to about 25 cycles. In an exemplary
embodiment, the number of freeze-thaw pumping cycles can be varied
from about 1 to about 7 cycles.
[0074] The first hydrogel can be the same or different as the
second hydrogel. Examples of the first and second hydrogels are
crosslinked polyhydroxyethylmethacrylate, polyethylene oxide,
polyacrylic acid, polyvinylpyrrole, chitosan, collagen, or the
like, or a combination comprising at least one of the foregoing
hydrogels.
[0075] In one embodiment, the first layer of the first hydrogel 118
and the second layer of the second hydrogel 122 both comprise
polyvinylalcohol (PVA). The PVA ensures a homogeneous coverage of
the immobilized glucoseoxidase (GO.sub.x) enzyme layer 110. It also
facilitates the storage of O.sub.2. In an exemplary embodiment, the
amount of stored O.sub.2 is controlled by varying the number of
freeze-thaw cycles for the PVA. In another exemplary embodiment,
this PVA layer can be loaded with various oxygen storing enzymes
(e.g., myoglobin) and oxygen producing enzymes (e.g.,
catalase).
[0076] The hydrogel membrane is spun coated, crosslinked and
patterned on top of the working electrode. In one embodiment, the
hydrogel membrane is inkjet printed on top of the working
electrode.
[0077] The second layer of the second hydrogel 122 is in operative
communication with the first layer of the first hydrogel 118. In
one embodiment, the second layer of the second hydrogel 122 is in
physical communication with the first layer of the first hydrogel
118. In an exemplary embodiment, the second layer of the second
hydrogel is disposed upon and in intimate contact with a surface of
the first layer of the first hydrogel 118 that is opposed to the
surface in contact with the semi-permeable membrane 114.
[0078] The second layer of the second hydrogel 122 contains tissue
response modifying (TRM) release agents. The TRM can be a composite
of PVA and TRM containing PLGA microspheres. The second layer of
hydrogel can also be crosslinked by freeze-thaw pumping. In one
embodiment, the second layer of hydrogel 122 is a composite that
comprises a water-soluble polymer in addition to TRM microspheres.
In another embodiment, the water-soluble polymer of the second
layer of hydrogel 122 is also polyvinylalcohol. A surface of the
second layer of hydrogel 122 generally contacts the tissue 126 of a
living being. The gradual release of TRM is assisted by the
degradation of microspheres (7) that contain various drugs. As can
be seen in the FIG. 1, the TRM releases these drugs over a period
of time (see 7, 8 and 9 in the FIG. 1). The concentration of the
TRM in the second layer of hydrogel 118 can be varied.
[0079] In general, the first layer of hydrogel 118 and/or the
second layer of hydrogel 122 can comprise a variety of enzymes to
eliminate endogenous species. Examples of the enzymes are catalase,
transferase, hydrolase, oxidase, peroxidase, kinases, superoxidase,
phosphatase, transferase, hydrolase, pyrophosphatase, oxygenase,
nuclease, lipase, peptidase, trancacetylase, hydroxylase,
dioxygenase, dehydrogenase, carboxylase, aminase, catalase,
phosphohydrolase, diaminase, reductase, synthase, kinase, caspase,
methionine synthase, cystathionase, or the like, or a combination
comprising at least one of the foregoing enzymes.
[0080] In one embodiment, the PVA layer can contain a variety of
different additives. Examples of such additives are myoglobin,
nanotubes, nanorods, or the like, or a combination comprising at
least one of the foregoing additives. In one embodiment, the
combination of catalase and myoglobin is varied in an amount of
about 1 to about 99 weight percent (wt %), based on the total
weight of the first layer of hydrogel 118 and/or the second layer
of hydrogel 122.
[0081] In general in comparative devices, the sensor is subjected
to an amperometric testing methodology that relies on a continuous
biasing of the working electrode. This testing is conducted to
determine the calibration of the sensor, which in turn dictates the
working of the sensor and the infusion of insulin or glucose into
the body of a living being. In general, in order to effect
calibration and functioning of the sensor in this manner, the
electrical current response with time is used to determine the
amount of insulin or glucose that should be injected into the
living being.
[0082] It is believed that this testing methodology ensures that
all transient effects have been eliminated and proper diffusion
gradients have been established between the reactants and products
of the electrochemical half reaction that are brought about by the
biosensor. However, continuous biased amperometry subjects the
sensor to sensor-stress that eventually leads to signal decay. This
signal decay occurs because of drift in readings at the working
electrode. However, additional drifts can originate from changes in
permeability of the semi-permeable membranes that are located
between the working electrode and the tissue as well as because of
changes in the electro-activity of the working electrode.
[0083] In general, continuous biased amperometry leads to a much
smaller signal that is the result of the reduced availability of
the analyte(s), over saturation of the half reaction byproduct(s)
and/or suppressed electroactivity of the electrode due to the
higher presence of these byproduct(s) is produced. As a result,
this smaller signal is generally compensated for by using a (i)
higher applied potential (ii) double pulsed amperometry and/or
(iii) pulsed amperometric detection. Each of these methods of
compensation, however, uses higher voltages (which exacerbate
signals from exogenous species). In addition, double pulse
amperometry utilizes complex driving electronics, difficult to
attain for miniaturized implantable devices.
[0084] An exemplary response of an unused new sensor (or a sensor
that has not been operated for a long period of time) that is
subjected to this type of amperometric testing involving voltage
biasing as function of time is illustrated in FIG. 2. This response
exhibits two operational regimes: (A) non-equilibrated regime that
shows a rapid "run in" signal decay followed by a gradual
equilibration to enter regime B, where its response is
equilibrated. The signal displayed in the FIG. 2 can be a response
to exposure to any redox-active agents such as H.sub.2O.sub.2,
O.sub.2, H.sup.+, and the like. Upon termination of voltage
biasing, along with exposure to redox-active agents (i.e.
H.sub.2O.sub.2, O.sub.2, H.sub.2O, and the like), the amperometric
response of the sensor gradually starts moving backwards to the
non-equilibrated regime A. This departure from the equilibrium
depends on the time duration for which the voltage biasing has been
switched off. The departure from the equilibrium also depends on
the concentration of redox-active species in the vicinity of the
electrode. Since these redox-active species are in constant
equilibration with the subcutaneous tissue (e.g., 126 in the FIG.
1) (via diffusion through a number of semi-permeable membranes,
which also includes those generated by an immuno-response) it
provides a means of accessing and assessing the diffusion
characteristics of the membrane while the sensor is in an in vivo
operation.
[0085] In order to avoid the signal run-up and the subsequent decay
during calibration of the sensor, it is desirable for the sensor to
be operated in a periodic biasing mode with time periods long
enough to allow for equilibration. FIG. 3, illustrates one
exemplary periodic biasing operation which permits equilibration of
the sensor so that a new normal sensor operation state is
developed. This operation state is used to avoid the aforementioned
run-ups and decays that are depicted in the FIG. 2.
[0086] For purposes of simplicity, responses are depicted for
glucose concentration kept the same throughout the time period of
experimentation shown in FIG. 3. Throughout the amperometric
testing that includes sensor biasing, the biasing steps are
performed for a short duration of time (e.g., a period of about 1
second) and preferably kept the same throughout the sensor
operation, unless otherwise stated. Initially, the sensor response
has to be "partially" equilibrated, as depicted in regime A of FIG.
2. This is performed with periodic biasing (as depicted by
t.sub.eq) shown in FIG. 3. The resulting signals from the periodic
biasing steps correspond to sections of the current-time curve
shown in FIG. 2, regime A. From the FIG. 3 (steps P.sub.1 through
P.sub.8), it can be seen that the signal at the end of a given
biasing period is slightly lower than the initial signal value of
the next biasing period. This effect arises from the aforementioned
exposure of the sensor to the redox-active species that are
constantly replenished at the vicinity of the electrode and
therefore attempt to relax the sensor to its initial state.
Eventually, after a number of equilibration biasing steps (herein
shown in nine steps, P.sub.1 to P.sub.9) the sensor response
reaches a steady state. At this point normal sensor operation (or
equilibration) (P.sub.10 to P.sub.12), is established and after the
establishing of this equilibration, the interval between biasing
steps can be varied or controlled to reduce sensor stress and
consequent error.
[0087] In an exemplary embodiment, the interval (e.g., t.sub.wait
as seen between P.sub.10 to P.sub.12 in the FIG. 3) between biasing
steps (can be increased from the intervals established initially
(e.g., t.sub.eq as seen between P.sub.1 to P.sub.9) to reduce
sensor stress. In another exemplary embodiment, the interval
between biasing steps can be decreased from the intervals
established initially to reduce sensor stress. When sensor stress
and the sequence of signal run-ups and decays are reduced as much
as possible, it is then desirable for a computer processing unit
(that it in communication with the sensor) to direct the sensor to
begin calibration operations. This will be discussed below.
[0088] FIG. 4 illustrates one embodiment of a sensor biasing
sequence in order to perform an internal calibration routine to
assess changes in permeability of the membrane(s) that are disposed
upon the sensor as seen in the FIG. 1. It is first desirable to
assess a time regime where glucose concentration levels are fairly
constant. In general, this time regime occurs during periods of
extensive rest where the patient has not consumed a meal prior to
the testing. Such a constant glucose concentration regime can be
established by assessing the similarity of responses shown for
biasing steps P.sub.13 to P.sub.15 and as described in the
aforementioned paragraphs. Assuming that the end points of
S.sub.13, S.sub.14, and S.sub.15 are fairly constant, then a
computer processing unit can instruct the sensor to proceed to the
calibration state.
[0089] For this, the sensor is left unbiased for an extended period
of time (e.g., t.sub.cal=nt.sub.wait, where n varies from about 2
to about 10, where t.sub.wait is the waiting time period between
measurements and t.sub.cal is the calibration time where the
periodic biasing is switched off). During this time, redox-active
species act on the sensor and increase changes its electrical
current response upwards along the lines of the current-time curve
shown in regime A of the FIG. 2. At the end of t.sub.cal, the
biasing step P.sub.16 interrogates in the relaxed sensor and
records the S.sub.16 sensor response. By comparing the end points
of S.sub.15 and S.sub.16, the magnitude of the departure from
equilibrium is established (shown by two horizontal lines in FIG.
4).
[0090] In the case depicted in the FIG. 4, the end-point at the
equilibrium state for S.sub.16 is greater than the end-point at the
equilibrium state for S.sub.15. The difference between the
end-points is considered to be a positive difference and can be
used to facilitate a calibration of the sensor. In a similar manner
it is possible for the end-point at the equilibrium state for
S.sub.16 to be lower than the end-point at the equilibrium state
for S.sub.15. This difference will be negative and can also be used
to facilitate calibration of the sensor.
[0091] At this point, the sensor is brought back to equilibrium to
assess if the glucose concentration remained the same during the
calibration step. Rapid equilibration takes place in an equivalent
fashion to sensor equilibration following implantation, as shown in
FIG. 3 (S.sub.1 to S.sub.9). For example, this rapid equilibration
is shown within five biasing steps (P.sub.16 to P.sub.20) in FIG.
4. Subsequently, normal sensor operation is established as shown by
biasing steps P.sub.21 to P.sub.23. Assuming that the sensor
response indicated by the end points of S.sub.21 to S.sub.23 curves
is comparable to end points of S.sub.13 to S.sub.15 curves, then
the calibration routine is accepted. If the latter in not true,
then the calibration routine is not accepted and the driving
computer processing unit is instructed to seek another constant
glucose regime to re-perform this calibration.
[0092] Upon acceptance of the calibration routine, the driving
computer processing unit stores the recorded difference between the
end point of S.sub.15 and S.sub.16 curves (A.sub.n, where
n=t.sub.cal/t.sub.wait and compares it to a calibration chart that
is already stored in its memory. This chart has been established by
a careful in vitro calibration study described below.
[0093] Starting with a sensor having a Pt/PPD/GO.sub.x/(LBL).sub.m
configuration, as represented by the layers 102, 106, 110 and 114
of the FIG. 1, the A.sub.n,m values (where n=t.sub.cal/t.sub.wait
and m corresponds to the number of LBL bilayers) will be determined
as a function of constant glucose concentration (i.e.,
S.sub.15=S.sub.23) over the entire physiological glucose range (of
about 2 to about 22 millimolar (mM)).
[0094] Subsequently, the number of LBL bilayers can be varied and a
similar study conducted to determine the A.sub.n,m+1 values to
emulate in vivo induced pore clogging of semipermeable membranes.
An independent determination of the permeability coefficients
(D.sub.m) of these (LBL).sub.m membranes, will permit a correlation
of A.sub.n,m values with D.sub.m for glucose and H.sub.2O.sub.2 and
derive an empirical relationship between A.sub.n,m and related
sensitivity factors. This empirical function will be fed into the
operating program of the driving computer processing unit.
Following an internal calibration routine, the computer will assess
the obtained A.sub.n value with that of the previously stored one.
In the case of this calibration being performed for the first time,
the A.sub.n value will be compared to that obtained from a
calibration routine performed immediately before implantation. That
A.sub.n value will be correlated with the A.sub.n,m values
(obtained from the in vitro testing) to provide a direct
relationship to permeability of the semipermeable membranes.
Subsequently, and with the help of the stored A.sub.n,m values, the
sensor sensitivity factors will be updated immediately after every
successful internal calibration routine.
[0095] FIG. 5 illustrates one embodiment of performing an internal
calibration routine to assess changes in the electrode's activity
as a function of operation. First, as noted above, it is desirable
to assess a time regime where glucose concentration levels are
fairly constant. When this is achieved, a first internal
calibration routine is conducted upon a user having a first
constant glucose level (level-I) to acquire a value for A.sub.nI.
Upon acceptance of this routine at the first constant glucose level
the user is instructed to eat a particular meal that raises his/her
glucose levels to a second constant glucose level (level-II). A
second calibration routine is then performed at the second constant
glucose level to acquire a value for A.sub.nII. Following
acceptance of the second calibration routine, the computer compares
the A.sub.nI with the A.sub.nII values. Since the two calibration
routines have been performed within a short duration, it is safe to
assume that the permeability of the sensor's semipermeable
membranes remains constant. Based on this, the difference between
A.sub.nI and A.sub.nII provides an indication of changes in
electrode activity. Such changes in electrode activity will be
assessed by in vitro calibration where the sensors have been
interrogated for extended periods of time. This calibration chart
will be stored into the operating program of the driving computer
processing unit and utilized to re-adjust the sensor sensitivity
immediately after completion of this calibration sequence.
[0096] In another embodiment, performing an internal calibration
routine to assess changes in the electrode's activity as a function
of operating time can originate by tracking the shape and slope of
the sensor response during each measurement, and in particular
immediately after the calibration state. This time-resolved shape
and slope of the response curve is intimately dependent upon the
electrode activity. In order to accomplish this it is desirable to
have fast electronics and storage capability to (i) record, (ii)
store, and (iii) compare the time-resolved decay at each glucose
concentration. With the help of an advanced computer, this function
can be readily accomplished and compared with a calibration chart
(obtained as described above) to extract the electrode activity and
recalibrate the sensor. In yet another embodiment, a calibration
routine can be performed by using a single t.sub.bias measurement;
t.sub.bias the time that the sensor is subjected to the biasing
voltage. This single measurement is recorded, stored and compared
against the P.sub.16 results at the final and mid point of the
curve.
[0097] In yet another embodiment, an internal calibration routine
to assess changes in the electrode's activity as a function of
operation can originate by comparing two sequential calibration
steps shown in FIG. 5 performed on the same glucose concentration.
The difference between the two calibration steps arise by varying
the t.sup.bias of the two P.sub.16 steps. For example, the second
P.sub.16 step can be half or quarter of the first P.sub.16 step. By
recording the corresponding end points of the two S.sub.16 curves,
their difference corresponds to the slope of the S.sub.16 curve.
This slope is dependent on the activity of the electrode and can be
readily assessed by an in vitro calibration study routine obtained
by sensors that have been subjected for extended periods of time
and interrogated with two t.sub.bias at various glucose
concentrations. This calibration chart will also be stored in the
operating program of the driving computer processing unit and will
be used to compare and re-adjust sensor sensitivity with respect to
electrode activity.
[0098] The periodic biased chronoamperometry is performed by
biasing the working electrode for about 1 microsecond to about 10
seconds, specifically about 10 microseconds to about 5 seconds and
more specifically about 100 microseconds to about 1 second. In one
embodiment, the biasing of the working electrode is conducted at
regular periods of greater than or equal to about 5 minutes.
[0099] To further enhance the reliability of the sensor and to also
account for interferences from exogenous species, the sensor can be
interrogated at various potentials. FIG. 6 illustrates a typical
cyclic voltammogram of an electrochemical sensor containing a
composite of sensing elements (i.e. ascorbic acid at #3,
acetaminophen, uric and ascorbic acid at #2, glucose and all of
previous three at #1, O.sub.2 at # 4, O.sub.2 and H.sub.2O.sub.2 at
#5, and so on).
[0100] Based on this sequential interrogation of the sensor at
various biasing potentials (depicted by the broken lines in FIG.
6), it can be seen that since different analytes contribute to the
amperometric signal to different extents, this provides a means to
accessing the individual concentrations of various analytes. It is
to be noted that in order to perform amperometry at different
potentials, sensor equilibration is desirable at each potential, as
shown for the one potential in FIG. 2.
[0101] The multi-component complexity of such measurements can be
substantially simplified if the glucose concentration remains
fairly constant. Using this information and following similar
procedures to that described above in reference to the periodic
biasing for sensor equilibration and measurement of FIG. 3, one has
to repeat this for each interrogating potential. This is
schematically shown in FIG. 7. A matrix formulation routine can be
utilized to deduce and solve n.sup.th order parametric equations
involving the dependence of the response of the glucose sensor on
these interfering species. The response of the sensor at a
particular interrogating potential (following sensor equilibration)
is dependent on t.sub.bias, t.sub.wait, and the permeability of the
sensor's semi permeable membrane. This being the case, the response
of the sensor at a particular interrogating potential can be
written as follows:
i = 1 i = n C 1 i x i + j = 1 j = n C 1 j y j + k = 1 k = n C 1 k z
k + + i = 1 i = n j = 1 j = n C 1 ij x i y j + i = 1 i = n k = 1 k
= n C 1 ik x i z k + j = 1 j = n k = 1 k = n C 1 jk y j z k + i = 1
i = n j = 1 j = n k = 1 k = n C 1 ijk x i y j z k + + C 1 = 0 i = 1
i = n C 2 i x i + j = 1 j = n C 2 j y j + k = 1 k = n C 2 k z k + +
i = 1 i = n j = 1 j = n C 2 ij x i y j + i = 1 i = n k = 1 k = n C
2 ik x i z k + j = 1 j = n k = 1 k = n C 2 jk y j z k + i = 1 i = n
j = 1 j = n k = 1 k = n C 2 ijk x i y j z k + + C 2 = 0 i = 1 i = n
C mi x i + j = 1 j = n C m y j + k = 1 k = n C mk z k + + i = 1 i =
n j = 1 j = n C mij x i y j + i = 1 i = n k = 1 k = n C mik x i z k
+ j = 1 j = n k = 1 k = n C mjk y j z k + i = 1 i = n j = 1 j = n k
= 1 k = n C mijk x i y j z k + + C m = 0 ##EQU00002##
where m is the number of interrogating potentials, x, y, z, . . .
are the respective analyte concentrations, i, j, k are the power
law dependence (or any other mathematical functions) of the analyte
concentrations, and C.sub.m are constants for a particular analyte
or analyte overlapping sets and is given as follows:
C m = l ( t bias t wait , permeability of various analytes through
the sensor ` s semi permeable membranes ) ##EQU00003##
[0102] The x, y, z analyte concentrations are related to the
equilibrated amperometric responses at a given interrogating
potential through a function that involves signal contribution from
interfering analytes and diffusion related processes. This
diffusion related processes can be minimized by varying the
t.sub.bias and/or t.sub.wait to attain steady state. The t.sub.bias
and/or t.sub.wait at a given analyte concentration and at a given
interrogating potential, can provide secondary measuring data for
increasing reliability of the sensor. Such secondary measuring data
can also provide the ability to estimate and accurately correct the
natural decay of the sensor as a result of sensor drifts from
changes in permeability of outer membranes and
electrode-activity.
[0103] In an alternative methodology, the multi-potential
interrogation can be performed together with the calibration
sequences of FIGS. 4 and 5 to further increase the confidence level
of assessing interferences while the sensor is operational in
vivo.
[0104] In summary, the electro-active changes of the working
electrode can be determined by comparing the differential values
from two different internal calibration routines at two different
analyte concentrations. In another embodiment, the electro-active
changes of the said working electrode can be determined by
comparing the time dependent decay of the amperometric signal
immediately before and immediately after the time interval. In
addition, the time-dependent decay involves determining the slope
of the time-dependent decay curve. The slope of the time-dependent
curves involves interrogating the sensor at two different
t.sub.bias times. The slope of the time-dependent decay curve is
obtained by comparing the signal value at the final and mid point
of the curve.
[0105] The disclosed sensor along with the testing methodology
disclosed herein has a number of advantages. The implantable
glucose sensors can use periodically biased amperometry for
interrogation to improve the sensor's sensitivity and linearity
while at the same time enabling internal calibration against sensor
drifts that originate from changes in either electrode activity or
membrane permeability as a result of fouling, calcification and/or
fibrosis.
[0106] The aforementioned features provide numerous advantages over
other comparative biosensors in that they exhibit high linearity
and sensitivity. They take into account the contribution of
exogenous interfering species and provide internal calibration
routines to control and reduce sensor-induced drift based on in
vivo induced effects that change the permeability of semi-permeable
membrane. These features also account for the gradual decay in
electrode activity.
[0107] In addition, the implantable glucose sensor uses a hydrogel
layer that comprises PVA alone. The use of PVA ensures a
homogeneous coverage of the immobilized glucoseoxidase (GO.sub.x)
enzyme layer 110. It also facilitates the storage of O.sub.2 and
permits control of the amount of stored O.sub.2 by varying the
number of freeze-thaw cycles for the PVA.
[0108] The following examples, which are meant to be non-limiting
were conducted to demonstrate the method of manufacturing the
implantable glucose sensor disclosed herein. These examples also
demonstrate some of the methods of interrogation and calibration of
the implantable glucose sensor disclosed herein.
EXAMPLES
Example 1
[0109] This example was conducted to demonstrate the effect of the
application of a biasing voltage to the surface of the working
electrode of the glucose sensor. In order to increase the signal to
noise ratio for a spectroscopic technique to successfully
interrogate the effect of bias and presence of electro-active
species, a thin layer of poly (ortho-phenylene diamine) (PPD)
(which acts as the electrically conducting membrane 106) was
electropolymerized on top of an indium tin oxide (ITO) coated
substrate, a transparent conductor which can also be a material for
working electrode.
[0110] The approximate thickness of the electropolymerized PPD film
is ca. 10 nanometers (nm) and is comparable to that on an actual
glucose sensor. FIG. 8 illustrates five overlapping UV-Vis-NIR
absorption curves of the PPD films taken at different biasing
voltages (0 Volts (V), 0.2V, 0.4V, 0.6V and 0.8 V) with respect to
a Ag/AgCl reference electrode. The application of the biasing
voltage increases the intensity of the 450 nm absorption peak while
decreasing the broad NIR absorption, with an isosbestic point ca.
690 nm. This indicates that PPD changes its conductivity with the
application of higher bias.
[0111] FIG. 9 illustrates the UV-Vis-NIR absorption spectra for the
same PPD film that was first biased for 150 sec at 0.2 V followed
by removal of bias and exposure to 1 millimolar (mM) of
H.sub.2O.sub.2. The spectra in the FIG. 9 shows that the
application of a 0.2 V bias leads to a blue-shift in the 400 nm
absorption range and decrease in broad NIR absorption corresponding
to the changes in conductivity of PPD as shown in FIG. 8. Upon
exposure to 1 mM H.sub.2O.sub.2, the 400 nm absorption range was
red-shifted, thus opposing the action of the applied bias. This
demonstrates that the generated H.sub.2O.sub.2 interferes with the
oxidation state of PPD and therefore causes changes to the
electrochemical activity of the working electrode.
[0112] These spectroscopic results are in agreement with the
amperometric sensor behavior shown in FIG. 2, suggesting that the
rapid sensitivity decrease might originate from over-oxidation of
PPD that renders it less conductive and therefore less prone to
receive electrons from the reaction 2 detailed above. Upon the
removal of the positive bias, the H.sub.2O.sub.2 whose levels
depend on the permeability of the semi permeable membrane(s),
brings the electrode (i.e. PPD) to its conductive state.
Example 2
[0113] This example was conducted to demonstrate the different
reactivities of the working electrode depending on whether the
device is operated in a continuous or periodic biasing mode. The
working electrode for this example is denoted by the
nomenclature--Pt/PPD/GO.sub.x/(HAs--Fe.sup.3+).sub.5, where the
electrode comprises platinum, the electrically conducting membrane
comprises PPD, the enzyme layer comprises GO.sub.x and the
semipermeable membrane comprises of a humic acid
--Fe.sup.3+(HAs--Fe.sup.3+).sub.5 layer .
[0114] FIG. 10 demonstrates the different reactivities of the
working electrode when it is operated in continuous vs. periodic
biasing at a t.sub.bias=1 second. The biasing voltage was 0.7 V and
the reference electrode comprises Ag/AgCl. For the periodic biasing
mode, the operation took place in regime A (as witnessed in the
FIG. 2), which explains its higher recorded amperometric current
versus that for the continuous biasing mode that operates in regime
B. As can be seen from the FIG. 10, one of the advantages of
periodic biasing is that greater sensor linearity can be achieved
because the pristine state and consequently pristine activity of
the electrode is retained by virtue of subjecting it to periodic
biasing.
Example 3
[0115] This example was conducted to demonstrate the conditions for
equilibrium performance of the working electrode. The working
electrode used in this example has the same configuration as that
described in the Example 2. FIG. 11 illustrates the periodic
biasing operation of the device at a constant glucose concentration
of 2 mM when operated at a biasing voltage of 0.7 V. The reference
electrode comprised Ag/AgCl. The biasing time t.sub.bias was of 1
second duration and the waiting time t.sub.wait was varied from 3.5
to 5 to 7.5 minutes.
[0116] This device operates in regime A (see FIG. 2) where sensor
equilibration has not been attained. The sensor has been removed
from the test cell, washed with deionized (DI) water and
reconnected to the electrochemical potentiostat prior to commencing
the next set of t.sub.wait experiments. Amperometric response
stabilization is witnessed at 7.5 min. This indicates that the
departure from equilibrium by the application of a biasing time
(t.sub.bias) of 1 second requires approximately 7.5 minutes of
incubation in a H.sub.2O.sub.2 environment generated by 2 mM of
glucose.
Example 4
[0117] This example was conducted to demonstrate the advantages
provided by the use of a hydrogel layer that comprises only PVA.
This example compares the response of two electrodes--one that
contains PVA while the other does not contain the PVA.
[0118] As noted in the FIG. 1, a first layer of a first hydrogel
118 is disposed between the semi-permeable membrane 114 and a
second layer of the second hydrogel 122. In this example, the
semi-permeable membrane 114 comprises an LBL-grown membrane while
the second layer of the second hydrogel 122 comprises
TRM-containing microspheres.
[0119] In general, water-containing hydrogels act as poor absorbers
for oxygen as opposed to hydrophobic polymers. It is therefore
desirable to incorporate hydrophobic domains within the hydrogels
to increase their oxygen storing ability.
[0120] FIG. 12 shows the amperometric response for a continuous
biasing voltage of 0.7 V when used on two working electrodes. Both
electrodes are similar to that used in the Example 3, except that
one has no PVA, while the other has a layer of PVA hydrogel. Thus
in terms of the nomenclature adopted in the Example 2, one
electrode comprises Pt/PPD/GO.sub.x/(HAs--Fe.sup.3-).sub.5 (with no
PVA) while the other comprises
Pt/PPD/GO.sub.x/(HAs--Fe.sup.3+).sub.5/PVA (with PVA). The
amperometric response is measured versus a Ag/AgCl reference
electrode for varying glucose concentrations. As can be seen in the
FIG. 12, the incorporation of PVA dramatically increases linearity
and sensor sensitivity when compared to the device that does not
contain PVA.
[0121] In order to investigate the role of PVA in enhancing
sensor's linearity and sensitivity, the oxygen content in the PVA
layer was determined. These were determined by sealing a known
amount of a PVA sample in a glass tube under vacuum (hereinafter
"tube containing PVA" sample). This was achieved by filling the
actual tube with a 10% weight per unit volume of aqueous PVA
solution and performing a number of freeze-thaw cycles varying from
1 to 7 prior to flame sealing the tube under vacuum, while the gel
is frozen. This hydrogel containing glass capsule was then taken
into an air tight chamber containing 50 ml amount of DI water and a
commercial oxygen sensor. Following this, a test cell was sealed
from the atmosphere and N.sub.2 was bubbled in order to decrease
the oxygen concentration in the test cell. When the oxygen
concentration in the test cell reached 0 .mu.M, the tube containing
PVA was crushed, to let its oxygen level equilibrate with the
surrounding media. The increase in the oxygen level was recorded as
function of the number of freeze-thaw cycles, and the experiment
was repeated in triplicate. As shown in FIG. 13, the oxygen content
of PVA increases with increasing freeze-thaw cycles, which
indicates that the nature of the linearity and sensitivity increase
in FIG. 12, originates from the ability of PVA to store oxygen.
This relationship between the number of freeze thaw cycles and the
amount of stored oxygen was hitherto unknown.
Example 5
[0122] This example was conducted to demonstrate the effect of
incorporating carbon nanotubes in the PPD layer (see layer 106 of
FIG. 1). To enhance the ability of the present glucose sensor to
also detect oxygen, the PPD layer was formed (through
electropolymerization) in the presence of acid treated single
walled carbon nanotubes (SWNTs). The incorporation of SWNTs in PPD
is believed to originate from charge balancing the positively
charged PPI).
[0123] FIG. 14, illustrates the cyclic voltammogram of the
Pt/PPD+SWNT device in a 0.1 M Phosphate Buffer Saline (PBS) buffer
solution (pH=7.2) with and without the presence of O.sub.2. The
reference electrode is Ag/AgCl. In the presence of the SWNTs, a
shoulder at -0.1 V can be seen. This shoulder corresponds to the
reduction of oxygen, which is electro-catalyzed by the combination
of PPD and SWNTs. The corresponding continuously biased
amperometric response of the same device operated at -0.1 V vs. the
Ag/AgCl reference is shown in FIG. 15. As can be seen, the device
responds quickly and reproducibly to changes in oxygen
concentration.
Example 6
[0124] This example demonstrates a strategy to enable
simultaneously the "direct wiring" of the redox enzyme to the
working electrode (as in third generation biosensors) via SWNT
networks while at the same time utilizing the enhanced
electrocatalytic activity of PPD/SWNT composites to enable oxygen
sensing. As shown in FIG. 16, the working electrode is decorated
with SWNT networks via sequential immersion first in NAFION second
in an aqueous solution of FeCl.sub.3 and third in acid treated
SWNTs that have been washed and dispersed in N,N-dimethylformamide
(DMF). These carboxy-functionalized tips of the SWNT network can be
covalently reacted with either a variety of redox enzymes (i.e.,
GO.sub.x) or with its Flavin Adenine Dinucleotide (FAD) cofactor
followed by reconstitution with the apo-enzyme (i.e.,
apo-GO.sub.x). Subsequently the PPD can be grown within this SWNT
network via electropolymerization of OPD to PPD as shown in FIG.
16.
[0125] As can be seen in the aforementioned examples, the use of a
PVA hydrogel layer facilitates the storage of O.sub.2 and permits
control of the amount of stored O.sub.2 by varying the number of
freeze-thaw cycles for the PVA. The use of the PVA hydrogel layer
in the sensor when used in conjunction with intermittent biasing
permits the sensor to be self-calibrating which reduces maintenance
costs and replacements costs.
[0126] While the invention has been described in detail in
connection with a number of embodiments, the invention is not
limited to such disclosed embodiments. Rather, the invention can be
modified to incorporate any number of variations, alterations,
substitutions or equivalent arrangements not heretofore described,
but which are commensurate with the scope of the invention.
Additionally, while various embodiments of the invention have been
described, it is to be understood that aspects of the invention may
include only some of the described embodiments. Accordingly, the
invention is not to be seen as limited by the foregoing
description, but is only limited by the scope of the appended
claims.
* * * * *