U.S. patent application number 14/144497 was filed with the patent office on 2014-05-08 for device and method for determining analyte levels.
This patent application is currently assigned to DexCom, Inc.. The applicant listed for this patent is DexCom, Inc.. Invention is credited to Rathbun K. Rhodes, Mark C. Shults, Stuart J. Updike.
Application Number | 20140128700 14/144497 |
Document ID | / |
Family ID | 25206642 |
Filed Date | 2014-05-08 |
United States Patent
Application |
20140128700 |
Kind Code |
A1 |
Shults; Mark C. ; et
al. |
May 8, 2014 |
DEVICE AND METHOD FOR DETERMINING ANALYTE LEVELS
Abstract
Devices and methods for determining analyte levels are
described. The devices and methods allow for the implantation of
analyte-monitoring devices, such as glucose monitoring devices,
that result in the delivery of a dependable flow of blood to
deliver sample to the implanted device. The devices comprise a
unique microarchitectural arrangement in the sensor region that
allows accurate data to be obtained over long periods of time.
Inventors: |
Shults; Mark C.; (Madison,
WI) ; Updike; Stuart J.; (Madison, WI) ;
Rhodes; Rathbun K.; (Madison, WI) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
DexCom, Inc. |
San Diego |
CA |
US |
|
|
Assignee: |
DexCom, Inc.
San Diego
CA
|
Family ID: |
25206642 |
Appl. No.: |
14/144497 |
Filed: |
December 30, 2013 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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13166685 |
Jun 22, 2011 |
8676288 |
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14144497 |
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09447227 |
Nov 22, 1999 |
8527025 |
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13166685 |
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08811473 |
Mar 4, 1997 |
6001067 |
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09447227 |
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Current U.S.
Class: |
600/347 |
Current CPC
Class: |
A61B 5/1495 20130101;
A61B 5/14532 20130101; C12Q 1/006 20130101; A61B 5/0004 20130101;
A61B 5/14865 20130101 |
Class at
Publication: |
600/347 |
International
Class: |
A61B 5/1486 20060101
A61B005/1486; A61B 5/00 20060101 A61B005/00; A61B 5/1495 20060101
A61B005/1495; A61B 5/145 20060101 A61B005/145 |
Claims
1-20. (canceled)
21. A system for processing sensor data from a continuous glucose
sensor, the system comprising: a sensor comprising an electrode
configured for implantation in a subcutaneous tissue of a host,
wherein the sensor comprises a membrane configured to reduce a flux
of glucose therethrough, wherein the membrane comprises a product
of a reaction of an isocyanate compound with a compound comprising
a hydroxyl group and/or a compound comprising an amine group,
wherein the sensor is configured to continuously measure a glucose
concentration whereby sensor data indicative of the subcutaneous
glucose concentration is generated; and electronic circuitry
operably connected to the sensor; wherein the system is configured
to provide a substantially accurate measurement of the glucose
concentration in the host such that at least 95% of calibrated
sensor glucose values as determined by analysis of blood obtained
during a useful life of the sensor are within 25% of actual glucose
values.
22. The system of claim 21, wherein the product is
polyurethane.
23. The system of claim 21, wherein the product is polyurea.
24. The system of claim 21, wherein the product is a
polyurethaneurea copolymer.
25. The system of claim 21, wherein the useful life is greater than
about 3 days.
26. The system of claim 21, wherein the useful life is greater than
about 5 days.
27. The system of claim 21, wherein the useful life is greater than
about 7 days.
Description
INCORPORATION BY REFERENCE TO RELATED APPLICATIONS
[0001] Any and all priority claims identified in the Application
Data Sheet, or any correction thereto, are hereby incorporated by
reference under 37 CFR 1.57. This application is a continuation of
U.S. application Ser. No. 13/166,685, filed Jun. 22, 2011, which is
a continuation of U.S. application Ser. No. 09/447,227, filed Nov.
22, 1999, now U.S. Pat. No. 8,527,025, which is a division of U.S.
application Ser. No. 08/811,473, filed Mar. 4, 1997, now U.S. Pat.
No. 6,001,067. Each of the aforementioned applications is
incorporated by reference herein in its entirety, and each is
hereby expressly made a part of this specification.
FIELD OF THE INVENTION
[0002] The present invention relates generally to devices and
methods for determining analyte levels, and, more particularly, to
implantable devices and methods for monitoring glucose levels in a
biological fluid.
BACKGROUND OF THE INVENTION
[0003] The continuous measurement of substances in biological
fluids is of interest in the control and study of metabolic
disorders. Electrode systems have been developed for this purpose
whereby an enzyme-catalyzed reaction is monitored (e.g., by the
changing concentrations of reactants or products) by an
electrochemical sensor. In such electrode systems, the
electrochemical sensor comprises an electrode with potentiometric
or amperometric function in close contact with a thin layer
containing an enzyme in dissolved or insoluble form. Generally, a
semipermeable membrane separates the thin layer of the electrode
containing the enzyme from the sample of biological fluid that
includes the substance to be measured.
[0004] Electrode systems that include enzymes have been used to
convert amperometrically inactive substances into reaction products
which are amperometrically active. For example, in the analysis of
blood for glucose content, glucose (which is relatively inactive
amperometrically) may be catalytically converted by the enzyme
glucose oxidase in the presence of oxygen and water to gluconic
acid and hydrogen peroxide. Tracking the concentration of glucose
is possible since for every glucose molecule converted a
proportional change in either oxygen or hydrogen peroxide sensor
current will occur [U.S. Pat. Nos. 4,757,022 and 4,994,167 to
Shults et al., both of which are hereby incorporated by reference].
Hydrogen peroxide is anodically active and produces a current which
is proportional to the concentration of hydrogen peroxide, which is
directly related to the concentration of glucose in the sample.
[Updike et al., Diabetes Care, 11:801-807 (1988)].
[0005] Despite recent advances in the field of implantable glucose
monitoring devices, presently used devices are unable to provide
data safely and reliably for long periods of time (e.g., months or
years) [See, e.g., Moatti-Sirat et al., Diabetologia 35:224-30
(1992)]. For example, Armour et al., Diabetes 39:1519-26 (1990),
describes a miniaturized sensor that is placed intravascularly,
thereby allowing the tip of the sensor to be in continuous contact
with the blood. Unfortunately, probes that are placed directly into
the vasculature put the recipient at risk for thrombophlebosis,
thromboembolism, and thrombophlebitis.
[0006] Currently available glucose monitoring devices that may be
implanted in tissue (e.g., subcutaneously) are also associated with
several shortcomings. For example, there is no dependable flow of
blood to deliver sample to the tip of the probe of the implanted
device. Similarly, in order to be effective, the probe must consume
some oxygen and glucose, but not enough to perturb the available
glucose which it is intended to measure; subcutaneously implanted
probes often reside in a relatively stagnant environment in which
oxygen or glucose depletion zones around the probe tip may result
in erroneously low measured glucose levels. Finally, the probe may
be subject to "motion artifact" because the device is not
adequately secured to the tissue, thus contributing to unreliable
results. Partly because of these limitations, it has previously
been difficult to obtain accurate information regarding the changes
in the amounts of analytes (e.g., whether blood glucose levels are
increasing or decreasing); this information is often extremely
important, for example, in ascertaining whether immediate
corrective action is needed in the treatment of diabetic
patients.
[0007] There is a need for a device that accurately and
continuously determines the presence and the amounts of a
particular analyte, such as glucose, in biological fluids. The
device should be easy to use, be capable of accurate measurement of
the analyte over long periods of time, and should not readily be
susceptible to motion artifact.
SUMMARY OF THE INVENTION
[0008] The present invention relates generally to devices and
methods for determining analyte levels, and, more particularly, to
implantable devices and methods for monitoring glucose levels in a
biological fluid.
[0009] The devices and methods of the present invention allow for
the implantation of analyte-monitoring devices such as glucose
monitoring devices that result in a dependable flow of blood to
deliver sample to the implanted device at a concentration
representative of that in the vasculature. Moreover, the devices of
the present invention become secured within the tissue of the
subject, thereby greatly reducing or eliminating the phenomenon of
"motion artifact". In addition, the devices of the present
invention utilize materials that eliminate or significantly delay
environmental stress cracking at the sensor interface, resulting in
the ability to obtain accurate, long-term data.
[0010] These effects result, in part, from the use of materials
that enhance the formation of a foreign body capsule (FBC).
Previously, FBC formation has been viewed as being adverse to
sensor function, and researchers have attempted to minimize FBC
formation (see, e.g., U.S. Pat. No. 5,380,536 to Hubbell et al.).
However, the methods and devices of the present invention utilize
specific materials and microarchitecture that elicit a type of FBC
that does not hamper the generation of reliable data for long
periods. The devices of the present invention are capable of
accurate operation in the approximately 37.degree. C., low
PO.sub.2, environment characteristic of living tissue for extended
lengths of time (e.g., months to years).
[0011] The electrode-membrane region of the devices of the present
invention comprises a unique microarchitectural arrangement. In
preferred embodiments, the electrode surfaces are in contact with
(or operably connected with) a thin electrolyte phase, which in
turn is covered by an enzyme membrane that contains an enzyme,
e.g., glucose oxidase, and a polymer system. A bioprotective
membrane covers this enzyme membrane system and serves, in part, to
protect the sensor from external forces and factors that may result
in environmental stress cracking. Finally, an angiogenic layer is
placed over the bioprotective membrane and serves to promote
vascularization in the sensor interface region. It is to be
understood that other configurations (e.g., variations of that
described above) are contemplated by the present invention and are
within the scope thereof.
[0012] The present invention contemplates a biological fluid
measuring device, comprising a) a housing comprising electronic
circuit means and at least two electrodes operably connected to the
electronic circuit means; and b) a sensor means operably connected
to the electrodes of the housing, the sensor means comprising i) a
bioprotective membrane, and ii) an angiogenic layer, the angiogenic
layer positioned more distal to the housing than the bioprotective
membrane. In particular embodiments, the bioprotective membrane is
substantially impermeable to macrophages. In some embodiments, the
bioprotective membrane comprises pores having diameters ranging
from about 0.1 micron to about 1.0 micron. In certain embodiments,
the bioprotective membrane comprises polytetrafluoroethylene, and
in particular embodiments, the angiogenic layer also comprises
polytetrafluoroethylene.
[0013] Particular embodiments of the biological fluid measuring
device further comprise c) means for securing the device to
biological tissue, the securing means associated with the housing.
In some embodiments, the securing means comprises a polyester
velour jacket. In preferred embodiments, the securing means covers
the top surface (e.g., the top member or the top member sheath, as
described further below) and a portion of the sensor interface; it
should be noted that the securing means generally should not cover
the entire sensor interface, as this would interfere with the
ability of blood vessels to deliver sample to the biological fluid
measuring device. In preferred embodiments, the securing means
comprises poly(ethylene terephthalate).
[0014] In further embodiments, the sensor means of the biological
fluid measuring device further comprises means for determining the
amount of glucose in a biological sample. In some embodiments, the
glucose determining means comprises a membrane containing glucose
oxidase, the glucose oxidase-containing membrane positioned more
proximal to the housing than the bioprotective membrane. In
additional embodiments, the housing further comprises means for
transmitting data to a location external to the device (e.g., a
radiotelemetry device).
[0015] The present invention also contemplates a device for
measuring glucose in a biological fluid, comprising a) a housing
comprising electronic circuit means and at least one electrode
operably connected to the electronic circuit means; and b) a sensor
means operably connected to the electrode of the housing, the
sensor means comprising i) means for determining the amount of
glucose in a biological sample, the glucose determining means
operably associated with the electrode, ii) a bioprotective
membrane, the bioprotective membrane positioned more distal to the
housing than the glucose determining means and substantially
impermeable to macrophages, and iii) an angiogenic layer, the
angiogenic layer positioned more distal to the housing than the
bioprotective membrane.
[0016] In particular embodiments, the glucose determining means
comprises a membrane containing glucose oxidase. In some
embodiments, the angiogenic layer comprises
polytetrafluoroethylene.
[0017] In some embodiments, the pores of the bioprotective membrane
have diameters ranging from about 0.1 micron to about 1.0 micron,
while in other embodiments the pores have diameters ranging from
about 0.2 micron to about 0.5 micron. In certain embodiments, the
bioprotective membrane comprises polytetrafluoroethylene.
[0018] Still other embodiments further comprise c) means for
securing the device to biological tissue, the securing means
associated with the housing. In particular embodiments, the
securing means comprises poly(ethylene terephthalate). Additional
embodiments comprise means for transmitting data to a location
external to the device; in some embodiments, the data transmitting
means comprises a radiotelemetric device.
[0019] The present invention also contemplates a method for
monitoring glucose levels, comprising a) providing i) a host, and
ii) a device comprising a housing and means for determining the
amount of glucose in a biological fluid; and b) implanting the
device in the host under conditions such that the device measures
the glucose accurately for a period exceeding 90 days. In some
embodiments, the device measures glucose accurately for a period
exceeding 150 days, while in other embodiments, the device measures
glucose accurately for a period exceeding 360 days.
[0020] The present invention also contemplates a method of
measuring glucose in a biological fluid, comprising a) providing i)
a host, and ii) a device comprising a housing and means for
determining the amount of glucose in a biological fluid, the
glucose determining means capable of accurate continuous glucose
sensing; and b) implanting the device in the host under conditions
such that the continuous glucose sensing begins between
approximately day 2 and approximately day 25. In some embodiments,
the continuous glucose sensing begins between approximately day 3
and approximately day 21. In particular embodiments, the implanting
is subcutaneous.
[0021] The devices of the present invention allow continuous
information regarding, for example, glucose levels. Such continuous
information enables the determination of trends in glucose levels,
which can be extremely important in the management of diabetic
patients.
Definitions
[0022] In order to facilitate an understanding of the present
invention, a number of terms are defined below.
[0023] The term "accurately" means, for example, 95% of measured
values within 25% of the actual value as determined by analysis of
blood plasma, preferably within 15% of the actual value, and most
preferably within 5% of the actual value. It is understood that
like any analytical device, calibration, calibration check and
recalibration are required for the most accurate operation of the
device.
[0024] The term "analyte" refers to a substance or chemical
constituent in a biological fluid (e.g., blood or urine) that can
be analyzed. A preferred analyte for measurement by the devices and
methods of the present invention is glucose.
[0025] The terms "sensor interface," "sensor means," and the like
refer to the region of a monitoring device responsible for the
detection of a particular analyte. For example, in some embodiments
of a glucose monitoring device, the sensor interface refers to that
region wherein a biological sample (e.g., blood or interstitial
fluid) or a portion thereof contacts (directly or after passage
through one or more membranes or layers) an enzyme (e.g., glucose
oxidase); the reaction of the biological sample (or portion
thereof) results in the formation of reaction products that allow a
determination of the glucose level in the biological sample. In
preferred embodiments of the present invention, the sensor means
comprises an angiogenic layer, a bioprotective layer, an enzyme
layer, and an electrolyte phase (i.e., a free-flowing liquid phase
comprising an electrolyte-containing fluid [described further
below]). In some preferred embodiments, the sensor interface
protrudes beyond the plane of the housing.
[0026] The terms "operably connected," "operably linked," and the
like refer to one or more components being linked to another
component(s) in a manner that allows transmission of, e.g., signals
between the components. For example, one or more electrodes may be
used to detect the amount of analyte in a sample and convert that
information into a signal; the signal may then be transmitted to
electronic circuit means (i.e., the electrode is "operably linked"
to the electronic circuit means), which may convert the signal into
a numerical value in the form of known standard values.
[0027] The term "electronic circuit means" refers to the electronic
circuitry components of a biological fluid measuring device
required to process information obtained by a sensor means
regarding a particular analyte in a biological fluid, thereby
providing data regarding the amount of that analyte in the fluid.
U.S. Pat. No. 4,757,022 to Shults et al., previously incorporated
by reference, describes suitable electronic circuit means (see,
e.g., FIG. 7); of course, the present invention is not limited to
use with the electronic circuit means described therein. A variety
of circuits are contemplated, including but not limited to those
circuits described in U.S. Pat. Nos. 5,497,772 and 4,787,398,
hereby incorporated by reference.
[0028] The terms "angiogenic layer," "angiogenic membrane," and the
like refer to a region, membrane, etc. of a biological fluid
measuring device that promotes and maintains the development of
blood vessels microcirculation around the sensor region of the
device. As described in detail below, the angiogenic layer of the
devices of the present invention may be constructed of membrane
materials alone or in combination such as polytetrafluoroethylene,
hydrophilic polyvinylidene fluoride, mixed cellulose esters,
polyvinyl chloride, and other polymers including, but not limited
to, polypropylene, polysulphone, and polymethacrylate.
[0029] The phrase "positioned more distal" refers to the spatial
relationship between various elements in comparison to a particular
point of reference. For example, some embodiments of a biological
fluid measuring device comprise both a bioprotective membrane and
an angiogenic layer/membrane. If the housing of the biological
fluid measuring device is deemed to be the point of reference and
the angiogenic layer is positioned more distal to the housing than
the bioprotective layer, then the bioprotective layer is closer to
the housing than the angiogenic layer.
[0030] The terms "bioprotective membrane," "bioprotective layer,"
and the like refer to a semipermeable membrane comprised of
protective biomaterials of a few microns thickness or more which
are permeable to oxygen and glucose and are placed over the tip of
the sensor to keep the white blood cells (e.g., tissue macrophages)
from gaining proximity to and then damaging the enzyme membrane. In
some embodiments, the bioprotective membrane has pores (typically
from approximately 0.1 to approximately 1.0 micron). In preferred
embodiments, a bioprotective membrane comprises
polytetrafluoroethylene and contains pores of approximately 0.4
microns in diameter. Pore size is defined as the pore size provided
by the manufacturer or supplier.
[0031] The phrase "substantially impermeable to macrophages" means
that few, if any, macrophages are able to cross a barrier (e.g.,
the bioprotective membrane). In preferred embodiments, fewer than
1% of the macrophages that come in contact with the bioprotective
membrane are able to cross.
[0032] The phrase "means for securing said device to biological
tissue" refers to materials suitable for attaching the devices of
the present invention to, e.g., the fibrous tissue of a foreign
body capsule. Suitable materials include, but are not limited to,
poly(ethylene terephthalate). In preferred embodiments, the top of
the housing is covered with the materials in the form of surgical
grade fabrics; more preferred embodiments also contain material in
the sensor interface region (see FIG. 1B).
[0033] The phrase "means for determining the amount of glucose in a
biological sample" refers broadly to any mechanism (e.g., enzymatic
or non-enzymatic) by which glucose can be quantitated. For example,
some embodiments of the present invention utilize a membrane that
contains glucose oxidase that catalyzes the conversion of glucose
to gluconate:
Glucose+O.sub.2.fwdarw.Gluconate+H.sub.2O.sub.2. Because for each
glucose molecule converted to gluconate, there is a proportional
change in the co-reactant O.sub.2 and the product H.sub.2O.sub.2,
one can monitor the current change in either the co-reactant or the
product to determine glucose concentration.
[0034] The phrase "means for transmitting data to a location
external to said device" refers broadly to any mechanism by which
data collected by a biological fluid measuring device implanted
within a subject may be transferred to a location external to the
subject. In preferred embodiments of the present invention,
radiotelemetry is used to provide data regarding blood glucose
levels, trends, and the like. The terms "radiotelemetry,"
"radiotelemetric device," and the like refer to the transmission by
radio waves of the data recorded by the implanted device to an ex
vivo recording station (e.g., a computer), where the data is
recorded and, if desired, further processed (see, e.g., U.S. Pat.
Nos. 5,321,414 and 4,823,808, hereby incorporated by reference; PCT
Patent Publication WO 9422367).
[0035] The term "host" refers to both humans and animals.
[0036] The phrase "continuous glucose sensing" refers to the period
in which monitoring of plasma glucose concentration is continuously
carried out. More specifically, at the beginning of the period in
which continuous glucose sensing is effected, the background sensor
output noise disappears, and the sensor output stabilizes (e.g.,
over several days) to a long-term level reflecting adequate
microcirculatory delivery of glucose and oxygen to the tip of the
sensor (see FIG. 2). Though an understanding of this effect is not
required in order to practice the present invention, it is believed
to be due to adequately vascularized foreign body capsule tissue in
consistent contact with the sensor interface of the blood glucose
monitoring device. Failure of adequate vascularization or
consistent contact of tissue with sensor will result in failure of
continuous glucose sensing.
BRIEF DESCRIPTION OF THE DRAWINGS
[0037] FIG. 1A depicts a cross-sectional drawing of one embodiment
of an implantable analyte measuring device of the present
invention.
[0038] FIG. 1B depicts a cross-sectional exploded view of the
sensor interface dome of FIG. 1A.
[0039] FIG. 1C depicts a cross-sectional exploded view of the
electrode-membrane region of FIG. 1B detailing the sensor tip and
the functional membrane layers.
[0040] FIG. 2 graphically depicts glucose levels as a function of
the number of days post-implant.
[0041] FIG. 3 graphically depicts a correlation plot (days 21 to
62) of a glucose infusion study with one device of the present
invention.
[0042] FIG. 4 depicts a typical response to in vitro calibration to
glucose of a device of the present invention.
[0043] FIGS. 5A, 5B, and 5C graphically depict three in vivo sensor
response curves plotted in conjunction with the reference blood
glucose values for one device of the present invention at
post-implant times of 25, 88, and 109 days.
[0044] FIG. 6 graphically depicts sensor glucose versus reference
glucose for one device of the present invention using the single
set of calibration factors from day 88 of FIG. 5B.
DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENT
[0045] The present invention relates generally to devices and
methods for determining analyte levels, and, more particularly, to
implantable devices and methods for monitoring glucose levels in a
biological fluid. In a preferred embodiment, the device and methods
of the present invention are used to determine the level of glucose
in a subject, a particularly important measurement for individuals
having diabetes.
[0046] Although the description that follows is primarily directed
at glucose monitoring devices and methods for their use, the
devices and methods of the present invention are not limited to
glucose measurement. Rather, the devices and methods may be applied
to detect and quantitate other analytes present in biological
fluids (including, but not limited to, amino acids and lactate),
especially those analytes that are substrates for oxidase enzymes
[see, e.g., U.S. Pat. No. 4,703,756 to Gough et al., hereby
incorporated by reference]. Moreover, the devices and methods of
the present invention may be utilized to present components of
biological fluids to measurement methods which are not
enzyme-based, including, but not limited to, those based on surface
plasmon resonance, surface acoustic waves, optical absorbance in
the long wave infrared region, and optical rotation of polarized
light.
I. Nature of the Foreign Body Capsule
[0047] Probes that are implanted (e.g., subcutaneously) into tissue
will almost always elicit a foreign body capsule (FBC) as part of
the body's response to the introduction of a foreign material.
Though a precise understanding of the nature of a FBC is not
required in order to practice the present invention, generally
speaking, upon implantation of a glucose sensor, there is initially
an acute inflammatory reaction (which includes invasion of tissue
macrophages), followed by building of fibrotic tissue. A mature
capsule (i.e., the FBC) comprising primarily avascular fibrous
tissue forms around the device [Woodward, Diabetes Care, 5:278-281
(1982)]. Although fluid is frequently found within the capsular
space between the sensor and the capsule, levels of analytes (e.g.,
glucose and oxygen) within the fluid often do not mimic levels in
the body's vasculature, making accurate measurement difficult.
Example 4 below describes typically identifiable phases in FBC
formation as reflected by response of an implanted glucose
sensor.
[0048] In general, the formation of FBCs has precluded the
collection of reliable, continuous information because they isolate
the sensor of the implanted device from biological fluids, fully
equilibrated with at least the low molecular weight components
found in the circulation. Similarly, the composition of FBCs has
prevented stabilization of the implanted device, contributing to
motion artifact that renders unreliable results. Thus,
conventionally, it has been the practice of those skilled in the
art to attempt to minimize FBC formation by, for example, using a
short lived needle geometry or sensor coatings to minimize the
foreign body reaction.
[0049] In contrast to the prior art, the teachings of the present
invention recognize that FBC formation is the dominant event
surrounding long term implantation of any sensor and must be
orchestrated to support rather than hinder or block sensor
performance. For example, sensors often do not perform well until
the FBC has matured sufficiently to provide ingrowth of well
attached tissue bearing a rich supply of capillaries directly to
the surface of the sensor. This maturation process takes at least
several days and, when initiated according to the present
invention, is a function of biomaterial and host factors which
initiate and modulate angiogenesis, and promote and control
fibrocyte ingrowth. The present invention contemplates the use of
particular materials to promote angiogenesis adjacent to the sensor
interface (also referred to as the electrode-membrane region,
described below) and to anchor the device within the FBC.
II. The Implantable Glucose Monitoring Devices of the Present
Invention
[0050] The present invention contemplates the use of a unique
microarchitectural organization around the sensor interface of an
implantable device. Moreover, the present invention contemplates
the use of materials covering all or a portion of the device to
assist in the stabilization of the device following implantation.
However, it should be pointed out that the present invention does
not require a device comprising particular electronic components
(e.g., electrodes, circuitry, etc). Indeed, the teachings of the
present invention can be used with virtually any monitoring device
suitable for implantation (or subject to modification allowing
implantation); suitable devices include, but are not limited, to
those described in U.S. Pat. Nos. 4,703,756 and 4,994,167 to Shults
et al.; U.S. Pat. No. 4,703,756 to Gough et al., and U.S. Pat. No.
4,431,004 to Bessman et al.; the contents of each being hereby
incorporated by reference, and Bindra et al., Anal. Chem.
63:1692-96 (1991).
[0051] In the discussion that follows, an example of an implantable
device that includes the features of the present invention is first
described. Thereafter, the specific characteristics of, for
example, the sensor interface contemplated by the present invention
will be described in detail.
[0052] Generally speaking, the implantable devices contemplated for
use with the present invention are oval shaped; of course, devices
with other shapes may also be used with the present invention. The
sample device includes a housing having an upper portion and a
lower portion which together define a cavity. FIG. 1A depicts a
cross-sectional drawing of one embodiment of an implantable
measuring device. Referring to FIG. 1A, the device comprises a main
housing (also referred to as casing or packaging) consisting of a
bottom member 1 with upwardly angled projecting extensions along
its perimeter. The four downwardly projecting extensions of a
similarly-shaped top member 2 engage the upwardly projecting
extensions of the bottom member 1. As indicated in FIG. 1A, there
is an aperture in top member 2 that allows for protrusion of the
sensor interface dome 30. Preferred embodiments of the present
invention entail such a protrusion of the sensor interface dome 30;
in some embodiments, though a precise understanding of the effect
of the protrusion is not required in order to practice the present
invention, the protrusion is believed to assist in the formation of
vasculature in the sensor interface dome 30 region, and hence
presentation of sample to the electrodes.
[0053] In certain embodiments, a top member sheath 4 covers the top
member 2; like the top member 2, the top member sheath 4 has an
aperture which allows the sensor interface dome 30 to protrude
therethrough. As indicated in detail in FIG. 1B, the top member
sheath 4 angles upward as it approaches the aperture, allowing the
sensor interface capsular attachment layer 15 to be secured
thereto. The top member sheath 4 may be coated with a sheath
capsular attachment layer 16; in some embodiments, the sheath
capsular attachment layer extends beyond the top member sheath
(e.g., it may jacket the sides of the device or the bottom
member).
[0054] Maintaining the blood supply near an implanted foreign body
like an implanted analyte-monitoring sensor requires stable
fixation of FBC tissue on the surface of the foreign body. This can
be achieved, for example, by using capsular attachment membrane
materials (e.g., those materials that comprise the sensor interface
and top member capsular attachment layers) developed to repair or
reinforce tissues, including, but not limited to, polyester
(DACRON.RTM.; DuPont; poly(ethylene terephthalate)) velour,
expanded polytetrafluoroethylene (TEFLON.RTM.; Gore),
polytetrafluoroethylene felts, polypropylene cloth, and related
porous implant materials. The preferred material for FBC attachment
is surgical-grade polyester velour. FBC tissue tends to
aggressively grow into the materials disclosed above and form a
strong mechanical bond (i.e., capsular attachment); this fixation
of the implant in its capsule is essential to prevent motion
artifact or disturbance of the newly-developed capillary blood
supply. In preferred embodiments, capsular attachment materials are
not used in the region of the sensor interface so as not to
interfere with the vasculature development in that region.
[0055] Side braces 3 secure the top member sheath 4 to the bottom
member 1 (see FIG. 1A). A conventional O-ring 7 or other suitable
mechanical means may be used to assist in the attachment of the
membrane layers (e.g., the enzyme layer). In a preferred
embodiment, the housing is approximately 1.4 cm from the base of
the bottom member 1 to the top of the sheath capsular attachment
layer 16, and approximately 7.0 cm in length.
[0056] The interior (i.e., the cavity) of the housing comprises one
or more batteries 9 operably connected to an electronic circuit
means (e.g., a circuit board 8), which, in turn, is operably
connected to at least one electrode (described below); in preferred
embodiments, at least two electrodes are carried by the housing.
Any electronic circuitry and batteries that renders reliable,
continuous, long-term (e.g., months to years) results may be used
in conjunction with the devices of the present invention.
[0057] The housing of the devices of the present invention
preferably utilize a simple, low-cost packaging technique which
protects the components of the device for at least one year in
aqueous media. In preferred embodiments, the components of the
housing (e.g., the top and bottom members) comprise thermoformed
high-density polyethylene. The area in the cavity of the housing
that surrounds the batteries, electronic circuitry, etc., may be
filled with an encapsulant 40 (see FIG. 1A), a material that
secures in place the components within the cavity but that does not
interfere with the operation of those components. In preferred
embodiments, the encapsulant 40 is based on mixtures of petroleum
wax and low melting temperature resins developed for the hot-melt
glue industry [Shults et al., IEEE Trans. Biomed. Eng. 41:937-942
(1994)]. In addition to the high-quality moisture barrier formed
with this approach, the electronics (e.g., the circuit board 8) can
be recycled by remelting and draining the encapsulant when the
battery expires.
[0058] The preferred encapsulant compositions of the present
invention comprise approximately 54% PW 130/35H wax (Astor Wax),
approximately 40% MVO 2528 resin (Exxon Chemical), and
approximately 6% XS 93.04 resin (Exxon Chemical, Houston, Tex.).
These pelletized compounds render a well-mixed solution after
heating and mixing at about 120.degree. C. for approximately one
hour. This solution is then poured into the polyethylene housing
containing the implant electronics, taking caution to not to exceed
the burst temperature of, e.g., approximately 160.degree. C. when
lithium batteries are used.
[0059] FIG. 1B depicts a cross-sectional exploded view of the
sensor interface dome 30 of FIG. 1A. Referring to FIG. 1B, the
sensor interface dome comprises a region of, for example, epoxy
insulation 10 in which is embedded a silver reference electrode 20,
a platinum working electrode 21, and a platinum counter electrode
22. The present invention is neither limited by the composition of
the electrodes nor their position within the sensor interface dome
30.
[0060] FIG. 1C depicts a cross-sectional exploded view of the
electrode-membrane region set forth in FIG. 1B detailing the sensor
tip and the functional membrane layers. As depicted in FIG. 1C, the
electrode-membrane region comprises several different membrane
layers, the compositions and functions of which are described in
detail below. The top ends of the electrodes are in contact with
the electrolyte phase 31, a free-flowing fluid phase. The
electrolyte phase is covered by the enzyme membrane 32 that
contains an enzyme, e.g., glucose oxidase, and several functional
polymer layers (as described below). In turn, a bioprotective
membrane 33 covers the enzyme membrane 32 and serves, in part, to
protect the sensor from external forces that may result in
environmental stress cracking of the enzyme membrane 32. Finally,
an angiogenic layer 34 is placed over the bioprotective membrane 33
and serves to promote vascularization in the sensor interface
region.
[0061] A retaining gasket 18 composed of, for example, silicone
rubber, is used to retain the sensor interface capsular attachment
layer 15 (FIGS. 1A-B) and the angiogenic layer 34 and the
bioprotective membrane 33 (not shown). In preferred embodiments,
the angiogenic layer 34 and the bioprotective membrane 33 pass over
the tip of the sensor interface dome 30, over the O-ring 7, and
then under the sensor interface capsular attachment layer 15 and
the retaining gasket 18.
[0062] The present invention contemplates the construction of the
membrane layers of the sensor interface region using standard film
coating techniques. This type of membrane fabrication facilitates
control of membrane properties and membrane testing.
III. Sensor Interface
[0063] As alluded to above and disclosed in FIG. 1C, in a preferred
embodiment, the sensor interface region comprises several different
layers and membranes that cover the electrodes of an implantable
analyte-measuring device. The characteristics of these layers and
membranes are now discussed in more detail. The layers and
membranes prevent direct contact of the biological fluid sample
with the electrodes, while permitting selected substances (e.g.,
analytes) of the fluid to pass therethrough for electrochemical
reaction with the electrodes.
[0064] The membranes used in the sensor interface region are
semipermeable membranes. Generally speaking, the two fundamental
diffusion processes by which a semipermeable membrane can limit the
amount of a substance that passes therethrough are i) diffusion
through the semipermeable membrane as a porous structure and ii)
diffusion through the semipermeable membrane as a monolithic,
homogeneous structure. The present invention is not limited by the
nature of the semipermeable membranes used in the sensor interface
region.
[0065] A semipermeable membrane that comprises a porous structure
consists of a relatively impermeable matrix that includes a
plurality of "microholes" or pores of molecular dimensions.
Transfer through these membranes is primarily due to passage of
substances through the pores (i.e., the membrane acts as a
microporous barrier or sieve). Examples of materials that may be
used to form porous, semipermeable membranes include, but are not
limited to, polyethylene, polyvinylchloride,
polytetrafluoroethylene, polypropylene, polyacrylamide, cellulose
acetate, polymethyl methacrylate, silicone polymers, polycarbonate,
and cellulosic polymers.
[0066] Because diffusion is primarily due to passage of the
substance through pores, the permeability is related to the
effective size of the pores, the membrane thickness, and to the
molecular size of the diffusing substance. As a result, there is
little selectivity in the separation of two chemically or
structurally related molecules, except when their molecular size is
approximately the same as the size of the pore; when this occurs,
forces acting between the substance and the surface of the pore
channel may influence the rate of transfer. In addition, the upper
size limit to diffusion is determined by the largest pore diameter,
and the overall diffusion rate depends on the total number of
pores.
[0067] In contrast, passage of a substance through a monolithic,
homogeneous membrane depends upon selective dissolution and
diffusion of the substance as a solute through a solid, non-porous
film. As used herein, the term "monolithic" means substantially
non-porous and having a generally unbroken surface. The term
"homogeneous", with reference to a membrane, means having
substantially uniform characteristics from one side of the membrane
to the other. However, a membrane may have heterogeneous structural
domains, for example, created by using block copolymers (i.e.,
polymers in which different blocks of identical monomer units
alternate with each other), and still be characterized functionally
as homogeneous with respect to its dependence upon dissolution
rather than sieving to effect separation of substances. A
monolithic membrane can thus be used to selectively separate
components of a solution on the basis of properties other than the
size, shape and density of the diffusing substances. Monolithic,
homogeneous membranes act as a barrier because of the preferential
diffusion therethrough of some substance. They may be formed from
materials such as those previously listed for porous membranes,
including, but not limited to, polyethylene, polyvinylchloride,
tetrafluorethylene, polypropylene, polyacrylamide, polymethyl
methacrylate, silicone polymers, polycarbonate, collagen,
polyurethanes and block copolymers thereof (block copolymers are
discussed in U.S. Pat. Nos. 4,803,243 and 4,686,044, hereby
incorporated by reference).
[0068] A. Angiogenic Layer
[0069] For implantable glucose monitoring devices, a sensor/tissue
interface must be created which provides the sensor with oxygen and
glucose concentrations comparable to that normally available to
tissue comprised of living cells. Absent such an interface, the
sensor is associated with unstable and chaotic performance
indicating that inadequate oxygen and/or glucose are reaching the
sensor. The development of suitable interfaces in other contexts
has been reported. For example, investigators have developed
techniques which stimulate and maintain blood vessels inside a FBC
to provide for the demanding oxygen needs of pancreatic islets
within an implanted membrane. [See, e.g., Brauker et al., Abstract
from 4th World Biomaterials Congress, Berlin (1992)]. These
techniques depend, in part, on the use of a vascularizing layer on
the exterior of the implanted membrane. However,
previously-described implantable analyte-monitoring devices have
not been able to successfully maintain sufficient blood flow to the
sensor interface.
[0070] As described above, the outermost layer of the
electrode-membrane region comprises an angiogenic material. The
angiogenic layer of the devices of the present invention may be
constructed of membrane materials such as hydrophilic
polyvinylidene fluoride (e.g., Duraporeo; Millipore), mixed
cellulose esters (e.g., MF; Millipore), polyvinyl chloride (e.g.,
PVC; Millipore), and other polymers including, but not limited to,
polypropylene, polysulphone, and polymethacrylate. Preferably, the
thickness of the angiogenic layer is about 10 .mu.m to about 20
.mu.m. The angiogenic layer comprises pores sizes of about 0.5 to
about 20 .mu.m, and more preferably about 1.0 .mu.m to about 10
.mu.m, sizes that allow most substances to pass through, including,
e.g., macrophages. The preferred material is expanded PTFE of a
thickness of about 15 .mu.m and pore sizes of about 5 .mu.m to
about 10 .mu.m.
[0071] To further promote stable foreign body capsule structure
without interfering with angiogenesis, an additional outermost
layer of material comprised of a thin low-density non-woven
polyester (e.g., manufactured by Gore) can be laminated over the
preferred PTFE described above. In preferred embodiments, the
thickness of this layer is about 120 .mu.m. This additional thin
layer of material does not interfere with angiogenesis and enhances
the manufacturability of the angiogenic layer. [See U.S. Pat. No.
5,453,278 to Brauker et al., hereby incorporated by reference; PCT
patent Publication Nos. 96/32076, 96/01611, and 92/07525 assigned
to Baxter].
[0072] B. Bioprotective Membrane
[0073] The inflammatory response that initiates and sustains a FBC
is associated with both advantages and disadvantages. Some
inflammatory response is needed to create a new capillary bed in
close proximity to the surface of the sensor in order to i)
continuously deliver adequate oxygen and glucose and ii) create
sufficient tissue ingrowth to anchor the implant and prevent motion
artifact. On the other hand, inflammation is associated with
invasion of tissue macrophages which have the ability to biodegrade
many artificial biomaterials (some of which were, until recently,
considered nonbiodegradable). When activated by a foreign body,
tissue macrophages degranulate, releasing from their cytoplasmic
myeloperoxidase system hypochlorite (bleach), H.sub.2O.sub.2 and
other oxidant species. Both hypochlorite and H.sub.2O.sub.2 are
known to break down a variety of polymers, including polyurethane,
by a phenomenon referred to as environmental stress cracking.
[Phillips et al., J. Biomat. Appl., 3:202-227 (1988); Stokes, J.
Biomat. Appl. 3:228-259 (1988)]. Indeed, environmental stress
cracking has been shown to limit the lifetime and performance of an
enzyme-active polyurethane membrane stretched over the tip of a
glucose sensor. [Updike et al., Am. Soc. Artificial Internal
Organs, 40:157-163 (1994)].
[0074] Because both hypochlorite and H.sub.2O.sub.2 are short-lived
chemical species in vivo, biodegradation will not occur if
macrophages are kept a sufficient distance from the enzyme active
membrane. The present invention contemplates the use of protective
biomaterials of a few microns thickness or more (i.e., a
bioprotective membrane) which are permeable to oxygen and glucose
and are placed over the tip of the sensor to keep the macrophages
from gaining proximity to the sensor membrane. The devices of the
present invention are not limited by the nature of the
bioprotective layer. However, the bioprotective layer should be
biostable for long periods of time (e.g., several years); the
present invention contemplates the use of polymers including, but
not limited to, polypropylene, polysulphone,
polytetrafluoroethylene (PTFE), and poly(ethylene terephthalate)
(PET).
[0075] Preferably, the bioprotective layer is constructed of
expanded PTFE with a pore size of about 0.2 .mu.m to about 0.5
.mu.m and a thickness of about 15 to about 35 .mu.m. Most
preferably, the bioprotective layer is constructed of expanded PTFE
with a pore size of 0.4 .mu.m and a thickness of approximately 25
.mu.m (e.g., Millicell CM-Biopore.RTM.; Millipore).
[0076] C. The Enzyme Membrane
[0077] The present invention contemplates membranes impregnated
with enzyme. It is not intended that the present invention be
limited by the nature of the enzyme membrane. The enzyme membrane
of a preferred embodiment is depicted in FIG. 1C as being a single,
homogeneous structure. However, in preferred embodiments, the
enzyme membrane comprises a plurality of distinct layers. In a
particularly preferred embodiment, the enzyme membrane comprises
the following four layers (in succession from the bioprotective
membrane to the electrolyte phase): i) a resistance layer; ii) an
enzyme layer; iii) an interference layer; and iv) an electrolyte
layer.
[0078] Resistance Layer
[0079] There is a molar excess of glucose relative to the amount of
oxygen in samples of blood. Indeed, for every free oxygen molecule
in extracellular fluid, there are typically more than 100 glucose
molecules present [Updike et al., Diabetes Care 5:207-21(1982)].
However, an immobilized enzyme-based sensor using oxygen (O.sub.2)
as cofactor must be supplied with oxygen in non-rate-limiting
excess in order to respond linearly to changes in glucose
concentration while-not responding to changes in oxygen tension.
More specifically, when a glucose-monitoring reaction is
oxygen-limited, linearity is not achieved above minimal
concentrations of glucose. Without a semipermeable membrane over
the enzyme layer, linear response to glucose levels can be obtained
only up to about 40 mg/dL; however, in a clinical setting, linear
response to glucose levels are desirable up to at least about 500
mg/dL.
[0080] The resistance layer comprises a semipermeable membrane that
controls the flux of oxygen and glucose to the underlying enzyme
layer (i.e., limits the flux of glucose), rendering the necessary
supply of oxygen in non-rate-limiting excess. As a result, the
upper limit of linearity of glucose measurement is extended to a
much higher value than that which could be achieved without the
resistance layer. The devices of the present invention contemplate
resistance layers comprising polymer membranes with
oxygen-to-glucose permeability ratios of approximately 200:1; as a
result, one-dimensional reactant diffusion is adequate to provide
excess oxygen at all reasonable glucose and oxygen concentrations
found in the subcutaneous matrix [Rhodes et al., Anal. Chem.,
66:1520-1529 (1994)].
[0081] In preferred embodiments, the resistance layer has a
thickness of less than about 45 .mu.m, more preferably in the range
of about 15 to about 40 .mu.m and most preferably in the range of
about 20 to about 35 .mu.m.
[0082] Enzyme Layer
[0083] In addition to glucose oxidase, the present invention
contemplates the use of a membrane layer impregnated with other
oxidases, e.g., galactose oxidase, uricase. For an enzyme-based
electrochemical glucose sensor to perform well, the sensor's
response must neither be limited by enzyme activity nor cofactor
concentration. Because enzymes, including the very robust glucose
oxidase, are subject to deactivation as a function of ambient
conditions, this behavior needs to be accounted for in constructing
sensors for long-term use.
[0084] The principle of losing half of the original enzyme activity
in a specific time may be used in calculating how much enzyme needs
to be included in the enzyme layer to provide a sensor of required
lifetime (see Experimental section). Previously, researchers have
found that, when placed in a saline solution at 37.degree. C.,
glucose electrodes lose half of their electrode enzyme activity in
85 to 105 days [See, e.g., Tse and Gough, Biotechnol. Bioeng.
29:705-713 (1987)]. Under reasonable diabetic conditions and normal
enzyme loading (e.g., 2.times.10.sup.-4 M glucose oxidase; see
Example 4), useful sensor lifetimes can last for at least one year.
However, exposure of the sensor to high levels of glucose in
combination with low oxygen levels for prolonged periods can result
in shortened sensor lifetimes. [Rhodes et al., Anal. Chem.,
66:1520-1529 (1994)].
[0085] Excess glucose oxidase loading is required for long sensor
life. The Experimental section provides a procedure that can be
used to determine the appropriate amount of enzyme to be included
in the enzyme layer. When excess glucose oxidase is used, up to two
years of performance is possible from the glucose-monitoring
devices contemplated by the present invention.
[0086] Interference Layer
[0087] The interference layer comprises a thin, hydrophobic
membrane that is non-swellable and has a low molecular weight
cut-off. The interference layer is permeable to relatively low
molecular weight substances, such as hydrogen peroxide, but
restricts the passage of higher molecular weight substances,
including glucose and ascorbic acid. The interference layer serves
to allow analytes and other substances that are to be measured by
the electrodes to pass through, while preventing passage of other
substances.
[0088] The interference layer has a preferred thickness of less
than about 5 .mu.m, more preferably in the range of about 0.1 to
about 5 .mu.m and most preferably in the range of about 0.5 to
about 3 .mu.m.
[0089] Electrolyte Layer
[0090] To ensure electrochemical reaction, the electrolyte layer
comprises a semipermeable coating that maintains hydrophilicity at
the electrode region of the sensor interface. The electrolyte layer
enhances the stability of the interference layer of the present
invention by protecting and supporting the membrane that makes up
the interference layer. Furthermore, the electrolyte layer assists
in stabilizing operation of the device by overcoming electrode
start-up problems and drifting problems caused by inadequate
electrolyte. The buffered electrolyte solution contained in the
electrolyte layer also protects against pH-mediated damage that may
result from the formation of a large pH gradient between the
hydrophobic interference layer and the electrode (or electrodes)
due to the electrochemical activity of the electrode.
[0091] Preferably the coating comprises a flexible,
water-swellable, substantially solid gel-like film having a "dry
film" thickness of about 2.5 .mu.m to about 12.5 .mu.m, preferably
about 6.0 .mu.m. "Dry film" thickness refers to the thickness of a
cured film cast from a coating formulation onto the surface of the
membrane by standard coating techniques. The coating formulation
comprises a premix of film-forming polymers and a crosslinking
agent and is curable upon the application of moderate heat.
[0092] Suitable coatings are formed of a curable copolymer of a
urethane polymer and a hydrophilic film-forming polymer.
Particularly preferred coatings are formed of a polyurethane
polymer having anionic carboxylate functional groups and non-ionic
hydrophilic polyether segments, which is crosslinked in the present
of polyvinylpyrrolidone and cured at a moderate temperature of
about 50.degree. C.
[0093] Particularly suitable for this purpose are aqueous
dispersions of fully-reacted colloidal polyurethane polymers having
cross-linkable carboxyl functionality (e.g., BAYBONDO; Mobay
Corporation). These polymers are supplied in dispersion grades
having a polycarbonate-polyurethane backbone containing carboxylate
groups identified as XW-121 and XW-123; and a
polyester-polyurethane backbone containing carboxylate groups,
identified as XW-110-2. Particularly preferred is BAYBOND.RTM. 123,
an aqueous anionic dispersion of an aliphate polycarbonate urethane
polymer sold as a 35 weight percent solution in water and
co-solvent N-methyl-2-pyrrolidone.
[0094] Polyvinylpyrrolidone is also particularly preferred as a
hydrophilic water-soluble polymer and is available commercially in
a range of viscosity grades and average molecular weights ranging
from about 18,000 to about 500,000, under the PVP K.RTM.
homopolymer series by BASF Wyandotte and by GAF Corporation.
Particularly preferred is the homopolymer having an average
molecular weight of about 360,000 identified as PVP-K90 (BASF
Wyandotte). Also suitable are hydrophilic, film-forming copolymers
of N-vinylpyrrolidone, such as a copolymer of N-vinylpyrrolidone
and vinyl acetate, a copolymer of N-vinylpyrrolidone,
ethylmethacrylate and methacrylic acid monomers, and the like.
[0095] The polyurethane polymer is crosslinked in the presence of
the polyvinylpyrrolidone by preparing a premix of the polymers and
adding a cross-linking agent just prior to the production of the
membrane. Suitable cross-linking agents can be carbodiimides,
epoxides and melamine/formaldehyde resins. Carbodiimide is
preferred, and a preferred carbodiimide crosslinker is UCARLNK.RTM.
XL-25 (Union Carbide).
[0096] The flexibility and hardness of the coating can be varied as
desired by varying the dry weight solids of the components in the
coating formulation. The term "dry weight solids" refers to the dry
weight percent based on the total coating composition after the
time the crosslinker is included. A preferred useful coating
formulation can contain about to about 20 dry weight percent,
preferably about 8 dry weight percent, polyvinylpyrrolidone; about
3 to about 10 dry weight percent preferably about 5 dry 5 weight
percent cross-linking agent; and about 70 to about 91 weight
percent, preferably about 87 weight percent of a polyurethane
polymer, preferably a polycarbonate-polyurethane polymer. The
reaction product of such a coating formulation is referred to
herein as a water-swellable cross-linked matrix of polyurethane and
polyvinylpyrrolidone.
[0097] D. The Electrolyte Phase
[0098] The electrolyte phase is a free-fluid phase comprising a
solution containing at least one compound, usually a soluble
chloride salt, that conducts electric current. The electrolyte
phase flows over the electrodes (see FIG. 1C) and is in contact
with the electrolyte layer of the enzyme membrane. The devices of
the present invention contemplate the use of any suitable
electrolyte solution, including standard, commercially available
solutions.
[0099] Generally speaking, the electrolyte phase should have the
same or less osmotic pressure than the sample being analyzed. In
preferred embodiments of the present invention, the electrolyte
phase comprises normal saline.
[0100] E. Electrode
[0101] The electrode assembly of this invention may also be used in
the manner commonly employed in the making of amperometric
measurements. A sample of the fluid being analyzed is placed in
contact with a reference electrode, e.g., silver/silver-chloride,
and the electrode of this invention which is preferably formed of
platinum. The electrodes are connected to a galvanometer or
polarographic instrument and the current is read or recorded upon
application of the desired D.C. bias voltage between the
electrodes.
[0102] The ability of the present device electrode assembly to
accurately measure the concentration of substances such as glucose
over a broad range of concentrations in fluids including undiluted
whole blood samples enables the rapid and accurate determination of
the concentration of those substances. That information can be
employed in the study and control of metabolic disorders including
diabetes.
IV. Sensor Implantation and Radiotelemetric Output
[0103] Long-term sensor performance is best achieved, and
transcutaneous bacterial infection is eliminated, with implanted
devices capable of radiotelemetric output. The present invention
contemplates the use of radiotelemetry to provide data regarding
blood glucose levels, trends, and the like. The term
"radiotelemetry" refers to the transmission by radio waves of the
data recorded by the implanted device to an ex vivo recording
station (e.g., a computer), where the data is recorded and, if
desired, further processed.
[0104] Although totally implanted glucose sensors of three month
lifetime, with radiotelemetric output, have been tested in animal
models at intravenous sites [see, e.g. Armour et al., Diabetes,
39:1519-1526 (1990)], subcutaneous implantation is the preferred
mode of implantation [see, e.g., Gilligan et al., Diabetes Care
17:882-887 (1994)]. The subcutaneous site has the advantage of
lowering the risk for thrombophlebitis with hematogenous spread of
infection and also lowers the risk of venous thrombosis with
pulmonary embolism. In addition, subcutaneous placement is
technically easier and more cost-effective than intravenous
placement, as it may be performed under local anesthesia by a
non-surgeon health care provider in an outpatient setting.
[0105] Preferably, the radiotelemetry devices contemplated for use
in conjunction with the present invention possess features
including small package size, adequate battery life, acceptable
noise-free transmission range, freedom from electrical
interference, and easy data collection and processing.
Radiotelemetry provides several advantages, one of the most
important of which is the ability of an implanted device to measure
analyte levels in a sealed-off, sterile environment.
[0106] The present invention is not limited by the nature of the
radiotelemetry equipment or methods for its use. Indeed,
commercially available equipment can be modified for use with the
devices of the present invention (e.g., devices manufactured by
Data Sciences). Similarly, custom-designed radiotelemetry devices
like those reported in the literature can be used in conjunction
with the implantable analyte-measuring devices of the present
invention [see, e.g., McKean and Gough, IEEE Trans. Biomed. Eng.
35:526-532 (1988); Shichiri et al., Diabetes Care 9:298-301 (1986);
and Shults et al., IEEE Trans. Biomed. Eng. 41:937-942 (1994)]. In
a preferred embodiment, transmitters are programmed with an
external magnet to transmit at 4-, 32-, or 256-second intervals
depending on the need of the subject; presently, battery lifetimes
at the current longest transmission intervals (about 256 seconds)
is approximately up to two years.
V. Response Time and Calibration
[0107] Every measurement method reports data with some delay after
the measured event. For data to be useful, this delay must be
smaller than some time depending on the needs of the method. Thus,
response time of the current invention has been carefully studied.
The use of the term "initial response" is not to be confused with
the term "response time." After a step function change in glucose
concentration, the time delay before the first unequivocal change
in sensor signal occurs is the "initial response," while the
following time delay to reach 90% of the steady-state signal
development is the "response time." "Response time" is the factor
which normally controls how quickly a sensor can track a
dynamically changing system.
[0108] Furthermore, the time required before a glucose sensor in a
FBC will indicate an initial response to a bolus intravenous
glucose injection is a function of the animal "circulation time"
and the sensor's "initial response". The circulation time is the
time required for a bolus glucose injection to reach the site of
sensor implantation.
[0109] Generally speaking, equilibration between vascular and
interstitial compartments for glucose is so rapid that it plays no
role in either the initial response or the observed response time.
If the tip of the sensor is in intimate contact with the
interstitial compartment (e.g., FBC), then there is no significant
delay in glucose diffusing from the capillary lumen to the tip of
the sensor. The inventors have found that the glucose sensors of
the present invention provide initial responses of about 30 seconds
in dogs about half of which is circulation time. The dog model
represents a useful and accepted model for determining the efficacy
of glucose monitoring devices.
[0110] While the devices of the present invention do not require a
specific response time, in preferred embodiments of the present
invention, the in vitro 90% response times at 37.degree. C. for
subsequently subcutaneously implanted devices are in the range of 2
to 5 minutes in dogs. Though the use of the devices of the present
invention does not require an understanding of the factors that
influence response time or the factors' mechanisms of action, in
vivo response times are believed to be primarily a function of
glucose diffusion through the sensor membrane (e.g., a 40-60 micron
membrane). Of note, response times of up to about 10 minutes do not
limit the clinical utility of tracking blood glucose in diabetic
patients because physiologic or pathologic glucose levels do not
change more rapidly than a few percent per minute.
[0111] In calibrating the glucose sensors of the present invention,
a single point recalibration of the sensor at four-week intervals
against an acceptable glucose reference method is preferred (e.g.,
calibration against blood obtained from a finger-prick). Generally
speaking, the recalibration amounts to a simple adjustment in
sensor gain. The sensor offset current (i.e., the current at 0
mg/dL glucose) needs to remain invariant over the duration of the
implant for the sensor to provide optimal data.
Experimental
[0112] The following examples serve to illustrate certain preferred
embodiments and aspects of the present invention and are not to be
construed as limiting the scope thereof.
[0113] In the preceding description and the experimental disclosure
which follows, the following abbreviations apply: Eq and Eqs
(equivalents); mEq (milliequivalents); M (molar); mM (millimolar)
.mu.M (micromolar); N (Normal); mol (moles); mmol (millimoles);
.mu.mol (micromoles); nmol (nanomoles); g (grams); mg (milligrams);
.mu.g (micrograms); Kg (kilograms); L (liters); mL (milliliters);
dL (deciliters); .mu.L (microliters); cm (centimeters); mm
(millimeters); sm (micrometers); nm (nanometers); h and hr (hours);
min. (minutes); s and sec. (seconds); .degree. C. (degrees
Centigrade); Astor Wax (Titusville, Pa.); BASF Wyandotte
Corporation (Parsippany, N.J.); Data Sciences, Inc. (St. Paul,
Minn.); DuPont (DuPont Co., Wilmington, Del.); Exxon Chemical
(Houston, Tex.); GAF Corporation (New York, N.Y.); Markwell Medical
(Racine, Wis.); Meadox Medical, Inc. (Oakland, N.J.); Mobay (Mobay
Corporation, Pittsburgh, Pa.); Sandoz (East Hanover, N.J.); and
Union Carbide (Union Carbide Corporation; Chicago, Ill.).
EXAMPLE 1
[0114] The polyurethanes are preferably prepared as block
copolymers by solution -polymerization techniques as generally
described in Lyman [J. Polymer Sci. 45:49 (1960)]. Specifically, a
two-step solution polymerization technique is used in which the
poly(oxyethylene) glycol is first "capped" by reaction with a
diisocyanate to form a macrodiisocyanate. The macrodiisocyanate is
then coupled with a diol (or diamine) and the diisocyanate to form
a block copolyetherurethane (or a block copolyurethaneurea). The
resulting block copolymers are tough and elastic and may be
solution-cast in N,N-dimethylformamide to yield clear films that
demonstrate good wet strength when swollen in water.
[0115] In particular, a mixture of 8.4 g (0.006 mol),
poly(oxyethylene) glycol (CARBOWAX.RTM. 1540, Union Carbide), and
3.0 g (0.012 mol) 4,4'-diphenylmethane diisocyanate in 20 mL
dimethyl sulfoxide/4-methyl-2-pentanone (50/50) is placed in a
three-necked flask equipped with a stirrer and condenser and
protected from moisture. The reaction mixture is stirred and heated
at 110.degree. C. for about one hour. To this clear solution is
added 1.5 g (0.014 mol) 1,5-pentanediol and 2.0 g (0.008 mol)
4,4'-diphenylmethane diisocyanate.
[0116] After heating at 110.degree. C. for an additional two hours,
the resulting viscous solution is poured into water. The tough,
rubbery, white polymer precipitate that forms is chopped in a
Waring Blender, washed with water and dried in a vacuum oven at
about 60.degree. C. The yield is essentially quantitative. The
inherent viscosity of the copolymer in N,N-dimethyl formamide is
0.59 at 30.degree. C. (at a concentration of about 0.05 percent by
weight).
EXAMPLE 2
[0117] As previously described, the electrolyte layer, the membrane
layer closest to the electrode, can be coated as a water-swellable
film. This example illustrates a coating comprising a polyurethane
having anionic carboxylate functional groups and hydrophilic
polyether groups and polyvinylpyrrolidone (PVP) that can be
cross-linked by carbodiimide.
[0118] A coating preparation is prepared comprising a premix of a
colloidal aqueous dispersion of particles of a urethane polymer
having a polycarbonate-polyurethane (PC-PU) backbone containing
carboxylate groups and the water-soluble hydrophilic polymer, PVP,
which is crosslinked by the addition of the cross-linking agent
just before production of the coated membrane. Example coating
formulations are illustrated in Table 1.
TABLE-US-00001 TABLE 1 A B C Dry Dry Dry Weight Weight Weight % % %
Weight Solids Weight Solids Weight Solids Premix PVP.sup.1 48 6 64
8 160 20 PC-PV.sup.2 260 91 248 87 200 70 Cross- Linking Agent
Carbodiimide.sup.3 6 3 10 5 20 10 Totals 314 100 322 100 380 100
.sup.1Aqueous solution containing 12.5 weight percent PVP prepared
from polyvinylpyrrolidone having a number average molecular weight
of about 360,000 sold as a powder under the trademark BASF K90 by
BASF Wyandotte Corporation. .sup.2Colloidal dispersion of a
polycarbonatepolyurethane (PCPU) polymer at about 35 weight percent
solids in a co-solvent mixture of about 53 weight percent water and
about 12 weight percent N-methyl-2-pyrrolidone (BAYBOND .RTM. 123
or XW123; Mobay Corporation). As supplied, the dispersion has a pH
of about 7.5-9.0. .sup.3Carbodiimide (UCARLNK .RTM. XL25SE, Union
Carbide Corporation) supplied at about 50 weight percent solids in
a solvent solution of propylene glycol monomethylether acetate.
[0119] The viscosity and pH of the premix can be controlled and
maintained during processing and to prolong its useful life by
adding water or adjusting the pH with dilute ammonia solution or an
equivalent base prior to adding the crosslinker.
[0120] For production, the coating is applied with a Mayer rod onto
the unbound surface of a multilayered membrane. The amount of
coating applied should cast a film having a "dry film" thickness of
about 2.5 .mu.m to about 12.5 .mu.m, preferably about 6.0 .mu.m.
The coating is dried above room temperature preferably at about
50.degree. C. This coating dries to a substantially solid gel-like
film that is water swellable to maintain electrolyte between the
membrane covering the electrode and the electrode in the electrode
assembly during use.
EXAMPLE 3
[0121] The following procedure was used to determine the amount of
enzyme to be included in the enzyme layer. It is to be understood
that the present invention is not limited to the use of this or a
similar procedure, but rather contemplates the use of other
techniques known in the art.
[0122] A starting glucose oxidase concentration of
2.times.10.sup.-4 M was calculated from the enzyme weight and the
final volume of the enzyme layer. Thereafter, a series of eight
additional membrane formulations was prepared by decrementing
enzyme concentration in 50% steps (referred to as a change of one
"half loading") down to 7.8.times.10.sup.-7 M. Sensor responses
were then collected for this range of enzyme loadings and compared
to computer-simulated sensor outputs. The simulation parameter set
used included previously-determined membrane permeabilities and the
literature mechanisms and kinetics for glucose oxidase. [Rhodes et
al., Anal. Chem., 66:1520-1529 (1994)].
[0123] There was a good match of real-to-simulated sensor output at
all loadings (data not shown). Approximately a six-to-seven "half
loading" drop in enzyme activity was required before the sensor
output dropped 10%; another two-to-three half loading drop in
enzyme activity was required to drop the sensor response to 50% of
the fully loaded sensor response. These results indicate that, at
the loading used and the decay rates measured, up to two years of
performance is possible from these sensors when the sensor does not
see extended periods of high glucose and physiologically low
O.sub.2 concentrations.
EXAMPLE 4
[0124] This example illustrates long-term glucose sensor device
response following subcutaneous implantation of sensor devices
contemplated by the present invention into a dog. The stages of FBC
development are indicated by the long term glucose sensor device
response.
[0125] FIG. 2 graphically depicts glucose levels as a function of
the number of days post-implant. The data in FIG. 2 was taken at
four-minute intervals for 60 days after implantation. Sensor
response is calculated from a single preimplant calibration at
37.degree. C. Normal canine fasting glucose concentration of 5.5 mM
is shown for comparison.
[0126] The data set forth in FIG. 2 can be used to illustrate the
four typically identifiable phases in FBC formation. Phase 1 shows
rapidly dropping response from the time of implant to, in this
case, day 3. Though an understanding of the mechanism for this drop
in sensor output is not required in order to practice the present
invention, it is believed to reflect low PO.sub.2 and low glucose
present in fluid contacting the sensor. Phase 2 shows intermittent
sensor-tissue contact in seroma fluid from, in this case, day 3 to
about day 13. During this phase, fragile new tissue and blood
supply intermittently make contact with the sensor (which is
surrounded by seroma fluid). Phase 3 shows stabilization of
capillary supply between, in this case, days 13 and 22. More
specifically, the noise disappears and sensor output rises over
approximately six days to a long term level associated with
tracking of FBC glucose. Again, though an understanding of this
effect is not required to practice the present invention, the
effect is believed to reflect consistent contact of FBC tissue with
the sensor surface. Phase 4 from, in this case, day 22 to day 60,
shows duration of useful sensor device life. While there are timing
variations of the stages from sensor device to sensor device,
generally speaking, the first three steps of this process take from
3 days to three weeks and continuous sensing has been observed for
periods thereafter (e.g., for periods of 150 days and beyond).
EXAMPLE 5
[0127] In addition to collecting normoglycemic or non-diabetic dog
data from the sensor of the present invention as shown in Example
4, calibration stability, dynamic range, freedom from oxygen
dependence, response time and linearity of the sensor can be
studied by artificial manipulation of the intravenous glucose of
the sensor host.
[0128] This was done in this example via infusion of a 15 g bolus
of 50% sterile Dextrose given intravenously in less than about 20
seconds. Reference blood glucose data was then taken from a
different vein at 2-5 minute intervals for up to 2 hours after
bolus infusion. FIG. 3 depicts correlation plots of six bolus
infusion studies, at intervals of 7-10 days on one sensor of the
present invention. Sensor glucose concentrations are calculated
using a single 37.degree. C. in vitro preimplantation calibration.
The sensor response time is accounted for in calculating the sensor
glucose concentrations at times of reference blood sampling by time
shifting the sensor data 4 minutes.
[0129] As with any analytical system, periodic calibration should
be performed with the devices of the present invention. Thus, the
present invention contemplates some interval of calibration and/or
control testing to meet analytical, clinical and regulatory
requirements.
EXAMPLE 6
[0130] This example describes experiments directed at sensor
accuracy and long-term glucose sensor response of several sensor
devices contemplated by the present invention.
[0131] Pre-Implant In Vitro Evaluation
[0132] In vitro testing of the sensor devices was accomplished in a
manner similar to that previously described. [Gilligan et al.,
Diabetes Care 17:882-887 (1994)]. Briefly, sensor performance was
verified by demonstrating linearity to 100 mg/dL glucose
concentration steps from 0 mg/dL through 400 mg/dL (22 mM) with a
90% time response to the glucose steps of less than 5 minutes. A
typical satisfactory response to this protocol is shown in FIG. 4.
Modulating dissolved oxygen concentration from a PO.sub.2 of 150
down to 30 mm Hg (0.25 to 0.05 mM) showed no more than a 10% drop
in sensor output at 400 mg/dL for the preferred sensor devices of
the present invention. Stability of calibration was maintained
within 10% for one week before the final bioprotective and
angiogenesis membranes were added to finalize the implant package.
A final calibration check was made and had to be within 20% of the
prior results for the sensor to be passed on to the implant stage.
These final calibration factors (linear least squares regression
for the zero glucose current and output to 100 mg/dL current) are
used for the initial in vivo calibration. Sensor devices were then
wet sterilized with 0.05% thimerosal for 24 hours just prior to
implantation.
[0133] In Vivo Testing
[0134] Six sensor devices meeting the parameters described above
were surgically implanted under general anesthesia (pentothal
induction to effect, followed by halothane maintenance) into the
paravertebral subcutaneous tissue of the same mongrel non-diabetic
dog. A two-inch skin incision was made several inches from the
spine for each implant allowing the creation of a tight-fitting
subcutaneous pouch by blunt dissection. The implant was then
inserted into the pouch in sensor-down configuration. Subcutaneous
tissue was then closed with 3-0 vicryl and skin with 2-0 nylon.
Animals were closely monitored for discomfort after surgery and
analgesics administered if necessary.
[0135] These sensor devices were implanted two-at-a-time in the
same dog at approximately six week intervals. Four of the sensor
devices were covered with a PTFE-comprising angiogenic layer (these
sensor devices were designated Sensors 1901, 1902, 1903, and 1905),
while two of the sensor devices served as-control sensor devices
and did not contain an angiogenic layer, i.e., they contained a
bioprotective membrane and the underlying sensor interface
structures, as previously described (these sensor devices were
designated Sensors 1904 and 1906). To insure anchoring of the
device into the subcutaneous tissue, the sensor-side of each
implant, except for just over the tip of the sensor, was jacketed
in surgical grade double velour polyester fabric (Meadox Medical,
Inc.). All sensor devices were tracked after implantation at
four-minute intervals using radiotelemetry to follow the long-term
sensor response to normoglycemia, allowing verification of the
long-term stability of the sensors. To screen for sensor response
to changing glucose on selected days following implantation, the
response to 0.5 mg glucagon administered subcutaneously was
assessed. Responding sensors were identified by a
characteristically stable signal prior to glucagon administration
followed by a substantial increase in signal within 20 minutes of
glucagon injection. The sensor transients then reversed and
returned to the prior signal levels within one hour after glucagon
injection.
[0136] To determine in vivo sensor response times, short-term
stability, linearity to glucose concentration, and possible oxygen
cofactor limitation effects, glucose infusion studies of up to five
hours duration were performed on the dog. These studies were run
approximately once every three weeks. The dog was pretrained to
rest comfortably and was fully alert during this testing. These
experiments used the somatostatin analog octreotide
(SANDOSTATIN.RTM., Sandoz) to inhibit the release of insulin,
allowing for a slow ramping of blood glucose to the 400-500 mg/dL
concentration range.
[0137] Sensors were monitored at 32-second intervals to allow
simultaneous tracking of up to six sensor devices. In this
protocol, octreotide was injected (36-50 .mu.g/kg) 15-20 minutes
before initiation of the glucose infusion. Two peripheral veins
were cannulated in the dog to allow for glucose infusion and blood
glucose sampling. Ten percent dextrose (0.55 mM) was continuously
infused at gradually increasing rates to provide smooth increases
in blood glucose from the approximate fasting glucose concentration
of about 100 mg/dL to greater than 400 mg/dL. This infusion
protocol provides sensor glucose concentration data which can be
correlated with reference plasma glucose values when blood samples
were drawn from the animal every 5-to-10 minutes. The primary
reference glucose determinations were made using a hexokinase
method on the DuPont Dimension AR.RTM.. A DIRECT 30/30.RTM. meter
(Markwell Medical) was also used during the course of the
experiment to serve as a secondary monitor for the reference blood
glucose values and estimate when 400 mg/dL had been reached. At
this point the glucose infusion pump was turned off and the blood
glucose allowed to return to its normal level.
[0138] An additional variation of the protocol described above
involved studying the effects of insulin administration on blood
glucose concentration prior to the octreotide injection. For these
studies 5 units of insulin were injected intravenously, the blood
glucose tracked down to 40 mg/dl with the DIRECT 30/30.RTM.
(Markwell Medical), the octreotide injection made as before, and
the infusion pump then started. While the initial glucose pump rate
was the same, it was increased faster than before to counteract the
insulin and to maintain the same experimental timing.
[0139] Once studies were completed, the data was initially analyzed
using the final in vitro sensor calibration factors to calculate
the implanted sensor glucose concentration. If changes were needed
in these factors to optimize the linear regression of sensor to
reference blood glucose they were made and noted and followed over
the lifetime of the sensor device.
[0140] At varying points in time, the implanted sensor devices
became less than optimal and were then explanted to determine the
underlying cause (less than optimal was defined as the inability to
accurately track glucose infusion during two successive tests).
Explantation surgical protocols were very similar to those used in
the implantation procedure except that the foreign body capsule was
opened around the perimeter of the oval implant. The back and sides
of the housing had no tissue attachment (as they were not covered
with polyester velour), and thus easily separated from the
surrounding tissue. The top of the sensor device with attached
capsule was then carefully cut free from the subcutaneous
tissues.
[0141] Once explanted, the sensor devices were carefully examined
under a dissecting microscope to look at the state of the capsule
tissue contacting the sensor membranes. Once this had been
characterized and documented, the tissue was carefully removed from
the membrane surface and saved for histological examination. If
sensor visualization demonstrated intact membrane layers an initial
in vitro calibration check was performed. The sensors were then
disassembled from the top membrane down (i.e., from the membrane
furthest from the electrodes) with a glucose and hydrogen peroxide
calibration check made after removal of each layer. This allowed
differentiation of the mechanisms leading to less than optimal
results in the membranes and determination of whether processes
such as environmental stress cracking, biofouling, or loss of
enzyme activity were occurring.
[0142] Results And Conclusions
[0143] Typical Glucose Infusion Studies: The six sensor devices
were tracked for 20-150 days and were evaluated using the
octreotide-glucose infusion protocol. FIGS. 5A, 5B, and 5C
graphically depict three in vivo sensor response curves (using best
case calibration factors) plotted in conjunction with the reference
blood glucose values for Sensor 1903 at post-implant times of 25,
88, and 109 days; this data is representative of the data
obtainable with the sensor devices of the present invention.
Referring to FIGS. 5A-C, the arrow labeled "#1" indicates
octreotide injection, the arrow labeled "#2" indicates the turning
on of the glucose infusion pump, and the arrow labeled "#3"
indicates the turning off of this pump. The 90% response time for
this sensor over its lifetime ranged from 5-to-10 minutes and was 5
minutes for the data shown. Such time responses are adequate, since
changes in diabetic patients occur at slower rates than used with
infusion protocols.
[0144] FIG. 6 graphically depicts sensor glucose versus reference
glucose (for Sensor 1903) using the single set of calibration
factors from day 88. As depicted in FIG. 6, when sensor glucose is
plotted versus reference glucose, the changes in sensor calibration
over the lifetime of the sensor become apparent. These changes are
reflected primarily in the output sensitivity to a known glucose
concentration step while the zero current remained quite stable.
The results suggest that in vivo recalibration every month would be
preferred for this sensor to provide optimal glucose tracking.
Performance Comparisons
[0145] Angiogenesis Stimulating Membrane Sensors vs. Control
Membrane Sensors:
[0146] Generally speaking, demonstration of improvement in a sensor
can be judged by noting whether significant improvements in sensor
start up time, increased yields of operating glucose sensors,
extension of sensor lifetimes, arid maintenance of calibration
factors occurs. The lifetime of a glucose sensor can be defined as
the time of first glucose sensing (in this case during a glucagon
challenge) to the last glucose infusion study which provides
correct glucose trends to concentration changes. All sensors showed
glucose tracking and only one sensor showed a duration of less than
10 days. Average sensor lifetimes of 84.+-.55 days were observed
with the sensors containing the angiogenesis-stimulating membrane,
values superior to the control sensors which only showed lifetimes
of 35.+-.10 days. In addition, one of the sensors incorporating the
angiogenic membrane provided optimal data to 150 days.
[0147] The description and experimental materials presented above
are intended to be illustrative of the present invention while not
limiting the scope thereof. It will be apparent to those skilled in
the art that variations and modifications can be made without
departing from the spirit and scope of the present invention.
* * * * *