U.S. patent application number 12/247137 was filed with the patent office on 2009-01-29 for implantable analyte sensor.
This patent application is currently assigned to DexCom, Inc.. Invention is credited to James H. Brauker, Mark Brister, Paul V. Neale, James R. Petisce, Mark Shults.
Application Number | 20090030294 12/247137 |
Document ID | / |
Family ID | 35186249 |
Filed Date | 2009-01-29 |
United States Patent
Application |
20090030294 |
Kind Code |
A1 |
Petisce; James R. ; et
al. |
January 29, 2009 |
IMPLANTABLE ANALYTE SENSOR
Abstract
An implantable analyte sensor including a sensing region for
measuring the analyte and a non-sensing region for immobilizing the
sensor body in the host. The sensor is implanted in a precisely
dimensioned pocket to stabilize the analyte sensor in vivo and
enable measurement of the concentration of the analyte in the host
before and after formation of a foreign body capsule around the
sensor. The sensor further provides a transmitter for RF
transmission through the sensor body, electronic circuitry, and a
power source optimized for long-term use in the miniaturized sensor
body.
Inventors: |
Petisce; James R.; (San
Diego, CA) ; Brister; Mark; (Encinitas, CA) ;
Shults; Mark; (Madison, WI) ; Brauker; James H.;
(Cement City, MI) ; Neale; Paul V.; (San Diego,
CA) |
Correspondence
Address: |
KNOBBE, MARTENS, OLSEN & BEAR, LLP
2040 MAIN STREET, FOURTEENTH FLOOR
IRVINE
CA
92614
US
|
Assignee: |
DexCom, Inc.
San Diego
CA
|
Family ID: |
35186249 |
Appl. No.: |
12/247137 |
Filed: |
October 7, 2008 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
|
|
10838658 |
May 3, 2004 |
|
|
|
12247137 |
|
|
|
|
Current U.S.
Class: |
600/302 ;
264/104 |
Current CPC
Class: |
A61B 5/14532 20130101;
A61B 5/14865 20130101; A61B 5/0031 20130101 |
Class at
Publication: |
600/302 ;
264/104 |
International
Class: |
A61B 5/07 20060101
A61B005/07; B28B 11/04 20060101 B28B011/04 |
Claims
1. An implantable analyte sensor for measuring an analyte
concentration in a host, the sensor comprising: a sensor body
substantially formed from a water vapor permeable material; and
electrical components encapsulated within the sensor body, wherein
the electrical components comprise RF circuitry and an antenna
adapted for RF transmission from the sensor in vivo to a receiver
ex vivo, wherein the RF circuitry is spaced a fixed distance from
the sensor body so as to support a dielectric constant that enables
RF transmission between the sensor in vivo to the receiver ex
vivo.
2. The sensor according to claim 1, wherein the fixed distance
comprises a configuration that reduces water permeability
therein.
3. The sensor according to claim 2, wherein the configuration
comprises conformal coating.
4. The sensor according to claim 3, wherein the conformal coating
comprises Parylene.
5. The sensor according to claim 2, wherein the configuration
comprises epoxy.
6. The sensor according to claim 2, wherein the configuration
comprises glass.
7. The sensor according to claim 2, the configuration comprises one
or more hermetic containers.
8. An analyte sensor for RF transmission between the analyte sensor
in vivo and a receiver ex vivo, the sensor comprising: a sensor
body comprising RF circuitry encapsulated within a substantially
water vapor permeable body that enables RF transmission
therethrough; a sensing region located on an outer surface of the
sensor body for measuring an analyte in soft tissue; a biointerface
material disposed adjacent to the sensing region that supports
vascularized tissue ingrowth for transport of the analyte to the
sensing region; and an anchoring material on a non-sensing outer
surface of the sensor body that supports tissue ingrowth for
immobilization of the sensor body in soft tissue.
9. The sensor according to claim 8, further comprising an antenna
encapsulated within the sensor body.
10. The sensor according to claim 9, further comprising a power
source encapsulated within the sensor body.
11. The sensor according to claim 8, wherein the sensor body is
formed from plastic.
12. The sensor according to claim 11, wherein the plastic comprises
epoxy.
13. The sensor according to claim 8, wherein the sensor body is
molded around the RF circuitry.
14. The sensor according to claim 8, further comprising an
electrode system exposed at the sensing region.
15. The sensor according to claim 14, wherein the electrode system
extends through the water vapor permeable body and is operably
connected to the RF circuitry.
16. The sensor according to claim 8, wherein the biointerface
material comprises a solid portion with a plurality of
interconnected cavities.
17. The sensor according to claim 8, wherein the biointerface
material further comprises a domain proximal to the sensing region
that is impermeable to cells or cell processes and is permeable to
the passage of the analyte.
18. A method for manufacturing an analyte sensor, the method
comprising: providing sensor electronics designed to process
signals from the sensor; conformally coating the sensor electronics
with a material that has a first water permeability rate; and
molding a water vapor permeable material that has a second water
permeability rate around the coated sensor electronics to form a
substantially seamless sensor body, wherein the second water
permeability rate is greater than the first water penetration
rate.
19. The method according to claim 18, wherein the molding comprises
a two-step molding process to form the seamless sensor body.
20. The method according to claim 19, wherein the two-step molding
process comprises: holding a first portion of the coated sensor
electronics and molding around a second portion of the coated
sensor electronics; and holding a portion of the cured sensor body
and molding around the first portion of the coated sensor
electronics.
21. An implantable analyte sensor comprising: a body comprising a
material which is permeable to water vapor, said body further
comprising a sensing region for measuring levels of an analyte; and
a transmitter within said body for transmitting the measurements
obtained by said sensing region, wherein at least a portion of said
transmitter is spaced from said body by a material adapted to
reduce water from penetrating therein.
22. The sensor of claim 21, wherein said transmitter comprises an
oscillator and at least a portion of said oscillator is spaced from
said body by said material adapted to inhibit fluid from
penetrating therein.
23. The sensor of claim 22, wherein said oscillator comprises an
inductor and wherein said inductor is spaced from said body by said
material adapted to inhibit fluid from penetrating therein.
24. The sensor of claim 22, wherein said oscillator comprises a
voltage controlled oscillator.
Description
CROSS-REFERENCE TO RELATED APPLICATION
[0001] This application is a division of U.S. application Ser. No.
10/838,658 filed May 3, 2004, which is incorporated by reference
herein in its entirety, and is hereby made a part of this
specification.
FIELD OF THE INVENTION
[0002] The present invention relates generally to systems and
methods for making and using an implantable analyte sensor.
BACKGROUND OF THE INVENTION
[0003] Diabetes mellitus is a disorder in which the pancreas cannot
create sufficient insulin (Type I or insulin dependent) and/or in
which insulin is not effective (Type 2 or non-insulin dependent).
In the diabetic state, the victim suffers from high blood sugar,
which may cause an array of physiological derangements (for
example, kidney failure, skin ulcers, or bleeding into the vitreous
of the eye) associated with the deterioration of small blood
vessels. A hypoglycemic reaction (low blood sugar) may be induced
by an inadvertent overdose of insulin, or after a normal dose of
insulin or glucose-lowering agent accompanied by extraordinary
exercise or insufficient food intake.
[0004] Conventionally, a diabetic person carries a self-monitoring
blood glucose (SMBG) monitor, which typically comprises
uncomfortable finger pricking methods. Due to the lack of comfort
and convenience, a diabetic will normally only measure his or her
glucose level two to four times per day. Unfortunately, these time
intervals are so far spread apart that the diabetic will likely
find out too late, sometimes incurring dangerous side effects, of a
hyper- or hypo-glycemic condition. In fact, it is not only unlikely
that a diabetic will take a timely SMBG value, but the diabetic
will not know if their blood glucose value is going up (higher) or
down (lower) based on conventional methods, inhibiting their
ability to make educated insulin therapy decisions.
[0005] The prior art discloses a variety of analyte sensors that
provide complex, short-term, transcutaneous, or partially
implantable analyte sensors. Unfortunately, each of these sensors
suffers from various disadvantages, such as lack of continuous care
(short-term sensors), discomfort (transcutaneous and partially
implantable sensors), and inconvenience (sensors with multiple
components).
SUMMARY OF THE INVENTION
[0006] There is a need for a device a long-term, implantable
analyte sensor that functions accurately and reliably, to provide
improved patient convenience and care.
[0007] Accordingly, in a first embodiment, an implantable analyte
sensor for measuring an analyte in a host is provided, the sensor
including: a sensor body including a sensing region for measuring
the analyte and a non-sensing region for immobilizing the sensor
body in the host; a first biointerface material adjacent to the
sensing region, wherein the first biointerface material includes a
porous architecture that promotes vascularized tissue ingrowth and
interferes with barrier cell layer formation, for allowing analyte
transport to the sensing region in vivo; and a second biointerface
material adjacent to at least a portion of the non-sensing region,
wherein the second biointerface material includes a porous
architecture that promotes tissue ingrowth for anchoring the sensor
in vivo.
[0008] In an aspect of the first embodiment, the first biointerface
material further includes a domain proximal to the sensing region
that is impermeable to cells or cell processes and is permeable to
the passage of the analyte.
[0009] In an aspect of the first embodiment, the second
biointerface material and the first biointerface material include
porous silicone, and wherein first biomaterial material includes
porous silicone with a through porosity substantially across the
entire material.
[0010] In an aspect of the first embodiment, the sensing region is
on a first surface of the sensor body, and wherein the sensor body
includes a second surface opposite the first surfaces, and wherein
the second biointerface material is disposed on a substantial
portion of the first and second surfaces of the sensor body.
[0011] In a second embodiment, an analyte sensor for short-term and
long-term immobilization in a host's soft tissue is provided, the
sensor including: a short-term anchoring mechanism for providing
immobilization of the sensor in the soft tissue prior to
substantial formation of the foreign body capsule; and a long-term
anchoring mechanism for providing immobilization of the sensor in
the soft tissue during and after substantial formation of the
foreign body capsule.
[0012] In an aspect of the second embodiment, the short-term
anchoring mechanism includes a suture tab on the sensor body. In an
aspect of the second embodiment, the short-term anchoring mechanism
includes a suture. In an aspect of the second embodiment, the
short-term anchoring mechanism includes at least one of prongs,
spines, barbs, wings, and hooks. In an aspect of the second
embodiment, the short-term anchoring mechanism includes a geometric
configuration of the sensor body. In an aspect of the second
embodiment, the geometric configuration includes at least one of a
helical, tapered, and winged configuration.
[0013] In an aspect of the second embodiment, the long-term
anchoring mechanism includes an anchoring material disposed on the
sensor body. In an aspect of the second embodiment, the anchoring
material includes a fibrous material. In an aspect of the second
embodiment, the anchoring material includes a porous material. In
an aspect of the second embodiment, the anchoring material includes
a material with a surface topography. In an aspect of the second
embodiment, the long-term anchoring mechanism includes a surface
topography formed on an outer surface of the sensor body.
[0014] In a third embodiment, a method for immobilizing an analyte
sensor in soft tissue is provided, the method including: implanting
the analyte sensor in a host; anchoring the sensor in the host
prior to formation of a foreign body capsule for at least
short-term immobilization of the sensor within the soft tissue of
the host; and anchoring the sensor within the foreign body capsule
for long-term immobilization of the sensor within the soft tissue
of the host.
[0015] In an aspect of the third embodiment, the short-term
immobilization step includes suturing the sensor to the host's
tissue. In an aspect of the third embodiment, the suturing step
includes suturing the sensor such that the sensor is in
compression. In an aspect of the third embodiment, the short term
immobilization step includes utilizing at least one of prongs,
spines, barbs, wings, and hooks on the sensor to anchor the sensor
into the host's tissue upon implantation.
[0016] In an aspect of the third embodiment, the long-term
immobilization step includes disposing an anchoring material on the
sensor body that allows host tissue ingrowth into the material. In
an aspect of the third embodiment, the anchoring material includes
a fibrous material. In an aspect of the third embodiment, the
anchoring material includes a porous material. In an aspect of the
third embodiment, the anchoring material includes a material with a
surface topography. In an aspect of the third embodiment, the
long-term immobilization step includes utilizing a surface
topography formed on an outer surface of the sensor body that
allows tissue ingrowth into the surface of the sensor body.
[0017] In a fourth embodiment, a method for continuous measurement
of an analyte in a host is provided, the method including:
implanting an analyte sensor in a host; and measuring the
concentration of the analyte in the host before and after formation
of a foreign body capsule around the sensor.
[0018] In an aspect of the fourth embodiment, the analyte sensor is
wholly implanted into the host. In an aspect of the fourth
embodiment, the method further includes explanting the analyte
sensor from the host. In an aspect of the fourth embodiment, the
method further includes implanting another analyte sensor in the
host. In an aspect of the fourth embodiment, the another analyte
sensor is wholly implanted into the host.
[0019] In an aspect of the fourth embodiment, the method further
includes anchoring the analyte sensor in the host prior to
formation of a foreign body capsule for at least short-term
immobilization of the sensor within the soft tissue of the host. In
an aspect of the fourth embodiment, the method further includes
anchoring the sensor within the foreign body capsule for long-term
immobilization of the sensor within the soft tissue of the
host.
[0020] In a fifth embodiment, a method for implantation of an
analyte sensor in a host is provided, the method including: forming
a precisely dimensioned pocket in the subcutaneous space of the
host, wherein the pocket is dimensioned no greater than the
dimensions of the analyte sensor; inserting the analyte sensor into
the precisely-dimensioned pocket so as to minimize movement of the
sensor within the pocket.
[0021] In an aspect of the fifth embodiment, the step of forming a
pocket includes using a tool that allows precise dimensioning of
the pocket. In an aspect of the fifth embodiment, the tool includes
a head dimensioned substantially similar to the dimensions of the
analyte sensor and a handle for guiding the head into the pocket.
In an aspect of the fifth embodiment, the tool includes a head
dimensioned smaller than the dimensions of the analyte sensor and a
handle for guiding the head into the pocket.
[0022] In an aspect of the fifth embodiment, the method further
includes suturing the analyte sensor to the host tissue.
[0023] In an aspect of the fifth embodiment, the pocket is formed
adjacent the fascia of the host. In an aspect of the fifth
embodiment, the analyte sensor includes a sensing region for
measuring an analyte concentration, and wherein the analyte sensor
is placed within the pocket such that the sensing region is located
adjacent to the fascia.
[0024] In an aspect of the fifth embodiment, the method further
includes a step of forming a vertical incision prior to the step of
forming a pocket. In an aspect of the fifth embodiment, the method
further includes a step of forming a horizontal incision prior to
the step of forming a pocket. In an aspect of the fifth embodiment,
the pocket is formed in the abdominal region of the host.
[0025] In a sixth embodiment, an implantable analyte sensor for
measuring an analyte concentration in a host is provided, the
sensor including: a sensor body substantially formed from a water
vapor permeable material; and electrical components encapsulated
within the sensor body, wherein the electrical components include
RF circuitry and an antenna adapted for RF transmission from the
sensor in vivo to a receiver ex vivo, wherein the RF circuitry is
spaced a fixed distance from the sensor body so as to support a
dielectric constant that enables RF transmission between the sensor
in vivo to the receiver ex vivo.
[0026] In an aspect of the sixth embodiment, the fixed distance
includes a configuration that reduces water permeability therein.
In an aspect of the sixth embodiment, the configuration includes
conformal coating. In an aspect of the sixth embodiment, the
conformal coating includes Parylene.
[0027] In an aspect of the sixth embodiment, the configuration
includes epoxy. In an aspect of the sixth embodiment, the
configuration includes glass. In an aspect of the sixth embodiment,
the configuration includes one or more hermetic containers.
[0028] In a seventh embodiment, an analyte sensor for RF
transmission between the analyte sensor in vivo and a receiver ex
vivo is provided, the sensor including: a sensor body including RF
circuitry encapsulated within a substantially water vapor permeable
body that enables RF transmission therethrough; a sensing region
located on an outer surface of the sensor body for measuring an
analyte in soft tissue; a biointerface material disposed adjacent
to the sensing region that supports vascularized tissue ingrowth
for transport of the analyte to the sensing region; and an
anchoring material on a non-sensing outer surface of the sensor
body that supports tissue ingrowth for immobilization of the sensor
body in soft tissue.
[0029] In an aspect of the seventh embodiment, the sensor further
includes an antenna encapsulated within the sensor body. In an
aspect of the seventh embodiment, the sensor further includes a
power source encapsulated within the sensor body.
[0030] In an aspect of the seventh embodiment, the sensor body is
formed from plastic. In an aspect of the seventh embodiment, the
plastic includes epoxy.
[0031] In an aspect of the seventh embodiment, the sensor body is
molded around the RF circuitry. In an aspect of the seventh
embodiment, the sensor further includes an electrode system exposed
at the sensing region. In an aspect of the seventh embodiment, the
electrode system extends through the water vapor permeable body and
is operably connected to the RF circuitry.
[0032] In an aspect of the seventh embodiment, the biointerface
material includes a solid portion with a plurality of
interconnected cavities. In an aspect of the seventh embodiment,
the biointerface material further includes a domain proximal to the
sensing region that is impermeable to cells or cell processes and
is permeable to the passage of the analyte.
[0033] In an eighth embodiment, an electrochemical analyte sensor
for measuring an analyte concentration is provided, the sensor
including: a sensor body including electronic circuitry
encapsulated within the sensor body; and a plurality of electrodes
that extend from an outer surface of the sensor body to the
encapsulated electronic circuitry, wherein the electrodes are
mechanically and electrically connected and aligned to the
electronic circuitry prior to encapsulation within the sensor
body.
[0034] In an aspect of the eighth embodiment, the electrodes are
swaged to the electronic circuitry. In an aspect of the eighth
embodiment, the electrodes are welded using a technique selected
from the group consisting of spot welding, ultrasonic welding, and
laser welding.
[0035] In an aspect of the eighth embodiment, the electrodes and
electronic circuitry are encapsulated in the sensor body by a
molding process. In an aspect of the eighth embodiment, the sensor
body includes a water vapor permeable material.
[0036] In an aspect of the eighth embodiment, the electronic
circuitry is spaced from the water vapor permeable sensor body,
such that water vapor penetration within a fixed distance from the
electronic circuitry is inhibited. In an aspect of the eighth
embodiment, the electronic circuitry is spaced from the water vapor
permeable sensor body by epoxy. In an aspect of the eighth
embodiment, the electronic circuitry is spaced from the water vapor
permeable sensor body by a glass tube. In an aspect of the eighth
embodiment, the electronic circuitry is spaced from the water vapor
permeable sensor body by Parylene. In an aspect of the eighth
embodiment, the electronic circuitry is spaced from the water vapor
permeable sensor body by one or more hermetic containers.
[0037] In an aspect of the eighth embodiment, the sensor body
includes a substantially seamless exterior with the electrodes
extending through the sensor body to the outer surface thereof.
[0038] In a ninth embodiment, a method for manufacturing an
electrochemical analyte sensor is provided, the method including:
providing electronic circuitry designed to process signals from the
sensor; swaging a plurality of electrodes to the electronic
circuitry; and molding a plastic material around the electronic
circuitry to form the sensor body.
[0039] In a tenth embodiment, a method for manufacturing an analyte
sensor is provided, the method including: providing sensor
electronics designed to process signals from the sensor;
conformally coating the sensor electronics with a material that has
a first water permeability rate; and molding a water vapor
permeable material that has a second water permeability rate around
the coated sensor electronics to form a substantially seamless
sensor body, wherein the second water permeability rate is greater
than the first water penetration rate.
[0040] In an aspect of the tenth embodiment, the molding includes a
two-step molding process to form the substantially seamless sensor
body.
[0041] In an aspect of the tenth embodiment, the two-step molding
process includes: holding a first portion of the coated sensor
electronics and molding around a second portion of the coated
sensor electronics; and holding a portion of the cured sensor body
and molding around the first portion of the coated sensor
electronics.
[0042] In an eleventh embodiment, a method for manufacturing a
multilayer membrane for an analyte sensor is provided, the method
including: serially casting and subsequently curing each of a
plurality of layers to form the multilayer membrane onto a liner,
wherein the layers include a resistance layer for limiting the
passage of an analyte and an enzyme layer including an enzyme for
reacting with the analyte; and releasing the multilayer membrane
from the liner for application onto the analyte sensor.
[0043] In an aspect of the eleventh embodiment, the multilayer
membrane further includes an interference layer that substantially
prevents passage of potentially electrochemically interfering
substances.
[0044] In an aspect of the eleventh embodiment, the multilayer
membrane further includes an electrolyte layer including a hydrogel
for maintaining hydrophilicity at electrochemically reactive
surfaces of the analyte sensor.
[0045] In a twelfth embodiment, a method for casting a membrane
that regulates the transport of glucose, the method including:
forming a solvent solution including a polymer blend and a solvent,
wherein the polymer blend includes hydrophilic and hydrophobic
components; maintaining the solution at a first elevated
temperature for a predetermined time period in order to mix the
hydrophilic and hydrophobic components with each other and the
solvent, wherein the elevated temperature is above room
temperature; applying the composition to a liner to form a film
thereon; and curing the film, wherein the curing is accomplished
while ramping the temperature at a predetermined ramp rate to a
second elevated temperature that is above the first
temperature.
[0046] In an aspect of the twelfth embodiment, the first elevated
temperature is between about 60.degree. C. and about 100.degree. C.
In an aspect of the twelfth embodiment, the first elevated
temperature is about 80.degree. C.
[0047] In an aspect of the twelfth embodiment, the predetermined
time period is at least about 24 hours. In an aspect of the twelfth
embodiment, the predetermined time period is at least about 44
hours.
[0048] In an aspect of the twelfth embodiment, the predetermined
ramp rate is between about 3.degree. C. per minute and 12.degree.
C. per minute. In an aspect of the twelfth embodiment, the
predetermined ramp rate is about 7.degree. C. per minute.
[0049] In an aspect of the twelfth embodiment, the second elevated
temperature is at least about 100.degree. C.
[0050] In a thirteenth embodiment, a method for casting a membrane
for use with an electrochemical glucose sensor is provided, wherein
the membrane substantially prevents passage of potentially
electrochemically interfering substances, the method including:
forming a sufficiently diluted solvent solution including a polymer
and a solvent, wherein sufficiently diluted solvent solution
includes a ratio of polymer to solvent of about 1 to 10 wt. %
polymer to about 90 to 99 wt. % solvent; and applying the solvent
solution at a sufficiently fast casting speed that substantially
avoids film thickness inhomogeneities due to evaporation during
casting of the sufficiently diluted solvent solution.
[0051] In an aspect of the thirteenth embodiment, the membrane
limits the diffusion of hydrophilic species and large molecular
weight species.
[0052] In an aspect of the thirteenth embodiment, the membrane
includes a thickness between about 0.1 and 5 microns. In an aspect
of the thirteenth embodiment, the membrane includes a thickness
between about 0.5 and 3 microns.
[0053] In an aspect of the thirteenth embodiment, the polymer
includes polyurethane.
[0054] In an aspect of the thirteenth embodiment, the sufficiently
fast casting speed is between about 8 to about 15 inches/second. In
an aspect of the thirteenth embodiment, the sufficiently fast
casting speed is about 11.5 inches/second.
[0055] In a fourteenth embodiment, an implantable analyte sensor is
provided, the sensor including: a body including a material which
is permeable to water vapor, the body further including a sensing
region for measuring levels of an analyte; and a transmitter within
the body for transmitting the measurements obtained by the sensing
region, wherein at least a portion of the transmitter is spaced
from the body by a material adapted to reduce water from
penetrating therein.
[0056] In an aspect of the fourteenth embodiment, the transmitter
includes an oscillator and at least a portion of the oscillator is
spaced from the body by the material adapted to inhibit fluid from
penetrating therein. In an aspect of the fourteenth embodiment, the
oscillator includes an inductor and wherein the inductor is spaced
from the body by the material adapted to inhibit fluid from
penetrating therein. In an aspect of the fourteenth embodiment, the
oscillator includes a voltage controlled oscillator.
[0057] In a fifteenth embodiment, an implantable analyte sensor is
provided, including: electronics encapsulated within a water vapor
permeable body, wherein the electronics include a microprocessor
module and an RF module that has an RF transceiver with a
phase-locked loop, and wherein the microprocessor module is
programmed to initiate re-calibration of the PLL responsive to
detection of off-frequency shift.
[0058] In an aspect of the fifteenth embodiment, an electrochemical
glucose sensor including a three-electrode system, the sensor
including: an electrochemical cell including a working electrode,
reference electrode, and counter electrode; and a potentiostat that
controls the potential between the working and reference
electrodes, wherein an allowable range for the counter electrode
voltage is set sufficiently wide such that the glucose sensor can
react with other reducible species when oxygen becomes limited and
sufficiently narrow to ensure the circuitry does not allow
excessive current draw or bubble formation to occur.
[0059] In an aspect of the fifteenth embodiment, limiting the
current of at least one of the working or counter electrode
amplifiers to a preset current value configures the allowable
range. In an aspect of the fifteenth embodiment, setting the op-amp
to be offset from battery ground configures the allowable range. In
an aspect of the fifteenth embodiment, a reference voltage setting
between about +0.6V and +0.8V with respect to battery ground
configures the allowable range. In an aspect of the fifteenth
embodiment, a reference voltage setting of about +0.7V with respect
to battery ground configures the allowable range.
[0060] In a sixteenth embodiment, a method for manufacturing an
analyte sensor is provided, the method including: providing a
sensor body, wherein the sensor body includes a sensing region for
measuring the analyte; forming a multilayer membrane on a liner;
releasing the multilayer membrane from the liner and onto the
sensor body; and attaching the multilayer membrane to the analyte
sensor body proximal to the sensing region.
[0061] In an aspect of the sixteenth embodiment, the attaching step
includes a mechanical attachment. In an aspect of the sixteenth
embodiment, the mechanical attachment includes a metal or plastic
O-ring adapted to fit around a raised sensing region. In an aspect
of the sixteenth embodiment, the mechanical attachment includes a
metal or plastic disc adapted to be press-fit into the sensor body.
In an aspect of the sixteenth embodiment, the mechanical attachment
includes a metal or plastic clip adapted to be snap-fit into the
sensor body.
[0062] In a seventeenth embodiment, a method for manufacturing an
analyte sensor is provided, the method including: providing a
sensor body, wherein the sensor body includes a sensing region for
measuring the analyte; forming a multilayer membrane on a liner;
releasing the multilayer membrane from the liner and placing onto
the sensor body; and attaching the multilayer membrane to the
analyte sensor body proximal to the sensing region.
[0063] In an aspect of the seventeenth embodiment, the attaching
step includes a mechanical attachment. In an aspect of the
seventeenth embodiment, the mechanical attachment includes a metal
or plastic O-ring adapted to fit around a raised sensing region. In
an aspect of the seventeenth embodiment, the mechanical attachment
includes a metal or plastic disc adapted to be press-fit into the
sensor body. In an aspect of the seventeenth embodiment, the
mechanical attachment includes a metal or plastic clip adapted to
be snap-fit into the sensor body.
BRIEF DESCRIPTION OF THE DRAWINGS
[0064] FIG. 1A is a perspective view of a system of the preferred
embodiments, including a continuous analyte sensor implanted within
a human and a receiver for processing and displaying sensor
data.
[0065] FIG. 1B is a perspective view of the implantable analyte
sensor of the preferred embodiments.
[0066] FIG. 2 is a block diagram that illustrates the electronics
associated with the implantable glucose sensor in one
embodiment.
[0067] FIG. 3A is a top view of the electronics subassembly of the
analyte sensor, which shows the electrode system.
[0068] FIG. 3B is a side cross-sectional view of the electronics
subassembly taken through line 3-3 on FIG. 3A.
[0069] FIG. 3C is a circuit diagram of the potentiostat that
controls the three-electrode system of the preferred
embodiments.
[0070] FIG. 3D is an expanded cross-sectional view of a swaged
electrode in one embodiment.
[0071] FIG. 3E is a bottom view of the PCB, showing the side of the
PCB that faces the antenna board.
[0072] FIG. 3F is a cross-sectional view of an inductor that
controls the magnetic field, shown on a cut-away portion of the PCB
in one embodiment.
[0073] FIG. 3G is a cross-sectional view of an inductor that
controls the magnetic field, shown on a cut-away portion of the PCB
in an alternative embodiment.
[0074] FIG. 4A is a perspective view of the primary mold which is
used in the primary casting in one embodiment.
[0075] FIG. 4B is a cross-sectional side view of the electronics
subassembly during the primary casting process.
[0076] FIG. 4C is a perspective view of the primary potted device
after the primary casting process.
[0077] FIG. 5A is a perspective view of inserting the primary
potted device into a secondary mold.
[0078] FIG. 5B is a perspective view of the secondary potted device
after the secondary casting process.
[0079] FIG. 6A is a perspective view of an analyte sensor in one
embodiment, including a thin substantially oval geometry, a curved
sensing region, and an overall curved surface on which the sensing
region is located, thereby causing contractile forces from the
foreign body capsule to press downward on the sensing region.
[0080] FIG. 6B is an end view of the analyte sensor of FIG. 6A
showing the contractile forces that would be caused by a foreign
body capsule.
[0081] FIG. 6C is a side view of the analyte sensor of FIG. 6A.
[0082] FIG. 7A is an illustration that represents a method of
forming the sensing membrane of the preferred embodiments.
[0083] FIG. 7B is a schematic side view of the sensing membrane in
one embodiment.
[0084] FIG. 8A is a cross-sectional schematic view of a
biointerface membrane in vivo in one embodiment, wherein the
membrane comprises a cell disruptive domain and cell impermeable
domain.
[0085] FIG. 8B is an illustration of the membrane of FIG. 8A,
showing contractile forces caused by the fibrous tissue of the
FBR.
[0086] FIG. 9A is an exploded perspective view of the analyte
sensor prior to membrane attachment.
[0087] FIG. 9B is a perspective view of the analyte sensor after
membrane systems and methods.
[0088] FIGS. 9C to 9H are exploded and collapsed perspective views
of a variety of alternative embodiments that utilize alternative
membrane attachment techniques.
[0089] FIG. 10A is an exploded perspective view of a sensor,
illustrating the tissue-facing components of the sensor prior to
attachment.
[0090] FIG. 10B is a perspective view of the assembled analyte
sensor, including the tissue-facing components attached
thereto.
[0091] FIG. 10C is a perspective view of the non-sensing side of
the assembled analyte sensor, showing a short-term anchoring device
in one embodiment.
[0092] FIG. 11A is a perspective view of a sizing tool in one
embodiment, including a head and a handle.
[0093] FIG. 11B is a side view of the sizing tool of FIG. 11A
showing the offset placement of the handle on the head in some
embodiments.
[0094] FIG. 11C is a top view of the sizing tool of FIG. 11A
showing a curvature substantially similar to that of the sensor
body.
[0095] FIG. 12A is a perspective view of an abdominal region of a
human, showing an exploded view of the incision and sizing tool
that may be used for surgical implantation.
[0096] FIG. 12B is a perspective view of a portion of the abdominal
region of a human, showing an exploded view of sensor insertion
into the precisely formed pocket.
[0097] FIG. 13A is a schematic view of the sensor after insertion
into the pocket with the sensing region positioned adjacent to the
fascia.
[0098] FIG. 13B is a side view of an analyte sensor with long-term
anchoring component in the form of an anchoring material on both
sides of the sensor.
DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENT
[0099] The following description and examples illustrate some
exemplary embodiments of the disclosed invention in detail. Those
of skill in the art will recognize that there are numerous
variations and modifications of this invention that are encompassed
by its scope. Accordingly, the description of a certain exemplary
embodiment should not be deemed to limit the scope of the present
invention.
DEFINITIONS
[0100] In order to facilitate an understanding of the disclosed
invention, a number of terms are defined below.
[0101] The term "ROM," as used herein, is a broad term and is used
in its ordinary sense, including, without limitation, read-only
memory. The term is inclusive of various types of ROM, including
EEPROM, rewritable ROMs, flash memory, or the like.
[0102] The term "RAM," as used herein, is a broad term and is used
in its ordinary sense, including, without limitation, random access
memory. The term is inclusive of various types of RAM, including
dynamic-RAM, static-RAM, non-static RAM, or the like.
[0103] The term "A/D Converter," as used herein, is a broad term
and is used in its ordinary sense, including, without limitation,
hardware and/or software that converts analog electrical signals
into corresponding digital signals.
[0104] The term "microprocessor," as used herein, is a broad term
and is used in its ordinary sense, including, without limitation a
computer system or processor designed to perform arithmetic and
logic operations using logic circuitry that responds to and
processes the basic instructions that drive a computer.
[0105] The term "RF transceiver," as used herein, is a broad term
and is used in its ordinary sense, including, without limitation, a
radio frequency transmitter and/or receiver for transmitting and/or
receiving signals.
[0106] The terms "raw data stream" and "data stream," as used
herein, are broad terms and are used in their ordinary sense,
including, without limitation, an analog or digital signal directly
related to the measured glucose from the glucose sensor. In one
example, the raw data stream is digital data in "counts" converted
by an A/D converter from an analog signal (for example, voltage or
amps) representative of a glucose concentration. The terms broadly
encompass a plurality of time spaced data points from a
substantially continuous glucose sensor, which comprises individual
measurements taken at time intervals ranging from fractions of a
second up to, for example, 1, 2, or 5 minutes or longer.
[0107] The term "counts," as used herein, is a broad term and is
used in its ordinary sense, including, without limitation, a unit
of measurement of a digital signal. In one example, a raw data
stream measured in counts is directly related to a voltage (for
example, converted by an A/D converter), which is directly related
to current from the working electrode. In another example, counter
electrode voltage measured in counts is directly related to a
voltage.
[0108] The term "potentiostat," as used herein, is a broad term and
is used in its ordinary sense, including, without limitation, an
electrical system that controls the potential between the working
and reference electrodes of a three-electrode cell at a preset
value. It forces whatever current is necessary to flow between the
working and counter electrodes to keep the desired potential, as
long as the needed cell voltage and current do not exceed the
compliance limits of the potentiostat.
[0109] The term "electrical potential," as used herein, is a broad
term and is used in its ordinary sense, including, without
limitation, the electrical potential difference between two points
in a circuit which is the cause of the flow of a current.
[0110] The term "physiologically feasible," as used herein, is a
broad term and is used in its ordinary sense, including, without
limitation, the physiological parameters obtained from continuous
studies of glucose data in humans and/or animals. For example, a
maximal sustained rate of change of glucose in humans of about 4 to
5 mg/dL/min and a maximum acceleration of the rate of change of
about 0.1 to 0.2 mg/dL/min/min are deemed physiologically feasible
limits. Values outside of these limits would be considered
non-physiological and likely a result of signal error, for example.
As another example, the rate of change of glucose is lowest at the
maxima and minima of the daily glucose range, which are the areas
of greatest risk in patient treatment, thus a physiologically
feasible rate of change can be set at the maxima and minima based
on continuous studies of glucose data.
[0111] The term "ischemia," as used herein, is a broad term and is
used in its ordinary sense, including, without limitation, local
and temporary deficiency of blood supply due to obstruction of
circulation to a part (for example, sensor). Ischemia can be caused
by mechanical obstruction (for example, arterial narrowing or
disruption) of the blood supply, for example.
[0112] The term "system noise," as used herein, is a broad term and
is used in its ordinary sense, including, without limitation,
unwanted electronic or diffusion-related noise which can include
Gaussian, motion-related, flicker, kinetic, or other white noise,
for example.
[0113] The term "biointerface membrane" as used herein is a broad
term and is used in its ordinary sense, including, without
limitation, a permeable membrane that functions as a device-tissue
interface comprised of two or more domains. In some embodiments,
the biointerface membrane is composed of two domains. The first
domain supports tissue ingrowth, interferes with barrier cell layer
formation, and includes an open cell configuration having cavities
and a solid portion. The second domain is resistant to cellular
attachment and impermeable to cells (for example, macrophages). The
biointerface membrane is made of biostable materials and can be
constructed in layers, uniform or non-uniform gradients (i.e.,
anisotropic), or in a uniform or non-uniform cavity size
configuration.
[0114] The term "sensing membrane," as used herein, is a broad term
and is used in its ordinary sense, including, without limitation, a
permeable or semi-permeable membrane that can be comprised of two
or more domains and is typically constructed of materials of a few
microns thickness or more, which are permeable to oxygen and may or
may not be permeable to glucose. In one example, the sensing
membrane comprises an enzyme, for example immobilized glucose
oxidase enzyme, which enables an electrochemical reaction to occur
to measure a concentration of analyte.
[0115] The term "domain" as used herein is a broad term and is used
in its ordinary sense, including, without limitation, regions of a
membrane that can be layers, uniform or non-uniform gradients (for
example, anisotropic) or provided as portions of the membrane. The
term is broad enough to include one or more functions one or more
(combined) domains, or a plurality of layers or regions that each
provide one or more of the functions of each of the various
domains.
[0116] The term "barrier cell layer" as used herein is a broad term
and is used in its ordinary sense, including, without limitation, a
cohesive monolayer of cells (for example, macrophages and foreign
body giant cells) that substantially block the transport of at
least some molecules across the second domain and/or membrane.
[0117] The term "cellular attachment," as used herein is a broad
term and is used in its ordinary sense, including, without
limitation, adhesion of cells and/or mechanical attachment of cell
processes to a material at the molecular level, and/or attachment
of cells and/or cell processes to micro- (or macro-) porous
material surfaces. One example of a material used in the prior art
that allows cellular attachment due to porous surfaces is the
BIOPORE.TM. cell culture support marketed by Millipore (Bedford,
Mass.) (see Brauker '330, supra).
[0118] The phrase "distal to" as used herein is a broad term and is
used in its ordinary sense, including, without limitation, the
spatial relationship between various elements in comparison to a
particular point of reference. For example, some embodiments of a
device include a biointerface membrane having a cell disruptive
domain and a cell impermeable domain. If the sensor is deemed to be
the point of reference and the cell disruptive domain is positioned
farther from the sensor, then that domain is distal to the
sensor.
[0119] The term "proximal to" as used herein is a broad term and is
used in its ordinary sense, including, without limitation, the
spatial relationship between various elements in comparison to a
particular point of reference. For example, some embodiments of a
device include a biointerface membrane having a cell disruptive
domain and a cell impermeable domain. If the sensor is deemed to be
the point of reference and the cell impermeable domain is
positioned nearer to the sensor, then that domain is proximal to
the sensor.
[0120] The term "cell processes" as used herein is a broad term and
is used in its ordinary sense, including, without limitation,
pseudopodia of a cell.
[0121] The term "solid portions" as used herein is a broad term and
is used in its ordinary sense, including, without limitation, a
solid material having a mechanical structure that demarcates the
cavities, voids, or other non-solid portions.
[0122] The term "substantial" as used herein is a broad term and is
used in its ordinary sense, including, without limitation, a
sufficient amount that provides a desired function. For example, in
the micro-architecture of the preferred embodiments, a substantial
number of cavities have a size that allows a substantial number of
inflammatory cells to enter therein, which may include an amount
greater than 50 percent, an amount greater than 60 percent, an
amount greater than 70 percent, an amount greater than 80 percent,
and an amount greater than 90 percent of cavities within a
preferred nominal pore size range.
[0123] The term "co-continuous" as used herein is a broad term and
is used in its ordinary sense, including, without limitation, a
solid portion wherein an unbroken curved line in three dimensions
exists between any two points of the solid portion.
[0124] The term "biostable" as used herein is a broad term and is
used in its ordinary sense, including, without limitation,
materials that are relatively resistant to degradation by processes
that are encountered in vivo.
[0125] The term "analyte" as used herein is a broad term and is
used in its ordinary sense, including, without limitation, to refer
to a substance or chemical constituent in a biological fluid (for
example, blood, interstitial fluid, cerebral spinal fluid, lymph
fluid or urine) that can be analyzed. Analytes can include
naturally occurring substances, artificial substances, metabolites,
and/or reaction products. In some embodiments, the analyte for
measurement by the sensing regions, devices, and methods is
glucose. However, other analytes are contemplated as well,
including but not limited to acarboxyprothrombin; acylcarnitine;
adenine phosphoribosyl transferase; adenosine deaminase; albumin;
alpha-fetoprotein; amino acid profiles (arginine (Krebs cycle),
histidine/urocanic acid, homocysteine, phenylalanine/tyrosine,
tryptophan); andrenostenedione; antipyrine; arabinitol enantiomers;
arginase; benzoylecgonine (cocaine); biotinidase; biopterin;
c-reactive protein; carnitine; carnosinase; CD4; ceruloplasmin;
chenodeoxycholic acid; chloroquine; cholesterol; cholinesterase;
conjugated 1-.beta. hydroxy-cholic acid; cortisol; creatine kinase;
creatine kinase MM isoenzyme; cyclosporin A; d-penicillamine;
de-ethylchloroquine; dehydroepiandrosterone sulfate; DNA
(acetylator polymorphism, alcohol dehydrogenase, alpha
1-antitrypsin, cystic fibrosis, Duchenne/Becker muscular dystrophy,
glucose-6-phosphate dehydrogenase, hemoglobinopathies, A, S, C, E,
D-Punjab, beta-thalassemia, hepatitis B virus, HCMV, HIV-1, HTLV-1,
Leber hereditary optic neuropathy, MCAD, RNA, PKU, Plasmodium
vivax, sexual differentiation, 21-deoxycortisol);
desbutylhalofantrine; dihydropteridine reductase; diptheria/tetanus
antitoxin; erythrocyte arginase; erythrocyte protoporphyrin;
esterase D; fatty acids/acylglycines; free .beta.-human chorionic
gonadotropin; free erythrocyte porphyrin; free thyroxine (FT4);
free tri-iodothyronine (FT3); fumarylacetoacetase;
galactose/gal-1-phosphate; galactose-1-phosphate uridyltransferase;
gentamicin; glucose-6-phosphate dehydrogenase; glutathione;
glutathione perioxidase; glycocholic acid; glycosylated hemoglobin;
halofantrine; hemoglobin variants; hexosaminidase A; human
erythrocyte carbonic anhydrase I; 17 alpha-hydroxyprogesterone;
hypoxanthine phosphoribosyl transferase; immunoreactive trypsin;
lactate; lead; lipoproteins ((a), B/A-1, .beta.); lysozyme;
mefloquine; netilmicin; phenobarbitone; phenytoin;
phytanic/pristanic acid; progesterone; prolactin; prolidase; purine
nucleoside phosphorylase; quinine; reverse tri-iodothyronine (rT3);
selenium; serum pancreatic lipase; sissomicin; somatomedin C;
specific antibodies (adenovirus, anti-nuclear antibody, anti-zeta
antibody, arbovirus, Aujeszky's disease virus, dengue virus,
Dracunculus medinensis, Echinococcus granulosus, Entamoeba
histolytica, enterovirus, Giardia duodenalisa, Helicobacter pylori,
hepatitis B virus, herpes virus, HIV-1, IgE (atopic disease),
influenza virus, Leishmania donovani, leptospira,
measles/mumps/rubella, Mycobacterium leprae, Mycoplasma pneumoniae,
Myoglobin, Onchocerca volvulus, parainfluenza virus, Plasmodium
falciparum, poliovirus, Pseudomonas aeruginosa, respiratory
syncytial virus, rickettsia (scrub typhus), Schistosoma mansoni,
Toxoplasma gondii, Trepenoma pallidium, Trypanosoma cruzi/rangeli,
vesicular stomatis virus, Wuchereria bancrofti, yellow fever
virus); specific antigens (hepatitis B virus, HIV-1);
succinylacetone; sulfadoxine; theophylline; thyrotropin (TSH);
thyroxine (T4); thyroxine-binding globulin; trace elements;
transferrin; UDP-galactose-4-epimerase; urea; uroporphyrinogen I
synthase; vitamin A; white blood cells; and zinc protoporphyrin.
Salts, sugar, protein, fat, vitamins and hormones naturally
occurring in blood or interstitial fluids can also constitute
analytes in certain embodiments. The analyte can be naturally
present in the biological fluid, for example, a metabolic product,
a hormone, an antigen, an antibody, and the like. Alternatively,
the analyte can be introduced into the body, for example, a
contrast agent for imaging, a radioisotope, a chemical agent, a
fluorocarbon-based synthetic blood, or a drug or pharmaceutical
composition, including but not limited to insulin; ethanol;
cannabis (marijuana, tetrahydrocannabinol, hashish); inhalants
(nitrous oxide, amyl nitrite, butyl nitrite, chlorohydrocarbons,
hydrocarbons); cocaine (crack cocaine); stimulants (amphetamines,
methamphetamines, Ritalin, Cylert, Preludin, Didrex, PreState,
Voranil, Sandrex, Plegine); depressants (barbituates, methaqualone,
tranquilizers such as Valium, Librium, Miltown, Serax, Equanil,
Tranxene); hallucinogens (phencyclidine, lysergic acid, mescaline,
peyote, psilocybin); narcotics (heroin, codeine, morphine, opium,
meperidine, Percocet, Percodan, Tussionex, Fentanyl, Darvon,
Talwin, Lomotil); designer drugs (analogs of fentanyl, meperidine,
amphetamines, methamphetamines, and phencyclidine, for example,
Ecstasy); anabolic steroids; and nicotine. The metabolic products
of drugs and pharmaceutical compositions are also contemplated
analytes. Analytes such as neurochemicals and other chemicals
generated within the body can also be analyzed, such as, for
example, ascorbic acid, uric acid, dopamine, noradrenaline,
3-methoxytyramine (3MT), 3,4-dihydroxyphenylacetic acid (DOPAC),
homovanillic acid (HVA), 5-hydroxytryptamine (5HT), and
5-hydroxyindoleacetic acid (FHIAA).
[0126] The terms "operably connected" and "operably linked" as used
herein are broad terms and are used in their ordinary sense,
including, without limitation, one or more components being linked
to another component(s) in a manner that allows transmission of
signals between the components. For example, one or more electrodes
can be used to detect the amount of analyte in a sample and convert
that information into a signal; the signal can then be transmitted
to a circuit. In this case, the electrode is "operably linked" to
the electronic circuitry.
[0127] The term "host" as used herein is a broad term and is used
in its ordinary sense, including, without limitation, mammals,
particularly humans.
[0128] The phrase "continuous (or continual) analyte sensing" as
used herein is a broad term and is used in its ordinary sense,
including, without limitation, the period in which monitoring of
analyte concentration is continuously, continually, and or
intermittently (regularly or irregularly) performed, for example,
about every 5 to 10 minutes.
[0129] The term "sensing region" as used herein is a broad term and
is used in its ordinary sense, including, without limitation, the
region of a monitoring device responsible for the detection of a
particular analyte. The sensing region generally comprises a
non-conductive body, a working electrode (anode), a reference
electrode and a counter electrode (cathode) passing through and
secured within the body forming an electrochemically reactive
surface at one location on the body and an electronic connective
means at another location on the body, and a multi-region membrane
affixed to the body and covering the electrochemically reactive
surface. The counter electrode has a greater electrochemically
reactive surface area than the working electrode. During general
operation of the sensor a biological sample (for example, blood or
interstitial fluid) or a portion thereof contacts (directly or
after passage through one or more membranes or domains) an enzyme
(for example, glucose oxidase); the reaction of the biological
sample (or portion thereof) results in the formation of reaction
products that allow a determination of the analyte (for example,
glucose) level in the biological sample. In some embodiments, the
multi-region membrane further comprises an enzyme domain (for
example, and enzyme layer), and an electrolyte phase (i.e., a
free-flowing liquid phase comprising an electrolyte-containing
fluid described further below).
[0130] The term "electrochemically reactive surface" as used herein
is a broad term and is used in its ordinary sense, including,
without limitation, the surface of an electrode where an
electrochemical reaction takes place. In the case of the working
electrode, the hydrogen peroxide produced by the enzyme catalyzed
reaction of the analyte being detected reacts creating a measurable
electric current (for example, detection of glucose analyte
utilizing glucose oxidase produces H.sub.2O.sub.2 peroxide as a by
product, H.sub.2O.sub.2 reacts with the surface of the working
electrode producing two protons (2H.sup.+), two electrons
(2e.sup.-) and one molecule of oxygen (O.sub.2) which produces the
electronic current being detected). In the case of the counter
electrode, a reducible species, for example, O.sub.2 is reduced at
the electrode surface in order to balance the current being
generated by the working electrode.
[0131] The term "electronic connection" as used herein is a broad
term and is used in its ordinary sense, including, without
limitation, any electronic connection known to those in the art
that can be utilized to interface the sensing region electrodes
with the electronic circuitry of a device such as mechanical (for
example, pin and socket) or soldered.
[0132] The term "oxygen antenna domain" as used herein is a broad
term and is used in its ordinary sense, including, without
limitation, a domain composed of a material that has higher oxygen
solubility than aqueous media so that it concentrates oxygen from
the biological fluid surrounding the biointerface membrane. The
domain can then act as an oxygen reservoir during times of minimal
oxygen need and has the capacity to provide on demand a higher
oxygen gradient to facilitate oxygen transport across the membrane.
This enhances function in the enzyme reaction domain and at the
counter electrode surface when glucose conversion to hydrogen
peroxide in the enzyme domain consumes oxygen from the surrounding
domains. Thus, this ability of the oxygen antenna domain to apply a
higher flux of oxygen to critical domains when needed improves
overall sensor function.
[0133] The term "casting" as used herein is a broad term and is
used in its ordinary sense, including, without limitation, a
process where a fluid material is applied to a surface or surfaces
and allowed to cure. The term is broad enough to encompass a
variety of coating techniques, for example, using a draw-down
machine, dip coating, or the like.
[0134] The term "water vapor permeable" as used herein is a broad
term and is used in its ordinary sense, including, without
limitation, characterized by permitting water vapor to permeate
therethrough.
[0135] The following abbreviations apply herein: 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); .mu.m (micrometers); nm (nanometers); h and hr
(hours); min. (minutes); s and sec. (seconds); and .degree. C.
(degrees Centigrade).
Overview
[0136] FIG. 1A is a perspective view of a system of the preferred
embodiments, including a continuous analyte sensor 12 implanted
within a human 10 and a receiver 14 for processing and displaying
sensor data. The system of the preferred embodiments provides
improved convenience and accuracy because of its discrete design
that enables acceptance into a host's tissue with minimum invasive
trauma, while providing reliable wireless transmissions through the
physiological environment, and thereby increases overall patient
comfort, confidence, safety, and convenience.
[0137] The continuous analyte sensor 12 measures a concentration of
an analyte or a substance indicative of the concentration or
presence of the analyte. Although some of the following description
is drawn to a glucose sensor, the analyte sensor 12 may be any
sensor capable of determining the level of any analyte in the body,
for example oxygen, lactase, insulin, hormones, cholesterol,
medicaments, viruses, or the like. Additionally, although much of
the description of the analyte sensor is focused on electrochemical
detection methods, the systems and methods may be applied to
analyte sensors that utilize other measurement techniques,
including enzymatic, chemical, physical, spectrophotometric,
polarimetric, calorimetric, radiometric, or the like.
[0138] FIGS. 2 to 13 describe the systems and methods associated
with the manufacture, configuration, and implantation of the
analyte sensor of the preferred embodiments. FIG. 2 describes
systems and methods for measuring an analyte concentration and
providing an output signal indicative of the concentration of the
analyte, in one embodiment. This output signal is typically a raw
data stream that is used to provide a useful value of the measured
analyte concentration to a patient or doctor, for example.
Accordingly, a receiver 14 is provided that receives and processes
the raw data stream, including calibrating, validating, and
displaying meaningful glucose values to a patient, such as
described in co-pending U.S. patent application Ser. No.
10/633,367, which is incorporated herein by reference in its
entirety.
[0139] Reference is now made to FIG. 1B, which is a perspective
view of the implantable analyte sensor 12 of the preferred
embodiments. In this embodiment, a sensing region 16 and a
non-sensing region 18 are shown on the analyte sensor 12.
Electronics associated with the analyte sensor 12 are described in
more detail below with reference to FIGS. 2 and 3. The sensor
electronics are preferably encapsulated within a molded plastic
body, for example, a thermoset, such as described in more detail
below with reference to FIGS. 4 and 5. The sensor electronics
include an electrode system that extends through the molded body
and is exposed at the sensing region 16, such as described in more
detail below with reference to FIG. 3. A sensing membrane covers
the exposed electrodes, which is described in more detail below
with reference to FIG. 7. A biointerface membrane, which covers the
sensing membrane, is configured to support tissue ingrowth, disrupt
contractile forces typically found in a foreign body response,
encourage vascularity, and interfere with barrier cell layer
formation, and is described in more detail below with reference to
FIG. 8. An anchoring material covers at least a portion of the
non-sensing region of the analyte sensor 12 for long-term anchoring
of the sensor to the host tissue, which is described in more detail
below with reference to FIGS. 10 and 13. A suture strip or other
component for short-term anchoring of the sensor to the host tissue
may optionally be provided. In some embodiments, the component for
short-term anchoring of the sensor to the host tissue may be on a
portion of the non-sensing region, as described in more detail
below with reference to FIGS. 10 and 12.
[0140] In one preferred embodiment, the analyte sensor is a glucose
sensor, wherein the sensing region 16 comprises electrode system
including a platinum working electrode, a platinum counter
electrode, and a silver/silver chloride reference electrode, for
example. However a variety of electrode materials and
configurations may be used with the implantable analyte sensor of
the preferred embodiments. The top ends of the electrodes are in
contact with an electrolyte phase (not shown), which is a
free-flowing fluid phase disposed between a sensing membrane and
the electrodes. In one embodiment, the counter electrode is
provided to balance the current generated by the species being
measured at the working electrode. In some embodiments, the sensing
membrane includes an enzyme, for example, glucose oxidase, and
covers the electrolyte phase. In the case of a glucose oxidase
based glucose sensor, the species being measured at the working
electrode is H.sub.2O.sub.2. Glucose oxidase catalyzes the
conversion of oxygen and glucose to hydrogen peroxide and gluconate
according to the following reaction:
Glucose+O.sub.2.fwdarw.Gluconate+H.sub.2O.sub.2
[0141] The change in H.sub.2O.sub.2 can be monitored to determine
glucose concentration because for each glucose molecule
metabolized, there is a proportional change in the product
H.sub.2O.sub.2. Oxidation of H.sub.2O.sub.2 by the working
electrode is balanced by reduction of ambient oxygen, enzyme
generated H.sub.2O.sub.2, or other reducible species at the counter
electrode. The H.sub.2O.sub.2 produced from the glucose oxidase
reaction further reacts at the surface of working electrode and
produces two protons (2H.sup.+), two electrons (2e.sup.-), and one
oxygen molecule (O.sub.2).
[0142] A potentiostat (FIG. 3C) is employed to monitor the
electrochemical reaction at the electroactive surface(s). The
potentiostat applies a constant potential to the working and
reference electrodes to determine a current value. The current that
is produced at the working electrode (and flows through the
circuitry to the counter electrode) is substantially proportional
to the amount of H.sub.2O.sub.2 that diffuses to the working
electrode. Accordingly, a raw signal can be produced that is
representative of the concentration of glucose in the user's body,
and therefore can be utilized to estimate a meaningful glucose
value.
Description
[0143] Sensor Electronics
[0144] FIG. 2 is a block diagram that illustrates the electronics
22 associated with the implantable glucose sensor 12 in one
embodiment. In this embodiment, a potentiostat 24 is shown, which
is operably connected to an electrode system (such as described
above) to obtain a current value, and includes a resistor (not
shown) that translates the current into voltage. An A/D converter
26 digitizes the analog signal into "counts" for processing.
Accordingly, the resulting raw data stream in counts is directly
related to the current measured by the potentiostat 24.
[0145] A microprocessor module 28 includes the central control unit
that houses ROM 30 and RAM 32 and controls the processing of the
sensor electronics 22. It is noted that certain alternative
embodiments can utilize a computer system other than a
microprocessor to process data as described herein. In some
alternative embodiments, an application-specific integrated circuit
(ASIC) can be used for some or all the sensor's central processing
as is appreciated by one skilled in the art. The ROM 30 provides
semi-permanent storage of data, for example, storing data such as
sensor identifier (ID) and programming to process data streams (for
example, programming for data smoothing and/or replacement of
signal artifacts such as described in copending U.S. patent
application Ser. No. 10/648,849 and entitled, "SYSTEMS AND METHODS
FOR REPLACING SIGNAL ARTIFACTS IN A GLUCOSE SENSOR DATA STREAM,"
filed Aug. 22, 2003, which is incorporated herein by reference in
its entirety). The RAM 32 can be used for the system's cache
memory, for example for temporarily storing recent sensor data. In
some alternative embodiments, memory storage components comparable
to ROM 30 and RAM 32 may be used instead of or in addition to the
preferred hardware, such as dynamic-RAM, static-RAM, non-static
RAM, EEPROM, rewritable ROMs, flash memory, or the like.
[0146] A battery 34 is operably connected to the sensor electronics
22 and provides the necessary power for the sensor 12. In one
embodiment, the battery is a Lithium Manganese Dioxide battery,
however any appropriately sized and powered battery can be used
(for example, AAA, Nickel-cadmium, Zinc-carbon, Alkaline, Lithium,
Nickel-metal hydride, Lithium-ion, Zinc-air, Zinc-mercury oxide,
Silver-zinc, and/or hermetically-sealed). In some embodiments the
battery is rechargeable. In some embodiments, a plurality of
batteries can be used to power the system. In yet other
embodiments, the sensor can be transcutaneously powered via an
inductive coupling, for example. In some embodiments, a quartz
crystal 36 is operably connected to the microprocessor 28 and
maintains system time for the computer system as a whole.
[0147] An RF module 38 is operably connected to the microprocessor
28 and transmits the sensor data from the sensor 12 to a receiver
14 within a wireless transmission 40 via antenna 42. In some
embodiments, a second quartz crystal 44 provides the system time
for synchronizing the data transmissions from the RF transceiver.
The RF transceiver generally includes a register and a phase-locked
loop (PLL) with an oscillator, phase discriminator (PD), loop
filter (LPF), and a voltage-controlled oscillator (VCO) as is
appreciated by one skilled in the art. In some alternative
embodiments, however, other mechanisms such as optical, infrared
radiation (IR), ultrasonic, or the like may be used to transmit
and/or receive data.
[0148] It is noted that the preferred embodiments advantageously
encapsulate the electronics in a water vapor permeable material,
such as described in more detail with reference to FIGS. 4 to 5,
below. It has been observed, however, that a change of the
electrical properties of the water permeable sensor body
contributes to a changing dielectric loading of the RF, causing
shifting of the carrier frequency that may prohibit a receiver (for
example, listening in a particular frequency range) from receiving
the data transmissions. Although conventional PLL's are designed to
run a standard calibration to compensate for dielectric loading
changes over time, it has been seen that certain situations occur
in an implantable water vapor permeable sensor, wherein the water
penetration rate increases more quickly than can be compensated for
in a standard calibration cycle. Accordingly, the preferred
embodiments are programmed to monitor the PLL to determine when the
carrier frequency has shifted outside of the predetermined range
and to run a re-calibration cycle upon an indication of
"off-frequency." While not wishing to be bound by theory, it is
believed that re-calibrating the PLL responsive to detection of
off-frequency reduces or eliminates missing data transmissions that
result from a shifting carrier frequency in an implantable water
vapor permeable sensor.
[0149] Electronics Subassembly
[0150] FIGS. 3A and 3B are top (FIG. 3A) and side cross-sectional
(FIG. 3B) views of the electronic subassembly 46 associated with
the sensor 12 in one exemplary embodiment, which includes the
hardware and software that provides for the functionality of sensor
electronics as described above. Particularly, FIGS. 3A and 3B
illustrate the electronics subassembly 46 prior to encapsulation in
a moldable plastic material.
[0151] The electronics subassembly 46 generally includes hardware
and software designed to support the functions described above;
additionally, the electronics subassembly of the preferred
embodiments is configured to accommodate certain preferred design
parameters described herein, which facilitate analyte sensor
immobilization within the subcutaneous pocket. Immobilization of
the sensor within the host tissue is advantageous because motion
(for example, acute and/or chronic movement of the sensor in the
host tissue) has been found to produce acute and/or chronic
inflammation, which has been shown to result in poor short-term
and/or long-term sensor performance. For example, during an
experiment wherein larger, bulkier versions of the analyte sensor
were implanted in humans for an average of 44 days +/-14 days [See
Garg, S.; Schwartz, S.; Edelman, S. "Improved Glucose Excursions
Using an Implantable Real-Time Continuous Glucose Sensor in Adults
with Type 1Diabetes." Diabetes Care 2004, 27, 734-738], it was
discovered that movement of the sensor resulted in thicker foreign
body capsule formation, which correlated with decreased sensor
performance. While not wishing to be bound by theory, it is
believed that size optimization (for example, miniaturization) of
the analyte sensor enables more discrete and secure implantation,
and is believed to reduce macro-motion of the sensor induced by the
patient and micro-motion caused by movement of the sensor within
the subcutaneous pocket, and thereby improve sensor
performance.
[0152] Additionally, in contrast to devices made from hermetic
materials, the preferred embodiments of the present invention
advantageously encapsulate the electronics in a material that is
water vapor permeable. The use of a water vapor permeable material,
for example moldable plastic, is advantageous for a variety of
reasons described elsewhere herein, for example, ease of design
changes, security and alignment of the electronics during and after
the molding process, and the ability to machine the device with
precise curvatures. In some embodiments, the electronic subassembly
possesses features that maintain the frequency of the voltage
controlled oscillator (VCO) if water vapor penetrates into the
water vapor permeable sensor body. For example, in some
embodiments, the electronics subassembly reduces changes in the
inductor parameters, which may otherwise occur as a result of water
vapor within the electromagnetic field of the inductor (for
example, which may result in a shift in the carrier frequency of
the VCO). Such field effects may cause the VCO to transmit off of
its tuned carrier frequency, which degrades the RF telemetry
capabilities of the sensor. Furthermore, the carrier frequency of
the sensor of the preferred embodiments is optimized for sensor
longevity.
[0153] Additionally, the analyte sensor of the preferred
embodiments supports a high frequency, low power operation, which
supports miniaturization of the analyte sensor with optimized
functionality. Accordingly, the design of the electronics
subassembly 46 of the preferred embodiments provides a discrete,
efficient configuration, while maintaining long-term power supply
and functional RF telemetry within the water vapor permeable body
in vivo.
[0154] FIGS. 3A and 3B generally show the electronics subassembly
46 of the preferred embodiments including a printed circuit board
(PCB) 48, an antenna board 50, a battery 34, and a plurality of
interconnections therebetween, which are configured to provide the
above described functionality, as is appreciated by one skilled in
the art. FIG. 3A is a top view of the electronics subassembly of
the analyte sensor 12, which shows the electrode system 54
described in the Overview section, above. FIG. 3B is a side
cross-sectional view of the electronics subassembly taken from line
3-3 on FIG. 3A.
[0155] The PCB 48 supports the components, for example
microprocessor module 28, including ROM 30 and RAM 32, the
potentiostat 24, A/D converter 26, RF module 38, two crystals 36,
44, and a variety of other supporting components, deposited bonding
pads, and conductors, which provide the necessary functionality
described above. Additionally, the electronics subassembly 46
supports an electrode system 54 including a working electrode 54a,
a reference electrode 54b, and a counter electrode 54c in one
embodiment, such as described in more detail above in the Overview
section, however alternative electrode systems and/or measurement
techniques may be implemented.
[0156] The antenna board 50, on which the antenna (42 in FIG. 2) is
disposed (not shown on FIG. 3A or FIG. 3B), is connected to the PCB
48 via antenna feed 56. Preferably the antenna 42 is surface
mounted to the antenna board 50 to further support the reduction of
size of the subassembly 46. The PCB 48 and antenna board 50 may be
formed in any typical manner such as epoxy-glass and polyamide flex
printed wiring boards, ceramic, or silicon substrates. One skilled
in the art may appreciate other hardware components, software
configurations, and interconnections not described herein.
[0157] Electrode System
[0158] Reference is now made to the electrode system 54 of the
preferred embodiments, including the working electrode (anode) 54a,
the reference electrode 54b, and the counter electrode (cathode)
54c, such as shown in FIGS. 3A and 3C. Although alternative
electrode configurations and measurement techniques may be used
with the preferred embodiments, the following description is
focused on the preferred three-electrode system, which is described
above in the Overview section.
[0159] The working electrode 54a and counter-electrode 54c of a
glucose oxidase-based glucose sensor 12 require oxygen in different
capacities. Within the enzyme layer above the working electrode
54a, oxygen is required for the production of H.sub.2O.sub.2 from
glucose. The H.sub.2O.sub.2 produced from the glucose oxidase
reaction further reacts at the surface of the working electrode 54a
and produces two electrons. The products of this reaction are two
protons (2H.sup.+), two electrons (2e.sup.-), and one oxygen
molecule (O.sub.2) (See Fraser, D. M. "An Introduction to In Vivo
Biosensing: Progress and problems." In "Biosensors and the Body,"
D. M. Fraser, ed., 1997, pp. 1-56 John P. Wiley and Sons, New
York). In theory, the oxygen concentration near the working
electrode 54a, which is consumed during the glucose oxidase
reaction, is replenished by the second reaction at the working
electrode 54a; therefore, the net consumption of oxygen is zero. In
practice, however, not all of the H.sub.2O.sub.2 produced by the
enzyme diffuses to the working electrode surface nor does all of
the oxygen produced at the electrode diffuse to the enzyme
domain.
[0160] Additionally, the counter electrode 54c utilizes oxygen as
an electron acceptor. The most likely reducible species for this
system are oxygen or enzyme generated peroxide (Fraser, D. M.
supra). There are two main pathways by which oxygen may be consumed
at the counter electrode 54c. These are a four-electron pathway to
produce hydroxide and a two-electron pathway to produce hydrogen
peroxide. Oxygen is further consumed above the counter electrode by
the glucose oxidase. Due to the oxygen consumption by both the
enzyme and the counter electrode, there is a net consumption of
oxygen at the surface of the counter electrode 54c. Thus, in the
domain of the working electrode 54a there may be significantly less
net loss of oxygen than in the region of the counter electrode 54c.
Furthermore, it is noted that there is a close correlation between
the ability of the counter electrode 54c to maintain current
balance and sensor performance. Taken together, it is believed that
counter electrode 54c function becomes limited before the enzyme
reaction becomes limited when oxygen concentration is lowered. When
this occurs, the counter electrode limitation begins to manifest
itself as this electrode moves to increasingly negative voltages in
the search for reducible species. Thus, when a sufficient supply of
reducible species, such as oxygen, is not available to a
conventional sensor, the counter electrode voltage reaches a
circuitry limit, resulting in compromised sensor performance.
[0161] In order to overcome the above-described limitations, the
diameter of the counter electrode is at least twice the diameter of
the working electrode, resulting in an approximately 6-fold
increase in the exposed surface area of the counter electrode of
the preferred embodiment. Preferably, the surface area of the
electrochemically reactive surface of the counter electrode is not
less than about 2 times the surface area of the working electrode.
More preferably, the surface area of the electrochemically reactive
surface of the counter electrode is between about 2 and about 50,
between about 2 and about 25, or between about 2 and about 10 times
the surface area of the working electrode.
[0162] Reference is now made to FIG. 3C, which is a circuit diagram
of the potentiostat 24 that controls the three-electrode system 54
of the preferred embodiments. The potentiostat includes electrical
connections to the working electrode 54a, the reference electrode
54b, and the counter electrode 54c. The voltage applied to the
working electrode 54a is a constant value and the voltage applied
to the reference electrode is also set at a constant value such
that the potential (V.sub.BIAS) applied between the working and
reference electrodes is maintained at a constant value. The counter
electrode 54c is configured to have a constant current (equal to
the current being measured by the working electrode 54a), which is
accomplished by varying the voltage at the counter electrode in
order to balance the current going through the working electrode
54a such that current does not pass through the reference electrode
54c. A negative feedback loop 52 is constructed from an operational
amplifier (OP AMP), the reference electrode 54b, the counter
electrode 54c, and a reference potential (VREF), to maintain the
reference electrode at a constant voltage.
[0163] Thus, potentiostat 24 creates current in the counter
electrode by controlling the voltage applied between the reference
and the working electrode. The reaction that occurs on the counter
electrode is determined by how much voltage is applied to the
counter electrode. By increasing the voltage applied to the counter
electrode, increased amount and type of species may react in order
to create the necessary current, which may be advantageous for the
same reasons as described above with reference to the electrode
configuration of the preferred embodiments.
[0164] In addition to the net oxygen loss described above,
implantable glucose sensors face an additional challenge in
maintaining sensor output during ischemic conditions, which may
occur either as short-term transient events (for example,
compression caused by postural effects on the device) or as
long-term low oxygen conditions (for example, caused by a thickened
FBC or barrier cells). When the sensor is in a low oxygen
environment, the potentiostat will react by decreasing the voltage
relative to the reference electrode voltage applied to the counter
electrode, which may result in other less electro-active species
reacting at the counter electrode.
[0165] In some embodiments, the potentiostat settings are
configured to allow the counter electrode to react with other
reducible species when oxygen concentration is low. In some
circumstances, glucose sensors may suffer from a negative voltage
setting that is too low, particularly in low oxygen environments.
For example, as the voltage on the counter electrode becomes more
negative, it will begin to create current, by reacting with other
reducible species, a byproduct of this reaction is H.sub.2. Two
potential problems can occur because of the production of Hydrogen
at the counter electrode: 1) bubble formation, which disconnects
the counter from the current carrying buffer and causes the sensor
to lose function and 2) an interfering signal at the working
electrode.
[0166] In order to overcome the potential problems, the preferred
embodiments optimize the potentiostat settings to enable
functionality of the potentiostat even in low oxygen conditions
while limiting the counter electrode to ensure that the sensor does
not create conditions that could damage it. Namely, such that when
oxygen concentration decreases, the counter electrode is pushed
negative enough to allow it to react with the next most abundant
reducible species, for example water, which is not typically
limited (for example, in vivo).
[0167] In one embodiment, the potentiostat settings are optimized
by setting the allowable range for the counter electrode voltage
sufficiently wide such that the sensor can react with other
reducible species when oxygen becomes limited, while setting the
range sufficiently narrow to ensure the circuitry does not allow
excessive current draw or bubble formation to occur. The counter
electrode is preferably restricted such that the species that will
react at the counter electrode electroactive surface do not
restrict the contact of the counter electrode, while causing excess
current flow and potential current damage. Thus, the negative
voltage range is preferably wide enough to function in low oxygen
environments, while being limited enough to prevent the sensor from
applying a voltage to the counter electrode that would cause bubble
formation. The limit also provides a fail-safe mechanism for
prevention of a H.sub.2 feedback loop. If hydrogen diffuses to the
working electrode and creates current the counter electrode would
be pushed to its electronic limit. When the potentiostat reaches
the electronic limit it is no longer able to maintain the potential
applied between the working and the reference electrode and the
applied potential is decreased. At this point, a maximum limiting
electrode current condition is attained. Additionally, the
optimized potentiostat settings of the preferred embodiments
provide a failsafe mechanism that prevents a cascade reaction that
could cause damage to the sensor.
[0168] In one implementation of an implantable glucose sensor, a
reference voltage between about +0.6V and +0.8V, and preferably
about +0.7V, with respect to battery ground is chosen to ensure
functionality even in low oxygen conditions yet limiting the
counter electrode to a minimum of potential equal to ground
potential to ensure that the sensor does not create conditions that
could damage it. However, one skilled in the art appreciates that
the ratio of the electroactive surface areas of the working and
counter electrodes will influence the voltage operating point of
the counter electrode with larger counter electrode areas requiring
a less negative voltage relative to the reference electrode voltage
for the same working electrode current. Additionally, one skilled
in the art appreciates that optimization of the potentiostat to
produce the above-described results can be attained by limitations
other than on the reference voltage, for example, by limiting the
current of the working or counter electrode amplifiers to a preset
current limit or by setting or the op-amp offset (V.sub.OFFSET)
from battery ground (see FIG. 3C).
[0169] Reference is now made to FIGS. 3B and 3D, which illustrate
the electrode system of the preferred embodiments configured for
secure electrical connection and secure mechanical alignment to the
PCB 48. Although other methods for forming electrodes may be used,
the preferred embodiments provide bulk metal electrodes, which
provide for optimum quality and function of the electrodes in the
analyte sensor. It is noted that even slightly insecure attachment
and alignment of the electrodes to the board (for example, as has
been seen with soldering) may jeopardize the superior performance
of bulk metal electrodes due to slippage, disconnect, or the like.
Additionally, there is a concern that heat may cause damage to the
sensitive PCB electronics if subjected to a heat-bonding type
process. Furthermore, the subsequent manufacturing steps (for
example, molding) could alter or damage even slightly insecure
components and interconnections.
[0170] Therefore, the preferred embodiments provide for a
mechanical and electrical connection of the electrodes 54 to the
PCB 48 that is extremely secure, robust, and easily reproducible in
manufacture. In preferred embodiments, the electrodes 54 are swaged
to the PCB 48 prior to assembling the electronics subassembly.
Referring to FIG. 3D, which is an expanded cross-sectional view of
the swaged electrode 54c shown in FIG. 3B, the lower portion 58 of
the electrode 54c has been shaped by force around the PCB 48 (FIG.
3B) such that the electrode is tightly held within the necessary
electrical connections 60 therein. It is noted that swaged
connections may additionally include a solder bead to provide
further reliability of the electrical connector.
[0171] Swaging is a process whereby metal is shaped by hammering or
pressure with the aid of a form or anvil called a swage block,
substantially without heating. Notably, swaging is a solderless
attachment with good mechanical accuracy, stability, orientation,
and provides a quick and clean method of manufacture. The resulting
swaged electrode-to-PCB connection 60 at least in part enables the
reliable encapsulation of the electronics subassembly in a molded
material and longevity of the device due to reliability and
reproducibility of a stable electrode system 54.
[0172] In some alternative embodiments the electrodes 54 are welded
to the PCB 48, which may include, for example, spot welding, laser
welding, ultrasonic welding, or the like. Although these techniques
include some heating, they are typically cleaner than conventional
soldering techniques, for example, and may be advantageous in the
some embodiments.
[0173] RF Telemetry
[0174] FIG. 3E is a bottom view of the PCB 48 only, showing the
side of the PCB 48 that faces the antenna board 50 in one
embodiment. Notably, the RF module 38 is provided on the PCB 48 and
includes VCO circuitry, for example, an inductor 62 that provides a
magnetic field suitable for RF telemetry. It is further noted that
in some preferred embodiments, the sensor body is substantially
formed from a water vapor permeable material, such as described in
more detail with reference to FIGS. 4 and 5, below. Unfortunately,
if water vapor penetrates through the sensor body to a location
that is within the magnetic field produced by the inductor 62,
distortion of the electromagnetic field effects may create shifts
in the carrier frequency. Generally speaking, when the VCO is
unable to provide a stable carrier frequency, the RF transmissions
are unlikely or unable to successfully reach their designated
receiver (for example, the receiver 14), which has been tuned to
the specified carrier frequency. Therefore it is advantageous to
reduce or prevent water vapor from entering and distorting the
magnetic field created by the inductor in order to maintain a
stable dielectric constant within a fixed distance from sensitive
RF electronic components.
[0175] FIGS. 3F and 3G are cross-sectional side views of the
inductor 62 on a cut-away portion of the PCB 48, which controls the
magnetic field described above. In preferred embodiments, a spacer
is provided that substantially prevents or inhibits water vapor
from permeating to the magnetic field described above. In one
embodiment, such as shown in FIG. 3F, the spacer includes a small
volume of a substance 64 applied over the inductor 62 prior to the
subsequent manufacturing steps. In one embodiment, the substance 64
is a plastic material, for example, epoxy or silicone. In this
case, the purpose of the spacer is to create a fixed space between
the inductor 62 and the water vapor permeable sensor body (FIGS. 4
and 5); although epoxy and silicone are known to be water vapor
permeable, if desired, an additional less- or minimally-water vapor
permeable coating 66 may be applied over part or all of the
electronics subassembly 46 including this spacer 64, prior to
subsequent manufacture, which is described in more detail
below.
[0176] In an alternative embodiment, the inductor 62 may be
shielded from water vapor by a metal dome, box, or cover 68, for
example, such as shown in FIG. 3G. In this embodiment, the metal
cover 68 provides the necessary protection from water vapor,
however it is noted that the metal cover 68 should be grounded to
provide a known, stable potential in the near field of the
inductor. Alternatively, the cover 68 may be glass or other
hermetic enclosure. Other alternatives for ensuring a stable
electromagnetic field may be implemented with the devices of the
present invention; for example, a toroidal inductor, which creates
a smaller electromagnetic field surrounding the inductor, may be
used, thereby decreasing the spacing required. Although two
examples are illustrated herein, in general, any design or
configuration that provides a substantially water vapor-free space
surrounding the magnetic field may be considered a spacer for
purposes of the preferred embodiments.
[0177] Referring now to the configuration of the RF telemetry
module of the preferred embodiments, the hardware and software are
designed for low power requirements to increase the longevity of
the device (for example, to enable a life of 3 to 24 months) with
maximum RF transmittance from the in vivo environment to the ex
vivo environment (for example, about one to ten meters).
Preferably, a high frequency carrier signal in the range of 402 to
405 MHz is employed in order to maintain lower power requirements.
Additionally, the carrier frequency is adapted for physiological
attenuation levels, which is accomplished by tuning the RF module
in a simulated in vivo environment to ensure RF functionality after
implantation. Accordingly, it is believed that the preferred
glucose sensor can sustain sensor function for greater than 3
months, greater than 6 months, greater than 12 months, and greater
than 24 months.
[0178] In some alternative embodiments, hermetic packaging
encompasses some parts of the implantable analyte sensor, while
water vapor permeable packaging encompasses other parts of the
implantable analyte sensor. For example, the implantable analyte
sensor body may be formed from a hermetic material (such as
Titanium), which encompasses the RF circuitry and/or other water
vapor-sensitive components; and a water vapor permeable insert or
piece may be incorporated onto the hermetic body, which encompasses
the antenna and/or other non-water vapor-sensitive components
operably connected thereto. In this way, the RF circuitry and other
sensitive components are protected from negative effects that water
vapor causes, while allowing unobstructed transmissions and
receiving via the antenna.
[0179] Protective Coating
[0180] In the preferred embodiments, a substantial portion of the
electronics 46 is coated with a conformal coating 66. This
conformal coating preferably has a water permeability rate that is
less than the water permeability rate of the sensor body and
enables sufficient spacing for the electromagnetic field as
described in more detail above with reference to FIGS. 3F and 3G;
additionally the coating protects the PCB 48 and any coated portion
of the electronics subassembly 46 from damage during the molding
process (FIGS. 4 and 5) and from water permeation that may imbibe
through the molded water vapor permeable sensor body over the
lifetime of the sensor in vivo.
[0181] In one preferred embodiment, one or more conformal Parylene
coatings are applied prior to encapsulation in the sensor body.
Parylene is known to have a slow water vapor permeability rate and
is suitable for biomedical applications. The Parylene coating
process exposes product to the gas-phase monomer at low pressure.
Through vacuum deposition, Parylene condenses on the object's
surface in a polycrystalline fashion, providing a coating that is
truly conformal and pinhole free. Compared to liquid processes, the
effects of gravity and surface tension are negligible so there is
no bridging, thin-out, pinholes, puddling, run-off or sagging;
additionally, the process takes place at room temperature so there
is no thermal or mechanical stress on the product. Parylene is
physically stable and chemically inert within its usable
temperature range. Parylene provides excellent protection from
water vapor, corrosive vapors, and solvents, for example. In
alternative embodiments, other conformal coatings (for example,
HumiSeal.RTM., Woodside, N.Y.), spray coatings, or the like, may be
used for the less- or minimally-water vapor penetrable layer, which
protect the PCB 48 and electronics subassembly 46 from damage
during the molding process (FIGS. 4 and 5) and from water
penetration through the molded water vapor permeable sensor body in
vivo.
[0182] In one alternative embodiment, a coating of a secondary
material, such as silicone, is applied after the protective coating
66. The secondary coating is preferably made from a material that
is able to absorb mechanical stresses that may be translated from
the molding process to the sensitive electrical components beneath
the protective coating. Thus, silicone, or other similar material
with sufficient elasticity or ductility, may be applied to the
coated electronics subassembly prior to forming the sensor body,
which is described in more detail below.
Sensor Body
[0183] In one embodiment, the body of the sensor is preferably
formed from a plastic material molded around the sensor
electronics, however in alternative embodiments, the body may be
formed from a variety of materials, including metals, ceramics,
plastics, resins, or composites thereof.
[0184] It is noted that conventional prior art implantable sensors
that have electronics therein generally use a hermetic material for
at least a portion of the body that houses the sensitive
electronics. However conventional hermetic implantable devices
suffer from numerous disadvantages including: difficulty in RF
transmissions through the hermetic material, seams that may allow
water vapor penetration if not perfectly sealed, minimal design or
shape changes without major manufacturing changes (inability to
rapidly iterate on design), and need to mechanically hold and
reinforce the electronics inside, increased weight and density, for
example.
[0185] To overcome the disadvantages of the prior art, the
preferred embodiments mold a plastic material around the
electronics subassembly 46 (FIG. 4B) to form the sensor body, which
enables rapid design iterations (for example, changes in design
geometry without mold changes), machining into precise dimensions
and curvatures, aids with RF transmissions, adds mechanical
integrity to components (for example, because the material fills
around the subassembly 46 to form a monolithic piece and hold
components in place), allows multiple cures (for example, to
provide a seamless exterior), and reinforces fragile electrical
components. In preferred embodiments, the material is epoxy,
however other plastics may also be used, for example, silicone,
urethane, or the like.
[0186] Referring now to the molding process for forming the body, a
two-step process is preferably used in order to provide a total
seal of the components within the device and for forming a seamless
device with a desired curvature thereon. The two-step molding
process decreases micro-fissures that may form during the initial
molding step, while mechanically securing and protecting the
electrical components. However, some alternative embodiments may
utilize a one-step molding process, for example by
injection-molding the device.
[0187] FIG. 4A is a perspective view of the primary mold which is
used in the primary casting in one embodiment. FIG. 4B is a
cross-sectional side view of the electronics subassembly during the
primary casting process. FIG. 4C is a perspective view of the
primary potted device after the primary casting process.
[0188] During the primary casting process, the primary mold 70 is
preferably pre-filled with a predetermined amount of a selected
plastic material in order to substantially cover the lower portion
72 of the primary mold 70. The electronics subassembly 46 is then
pressed into the pre-filled primary mold 70, after which the
material 74 is filled around the electronics subassembly 46 to a
predetermined fill line or weight amount, ensuring minimal or no
air bubbles exist within the material (FIG. 4B). Because of
pre-filling, the electrodes 54 are fully encapsulated within the
material during the primary cast, which will be machined later to
expose electroactive surfaces of the electrodes. This process
enables a secure and seamless encapsulation of the electrode system
54 in an insulating material and provides for mechanical alignment,
security, and reduces or eliminates leakage of water through the
sensing region. Finally, a primary hold down fixture is secured
over the device. The material is cured using standard techniques
known in the art, for example the plastic material may be placed
into a pressure vessel and heated, or the like.
[0189] FIG. 4C shows the cured material 74 substantially
encapsulating the electronics subassembly 46 after curing,
hereinafter referred to as the primary potted device 78. At this
point, the components of the electronics subassembly are
mechanically aligned and secured with sensitive parts protected
from external exposure or damage. However, some parts of the
subassembly 46 (for example those parts that were contacting the
lower portion of the primary mold) are not encapsulated or covered.
Therefore a secondary casting process is provided to fully
encapsulate the electronics subassembly, including those portions
that are exposed after the primary casting process. Secondary
casting enables a seamless and robust sensor body while reinforcing
micro-fissures or other micro-damage that may have occurred during
the primary casting.
[0190] FIG. 5A is a perspective view of inserting the primary
potted device into a secondary mold. FIG. 5B is a perspective view
of the secondary potted device after the secondary casting
process.
[0191] During the secondary casting process, the secondary mold 80
is pre-filled with a predetermined amount of the selected material
in order to substantially cover the lower portion of the secondary
mold 80. The primary potted device 78 is then pressed into the
pre-filled secondary mold 80, after which the material 74 is filled
around the primary potted device to a predetermined fill line or
weight amount, ensuring minimal or no air bubbles exist within the
material. Finally a secondary hold down fixture is secured over the
device. The material is cured using standard techniques known in
the art, for example the plastic material may be placed into a
pressure vessel and heated, or the like. It is noted that this
secondary casting is advantageous because it creates additional
strength over a primary casting alone, fills in micro-fissures or
other damage that may have occurred during or after the primary
process, and provides a seamless exterior to prevent leakage into
the device.
[0192] Reference is now made to FIG. 5B, which shows the secondary
potted device 82 prior to final machining. Because of the nature of
the plastic material, the sensor can be machined with precise shape
and dimensions. For example, the preferred embodiments are machined
to a sensor geometry that optimizes healing at the sensor-tissue
interface in vivo and is less amenable to accidental movement due
to shear and rotational forces than other sensor configurations,
such as described in more detail below with reference to FIGS. 6A
to 6C. Additionally, machining the molded material to expose the
electrodes 54 enables careful exposure of the electroactive
surfaces.
[0193] It is noted that additional outer coatings may be
advantageously applied to the secondary potted device 82, such as
one or more Parylene coatings, in order to decrease water vapor
penetration, for example. As another example of an outer coating, a
silicone layer may be applied to the non-sensing region of the
device, which serves to fill in any microfissures or micropores,
for example, within the molded material, to provide a smooth outer
surface, and/or to enable attachment of additional materials (for
example, a silicone anchoring material). Other coatings may be
applied as is appreciated by one skilled in the art.
[0194] In some alternative embodiments, the electronics subassembly
46 is placed within a pre-formed shell, rather than molding the
body around the subassembly. In these alternative embodiments, the
shell configuration advantageously provides air space surrounding
the electronics, which aids in maintaining a stable dielectric
constant surrounding the VCO circuitry, such as described in more
detail with reference to FIGS. 3E to 3G. Additionally, the shell
provides protection for the electronics subassembly 46 from damage
that may occur during a molding process. The pre-formed body shell
may be advantageous for manufacture, because it enables the molding
process to be separated from the electronics subassembly, reducing
the amount of error that may occur to the electronics subassembly
in process. It is noted that conformal coatings may be applied to
the electronics subassembly prior to encapsulation in the shell,
and/or may be applied to the shell itself Coatings provide a
variety of advantages discussed para supra, and with reference to
FIG. 3F, for example.
[0195] Sensor Geometry
[0196] FIG. 6A is a perspective view of an analyte sensor in one
embodiment, including a thin substantially oval geometry, a curved
sensing region, and an overall curved surface on which the sensing
region is located, thereby causing contractile forces from the
foreign body capsule to press downward on the sensing region. FIG.
6B is an end view of the analyte sensor of FIG. 6A showing the
contractile forces that would be caused by a foreign body capsule.
FIG. 6C is a side view of the analyte sensor of FIG. 6A.
[0197] In this illustration, the analyte sensor 12 is shown without
subsequent membrane and anchoring material thereon and is used to
illustrate sensor geometry. The analyte sensor 12 includes the
sensing region 16 located on a curved portion of the sensor body,
and including no abrupt edge or discontinuous surface in the
proximity of the sensing region. Additionally, the overall
curvature of the surface on which the sensing region is located,
including rounded edges, invokes a generally uniform FBC around
that surface, decreasing inflammatory response and increasing
analyte transport at the device-tissue interface.
[0198] Perpendicular forces 84, depicted in FIG. 6B by arrows
pointing down, represent the forces presented by the foreign body
capsule on the device in vivo which have been found to reduce or
eliminate shear forces with the tissue at the sensing region. While
lateral forces 86 may appear to create shear forces at the sensing
region, several features of the sensor mitigate these forces. For
example, the sensor comprises a relatively thin aspect ratio (low
profile) and preferably is implanted adjacent to the fascia,
underlying the fat, making it less prone to movement. As another
example, in some embodiments the sensor may include short-term
anchoring components, for example the sensor may be sutured to the
tough fascia, which aids in preventing lateral forces from being
conveyed to the sensing region (FIG. 13A). In some embodiments, the
sensor may include long-term anchoring components, for example an
anchoring material may be employed (FIG. 13B). As yet another
example, in order to facilitate proper healing, the side of the
sensor upon which the sensing region is situated preferably has a
curved radius extending from lateral side to lateral side. As
depicted in the side view and end view (FIGS. 6B and 6C), the
sensing region is positioned at the apex of the radius. When
surrounding tissue contracts as it heals, the radius serves to
optimize the forces 84 exerted down onto the curved surface,
especially the forces in the lateral directions 86, to keep the
tissue uniformly in contact with the surface and to produce a
thinner foreign body capsule. The curvature ensures that the head
is resting against the tissue and that when tissue contraction
occurs, forces are generated downward on the head so that the
tissue attachment is maintained. It may be noted that the downward
forces bring the tissue into contact with porous biointerface
materials used for ingrowth-mediated attachment and for
biointerface optimization, such as described in more detail with
reference to FIGS. 8A to 8B.
[0199] Sensing Membrane
[0200] In preferred embodiments, the sensing membrane is
constructed of two or more domains and is disposed adjacent to the
electroactive surfaces of the sensing region 16. The sensing
membrane provides functional domains that enable measurement of the
analyte at the electroactive surfaces. For example, the sensing
membrane includes an enzyme, which catalyzes the reaction of the
analyte being measured with a co-reactant (for example, glucose and
oxygen) in order to produce a species that in turn generates a
current value at the working electrode, such as described in more
detail above in the Overview section. The sensing membrane can be
formed from one or more distinct layers and can comprise the same
or different materials.
[0201] In some embodiments, the sensing membrane 88 includes an
enzyme, for example, glucose oxidase, and covers the electrolyte
phase. In one embodiment, the sensing membrane 88 generally
includes a resistance domain 90 most distal from the
electrochemically reactive surfaces, an enzyme domain 92 less
distal from the electrochemically reactive surfaces than the
resistance domain, and an electrolyte domain 96 adjacent to the
electrochemically reactive surfaces. However, it is understood that
a sensing membrane modified for other devices, for example, by
including fewer or additional domains, is within the scope of the
preferred embodiments. Co-pending U.S. patent application Ser. No.
09/916,711, entitled, "SENSOR HEAD FOR USE WITH IMPLANTABLE
DEVICES" and U.S. patent application Ser. No. 10/153,356 entitled,
"TECHNIQUES TO IMPROVE POLYURETHANE MEMBRANES FOR IMPLANTABLE
GLUCOSE SENSORS," which are incorporated herein by reference in
their entirety, describe membranes that can be used in the
preferred embodiments. It is noted that in some embodiments, the
sensing membrane 88 may additionally include an interference domain
94 that limits some interfering species; such as described in the
above-cited co-pending patent application. Co-pending U.S. patent
application Ser. No. 10/695,636, entitled, "SILICONE COMPOSITION
FOR BIOCOMPATIBLE MEMBRANE" also describes membranes that may be
used for the sensing membrane of the preferred embodiments, and is
incorporated herein by reference in its entirety.
[0202] In some embodiments, the domains of the sensing membrane are
formed from materials such as silicone, polytetrafluoroethylene,
polyethylene-co-tetrafluoroethylene, polyolefin, polyester,
polycarbonate, biostable polytetrafluoroethylene, homopolymers,
copolymers, terpolymers of polyurethanes, polypropylene (PP),
polyvinylchloride (PVC), polyvinylidene difluoride (PVDF),
polybutylene terephthalate (PBT), polymethylmethacrylate (PMMA),
polyether ether ketone (PEEK), polyurethanes, cellulosic polymers,
polysulfones and block copolymers thereof including, for example,
di-block, tri-block, alternating, random and graft copolymers, or
the like.
[0203] FIG. 7A is an illustration that represents a method of
forming the sensing membrane in one embodiment. FIG. 7B is a
schematic side view of the sensing membrane in one embodiment. In
this embodiment, the sensing membrane 88 includes a resistance
domain 90, an enzyme domain 92, an interference domain 94, and an
electrolyte domain 96. Preferably, the domains are serially cast
upon a liner 98, all of which are formed on a supporting platform
100; however alternative embodiments may form the membrane domains
directly on the sensing region 16, for example, by spin-, spray-,
or dip-coating.
[0204] Referring now to the function of the resistance domain 90,
it is noted that there exists a molar excess of glucose relative to
the amount of oxygen in blood; that is, for every free oxygen
molecule in extracellular fluid, there are typically more than 100
glucose molecules present (see Updike et al., Diabetes Care
5:207-21 (1982)). However, an immobilized enzyme-based sensor
employing oxygen as cofactor should 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
situated over the enzyme domain to control the flux of glucose and
oxygen, a linear response to glucose levels can be obtained only up
to about 40 mg/dL. However, in a clinical setting, a linear
response to glucose levels is desirable up to at least about 500
mg/dL.
[0205] The resistance domain 90 includes a semipermeable membrane
that controls the flux of oxygen and glucose to the underlying
enzyme domain 92, preferably rendering oxygen in a
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 is achieved without the resistance domain. In one embodiment,
the resistance domain 90 exhibits an oxygen-to-glucose permeability
ratio 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
(See Rhodes et al., Anal. Chem., 66:1520-1529 (1994)).
[0206] In some alternative embodiments, a lower ratio of
oxygen-to-glucose may be sufficient to provide excess oxygen by
using an oxygen antenna domain (for example, a silicone or
fluorocarbon based material or domain) to enhance the
supply/transport of oxygen to the enzyme domain. In other words, if
more oxygen is supplied to the enzyme, then more glucose may also
be supplied to the enzyme without creating an oxygen rate-limiting
excess. In some alternative embodiments, the resistance domain is
formed from a silicone composition, such as described in copending
U.S. application Ser. No. 10/685,636 filed Oct. 28, 2003 and
entitled, "SILICONE COMPOSITION FOR BIOCOMPATIBLE MEMBRANE," which
is incorporated herein by reference in its entirety.
[0207] In one preferred embodiment, the resistance layer includes a
homogenous polyurethane membrane with both hydrophilic and
hydrophobic regions to control the diffusion of glucose and oxygen
to an analyte sensor, the membrane being fabricated easily and
reproducibly from commercially available materials.
[0208] In the preferred embodiment, the hydrophobic polymer
component is a polyurethane. In a most preferred embodiment, the
polyurethane is polyetherurethaneurea. A polyurethane is a polymer
produced by the condensation reaction of a diisocyanate and a
difunctional hydroxyl-containing material. A polyurethaneurea is a
polymer produced by the condensation reaction of a diisocyanate and
a difunctional amine-containing material. Preferred diisocyanates
include aliphatic diisocyanates containing from 4 to 8 methylene
units. Diisocyanates containing cycloaliphatic moieties may also be
useful in the preparation of the polymer and copolymer components
of the membrane of the present invention. The material that forms
the basis of the hydrophobic matrix of the resistance domain may be
any of those known in the art as appropriate for use as membranes
in sensor devices and having sufficient permeability to allow
relevant compounds to pass through it, for example, to allow an
oxygen molecule to pass through the membrane from the sample under
examination in order to reach the active enzyme or electrochemical
electrodes. Examples of materials which may be used to make a
non-polyurethane type membrane include vinyl polymers, polyethers,
polyesters, polyamides, inorganic polymers such as polysiloxanes
and polycarbosiloxanes, natural polymers such as cellulosic and
protein based materials and mixtures or combinations thereof.
[0209] In a preferred embodiment, the hydrophilic polymer component
is polyethylene oxide. For example, one useful
hydrophobic-hydrophilic copolymer component is a polyurethane
polymer that includes about 20% hydrophilic polyethylene oxide. The
polyethylene oxide portion of the copolymer is thermodynamically
driven to separate from the hydrophobic portions of the copolymer
and the hydrophobic polymer component. The 20% polyethylene oxide
based soft segment portion of the copolymer used to form the final
blend controls the water pick-up and subsequent glucose
permeability of the membrane of the preferred embodiments.
[0210] The preferred embodiments additionally provide a method for
preparing the resistance domain, which provides a homogeneous and
uniform structure. The homogeneous, uniform structure is
advantageous in order to ensure that glucose traversing through the
resistance domain adequately reaches the electroactive surfaces of
the electrode system, which is described in more detail with
reference to co-pending U.S. patent application Ser. No. 09/916,711
filed Jul. 27, 2001, entitled "SENSOR HEAD FOR USE WITH IMPLANTABLE
DEVICE," and U.S. patent application Ser. No. 10/153,356, entitled,
"TECHNIQUES TO IMPROVE POLYURETHANE MEMBRANES FOR IMPLANTABLE
GLUCOSE SENSORS," both of which are incorporated herein by
reference in their entirety.
[0211] The preferred method of casting the resistance domain 90
includes the steps of: (a) forming a solvent solution of a
hydrophilic polymer and a hydrophobic polymer; (b) maintaining the
composition at a temperature sufficient to maintain the hydrophobic
polymer and the hydrophilic polymer substantially soluble; (c)
applying the composition at the temperature to a liner 98 to form a
film thereon; and (d) permitting the solvent to evaporate from the
resultant film to form the membrane.
[0212] In the preferred embodiments, the composition is applied to
a liner 98, for example, an uncoated Polyethylene Terephthalate
(PET), which provides low risk of variability and contamination,
for example, in contrast to a coated liner. It is noted that a
membrane cured on uncoated PET can be easily removed after 1-hour
hydration in room temperature PBS. In some alternative embodiments,
other liners and release layers may be used. The platform 100
provides a support for casting and may be for example, the base of
a drawdown machine, or the like. While not wishing to be bound by
theory, it is believed that the sensing membranes of the preferred
embodiments are more consistently produced by casting individual
layers serially rather than using a continuous web coating machine,
which will be described in more detail throughout the
description.
[0213] In one embodiment, the forming step includes forming a
mixture or a blend of material for the resistance domain. As
described above, in preferred embodiments, the first polymer is a
polyurethane and the second polymer is a polyurethane comprising a
polyethylene oxide. In general, the second polymer may be a random
or ordered block copolymer.
[0214] The blend is heated substantially above room temperature in
order to mix the hydrophilic and hydrophobic components with each
other and the solvent. In one embodiment, the composition of the
blend of the hydrophilic polymer and the hydrophobic polymer is
heated to a temperature of at least about 70.degree. C. for a
predetermined time period (for example, at least about 24 hours,
preferably at least about 44 hours) to ensure the first and second
polymers are substantially intermixed. One skilled in the art
appreciates that the level of heating is dependent upon the
relative miscibility of the polymer components and may be adjusted
accordingly.
[0215] The preferred embodiments cure the coated film formed on the
liner 98 to dry at an elevated temperature. Additionally, the
temperature is ramped up during the curing process. In one
embodiment, the coated film is placed in an oven wherein the
temperature is ramped at a preferred ramp rate of within the range
of 3.degree. C./minute and 12.degree. C./minute, more preferably
7.degree. C. per minute from a first elevated temperature to a
second elevated temperature. Preferably, the first elevated
temperature is between about 60.degree. C. and 100.degree. C., more
preferably about 80.degree. C. Preferably, the second elevated
temperature is at least about 100.degree. C. The elevated
temperature serves to drive the solvent from the coating as quickly
as possible. Ramping of the temperature serves to provide more
uniformity and fewer defects in the hydrophobic and hydrophilic
domain structures as they cure. While not wishing to be bound by
theory, it is believed that elevating the temperature prior to
coating and ramping up the temperature during curing inhibits the
hydrophilic and hydrophobic portions of the membrane from
segregating and forming large undesired structures. Membranes
prepared in this way have been shown to provide accurate sensor
operation at temperatures from about 30.degree. C. to about
45.degree. C. for a period of time exceeding about 30 days to
exceeding about 6 months.
[0216] In one experiment, a plurality of resistance membranes (n=5)
were prepared as described above, including heating at about
80.degree. C. for greater than 24 hours prior to curing. In this
experiment, each of the membranes was cured in an oven where the
temperature was ramped at a rate of about 3.degree. C./min.,
5.degree. C./min., or 7.degree. C./min. All of the membranes
provided sufficient glucose permeability when tested (about 1.24
nA/mg/dL to 2.5 nA/mg/dL). It was noted that the rate of
permeability of glucose (namely, the sensitivity) through the
membrane decreased as a function of temperature ramp rate. Namely,
the glucose permeability of a membrane decreased as the ramp rate
used to cure that membrane increased, with a correlation (R.sup.2)
of 0.58. While not wishing to be bound by theory it is believed
that the glucose permeability can be optimized for a variety of
design requirements by altering the ramp rate at which the
resistance membrane is cured.
[0217] In preferred embodiments, the thickness of the resistance
domain is from about 10 microns or less to about 200 microns or
more. In more preferred embodiments, the thickness of the
resistance domain is from about 15, 20, 25, 30, or 35 microns to
about 65, 70, 75, 80, 85, 90, 95, or 100 microns. In more preferred
embodiments, the thickness of the resistance domain is from about
30 or 35 microns to about 40 or 45 microns.
[0218] In the preferred embodiments, the enzyme domain 92 provides
a catalyst to catalyze the reaction of the analyte and its
co-reactant, as described in more detail above. Preferably, the
enzyme domain includes glucose oxidase, however other oxidases, for
example, galactose oxidase or uricase, may be used.
[0219] 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 glucose oxidase, are subject to deactivation as a
function of time even in ambient conditions, this behavior needs to
be accounted for in constructing sensors for long-term use.
Preferably, the enzyme domain is constructed of aqueous dispersions
of colloidal polyurethane polymers including the enzyme. However,
some alternative embodiments construct the enzyme domain from an
oxygen antenna material, for example, silicone or fluorocarbons in
order to provide a supply of excess oxygen during transient
ischemia. Preferably, the enzyme is immobilized within the domain
as is appreciated by one skilled in the art. Preferably, the domain
is coated onto the resistance domain using casting techniques for a
coating thickness between about 2.5 microns and about 22 microns,
preferably about 15 microns.
[0220] In the preferred embodiments, an interference domain 94 is
provided to allow analytes and other substances that are to be
measured by the electrodes to pass through, while preventing
passage of other substances, including potentially interfering
substances. In one embodiment, the interference domain limits the
diffusion of hydrophilic species, such as ascorbate, and large
molecular weight species. Preferably, the interference domain is
constructed of polyurethane, however other materials may be
used.
[0221] Casting of the interference domain may be important in that
an interference layer that is too thick may block desired species
from being measured, while an interference domain that is too thin
may not block appropriate interfering species. The interference
domain has a preferred thickness of not more than about 5 microns,
more preferably not less than about 0.1 microns, and not more than
about 5 microns and, most preferably, not less than about 0.5
microns, and not more than about 3 microns.
[0222] Because of the extremely thin nature of the interference
domain, consistently functional interference layers have
conventionally been difficult to manufacture due to their
susceptibility to variances in the underlying domain (for example,
enzyme domain) and casting processes. In order to obtain an
interference domain with appropriate and constant thickness, the
interference solution of the preferred embodiments is cast by
applying a sufficiently diluted interference solution and drawing
down at a sufficiently fast speed in order to maintain a constant
viscosity of the liquid film with minimal solvent evaporation.
[0223] In one embodiment, a "sufficiently diluted interference
solution" includes a ratio of about 5 wt. % polymer to about 95 wt.
% solvent. However, a ratio of about 1 to 10 wt. % polymer to about
90 to 99 wt. % solvent may be used. Additionally, due to the
volatility of solvents used for in the interference solution (for
example, solvents with a boiling point slightly greater than room
temperature (about 5 to 15.degree. C.)), a sufficiently fast
casting speed is advantageous to avoid invariabilities (for
example, film thickness inhomogeneities) due to evaporation during
casting. In one embodiment, the liquid film is drawn down at a
speed of about 8 to about 15 inches/second, and preferably about
11.5 inches/second. While not wishing to be bound by theory,
optimization of the solvent dilution and draw down speed limits
solvent evaporation and viscosity buildup, which enables a very
thin but constant interference domain. Variability in the
interference domain has been discovered by the inventors to be a
significant contributor in the variability of sensor function.
[0224] In preferred embodiments, an electrolyte domain 96 is
provided to ensure an electrochemical reaction occurs at the
electroactive surfaces. Preferably, the electrolyte domain includes
a semipermeable coating that maintains hydrophilicity at the
electrochemically reactive surfaces of the sensor interface. The
electrolyte domain enhances the stability of the interference
domain 94 by protecting and supporting the material that makes up
the interference domain. The electrolyte domain also assists in
stabilizing the operation of the sensor by overcoming electrode
start-up problems and drifting problems caused by inadequate
electrolyte. The buffered electrolyte solution contained in the
electrolyte domain also protects against pH-mediated damage that
may result from the formation of a large pH gradient between the
substantially hydrophobic interference domain and the electrodes
due to the electrochemical activity of the electrodes.
[0225] In one embodiment, the electrolyte domain 96 includes a
flexible, water-swellable, substantially solid hydrogel film having
a "dry film" thickness of from about 5 microns to about 15 microns,
more preferably from about 3, 3.5, 4, 4.5, 5, or 5.5 to about 6,
6.5, 7, 7.5, 8, 8.5, 9, 9.5, 10, 10.5, 11, 11.5, or 12 microns.
"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.
[0226] In some embodiments, the electrolyte domain is formed of a
curable mixture 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 presence of polyvinylpyrrolidone and cured at a moderate
temperature of about 50.degree. C. Underlying the electrolyte
domain is an electrolyte phase is a free-fluid phase including a
solution containing at least one compound, typically a soluble
chloride salt, which conducts electric current. In one embodiment
wherein the biocompatible membrane is used with an analyte sensor
such as is described herein, the electrolyte phase flows over the
electrodes and is in contact with the electrolyte domain. The
devices of the preferred embodiments contemplate the use of any
suitable electrolyte solution, including standard, commercially
available solutions. Generally, the electrolyte phase can have the
same osmotic pressure or a lower osmotic pressure than the sample
being analyzed. In preferred embodiments, the electrolyte phase
comprises normal saline.
[0227] Underlying the electrolyte domain is an electrolyte phase,
which is a free-fluid phase including a solution containing at
least one compound, typically a soluble chloride salt, which
conducts electric current. In one embodiment wherein the
biocompatible membrane is used with an analyte sensor such as is
described herein, the electrolyte phase flows over the electrodes
and is in contact with the electrolyte domain. The devices of the
preferred embodiments contemplate the use of any suitable
electrolyte solution, including standard, commercially available
solutions. Generally, the electrolyte phase can have the same
osmotic pressure or a lower osmotic pressure than the sample being
analyzed. In preferred embodiments, the electrolyte phase comprises
normal saline.
[0228] Although the preferred embodiments provide for serially
casting of the sensing membrane, alternative embodiments may
utilize known thin or thick film fabrication techniques known in
the art (for example, continuous web or deposition techniques). In
various embodiments, any of these domains may be omitted, altered,
substituted for, and/or incorporated together without departing
from the spirit of the preferred embodiments. For example, the
interference domain may not be necessary in some embodiments, such
as when the analyte sensor is designed to reduce interfering
species using electrochemical techniques. Additionally, the various
domains may be combined in function; for example, one discrete
layer may function both as the resistance and enzyme domain. In
another such example, an oxygen antenna domain may be individually
formed from an oxygen reserving material (for example, silicone or
fluorocarbon), or may be combined with some or all of the
biointerface membrane. Additionally, the sensing membrane may be
combined with some or all of the domains of the biointerface
membrane, such as the bioprotective (cell impermeable) domain,
which is described in more detail below.
[0229] Biointerface Membrane
[0230] The preferred embodiments provide a biointerface membrane
disposed more distal to the electroactive surface than the sensing
membrane. Preferably, the biointerface membrane 106 supports tissue
ingrowth, serves to interfere with the formation of a barrier cell
layer, and protects the sensitive regions of the device from host
inflammatory response. In some embodiments, the biointerface
membrane is composed of one or more domains.
[0231] In one embodiment, the biointerface membrane 106 generally
includes a cell disruptive domain 108 most distal from the
electrochemically reactive surfaces and a cell impermeable domain
110 less distal from the electrochemically reactive surfaces than
the cell disruptive domain 108. The cell disruptive domain 108
comprises an architecture, including a cavity size, configuration,
and overall thickness that encourages vascular tissue ingrowth and
disrupts barrier cell formation in vivo and a cell impermeable
domain that comprises a cell impermeable layer that is resistant to
cellular attachment and has a robust interface that inhibits
attachment of barrier cells and delamination of the domains.
[0232] FIG. 8A is a cross-sectional schematic view of a membrane
106 in vivo in one embodiment, wherein the membrane comprises a
cell disruptive domain 108 and cell impermeable domain 110. The
architecture of the membrane 106 provides a robust long-term
implantable membrane that allows the transport of analytes through
vascularized tissue ingrowth without the formation of a barrier
cell layer.
[0233] The cell disruptive domain 108 comprises a solid portion 112
and a plurality of interconnected three-dimensional cavities 114
formed therein. The cavities 114 have sufficient size and structure
to allow invasive cells, such as fibroblasts 116, fibrous matrix
118, and blood vessels 120 to completely enter into the apertures
40 that define the entryway into each cavity 114, and to pass
through the interconnected cavities toward the interface 122
between the cell disruptive and cell impermeable domains (cells and
blood vessels are disproportionately large in the illustration).
The cavities 114 comprise an architecture that encourages the
ingrowth of vascular tissue in vivo as indicated by the blood
vessels 120 formed throughout the cavities. Because of the
vascularization within the cavities, solutes 126 (for example,
oxygen, glucose and other analytes) can pass through the first
domain with relative ease and/or the diffusion distance (i.e.,
distance that the glucose diffuses) can be reduced.
[0234] The cell impermeable domain 110 comprises a cell impermeable
layer that may be resistant to cellular attachment and thus
provides another mechanism for resisting barrier cell layer
formation (indicated in FIG. 8A by few macrophages and/or giant
cells at the interface 122 between the domains). Because the cell
impermeable domain 110 is resistant to cellular attachment and
barrier cell layer formation, the transport of solutes such as
described above can also pass through with relative ease without
blockage by barrier cells as seen in the prior art.
[0235] Reference is now made to FIG. 8B, which is an illustration
of the membrane of FIG. 8A, showing contractile force lines caused
by the fibrous tissue (for example, from the fibroblasts and
fibrous matrix) of the FBR. Particularly, the architecture of the
cell disruptive domain 108, including the cavity interconnectivity
and multiple-cavity depth, (i.e., two or more cavities in three
dimensions throughout a substantial portion of the first domain)
can affect the tissue contracture that typically occurs around a
foreign body.
[0236] It is noted that a contraction of the FBC around the device
as a whole produces downward forces on the device, such as shown in
FIGS. 6B and 6C, which can be helpful in reducing motion artifacts
such as described with reference to copending U.S. patent
application Ser. No. 10/646,333 entitled "OPTIMIZED SENSOR GEOMETRY
FOR AN IMPLANTABLE GLUCOSE SENSOR," which is incorporated herein in
its entirety by reference. However, the architecture of the first
domain described herein, including the interconnected cavities and
solid portion, are advantageous because the contractile forces
caused by the downward tissue contracture that can otherwise cause
cells to flatten against the device and occlude the transport of
analytes, is instead translated to, disrupted by, and/or
counteracted by the forces 128 that contract around the solid
portions 112 (for example, throughout the interconnected cavities
114) away from the device. That is, the architecture of the solid
portions 112 and cavities 114 of the cell disruptive domain cause
contractile forces 128 to disperse away from the interface between
the cell disruptive domain 108 and cell impermeable domain 110.
Without the organized contracture of fibrous tissue toward the
tissue-device interface typically found in a FBC, macrophages and
foreign body giant cells substantially do not form a monolayer of
cohesive cells (i.e., barrier cell layer) and therefore the
transport of molecules across the second domain and/or membrane is
substantially not blocked (indicated by free transport of analytes
126 through the domains in FIG. 8A).
[0237] Co-pending U.S. patent application Ser. No. 09/916,386,
entitled, "MEMBRANE FOR USE WITH IMPLANTABLE DEVICES," U.S. patent
application Ser. No. 10/647,065, entitled, "POROUS MEMBRANES FOR
USE WITH IMPLANTABLE DEVICES," and U.S. Provisional Patent
Application 60/544,722, entitled "BIOINTERFACE WITH INTEGRATED
MACRO- AND MICRO-ARCHITECTURES," describe biointerface membranes
that may be used in conjunction with the preferred embodiments, and
are incorporated herein by reference in their entirety.
[0238] The cell disruptive and cell impermeable domains can be
formed from materials such as silicone, polytetrafluoroethylene,
polyethylene-co-tetrafluoroethylene, polyolefin, polyester,
polycarbonate, biostable, homopolymers, copolymers, terpolymers of
polyurethanes, polypropylene (PP), polyvinylchloride (PVC),
polyvinyl alcohol (PVA), polyvinylidene fluoride (PVDF),
polybutylene terephthalate (PBT), polymethylmethacrylate (PMMA),
polyether ether ketone (PEEK), polyurethanes, cellulosic polymers,
polysulfones or block copolymers thereof including, for example,
di-block, tri-block, alternating, random and graft copolymers.
[0239] The cell disruptive domain and cell impermeable domain of
the biocompatible membrane can be formed together as one unitary
structure. Alternatively, the cell disruptive and cell impermeable
domains of the biocompatible membrane can be formed as two layers
mechanically or chemically bonded together. In yet another
embodiment, the cell impermeable domain is chemically or
mechanically attached to the sensing membrane. In some embodiments,
the bioprotective function of the cell impermeable domain is
inherent in the structure of the sensing membrane and therefore no
discrete cell impermeable domain is required.
[0240] Membrane Attachment
[0241] FIG. 9A is an exploded perspective view of the analyte
sensor prior to membrane attachment. FIG. 9B is a perspective view
of the analyte sensor after membrane attachment. In preferred
embodiments, the membrane 130 is attached to the sensing region 16
of the sensor 12 via a mechanical fastener 132, which in the
illustrated embodiment of FIGS. 9A and 9B is a clip 132a.
[0242] In some embodiments, both the sensing membrane 88 and
biointerface membranes 106 are placed to together and the membrane
130 attached to the sensing region 16 via the mechanical fastener
132. In some alternative embodiments, some portion of the
biointerface membrane 106 (for example, the cell disruptive domain
108) may be attached without the mechanical fastener 132. In
preferred embodiments, the cell impermeable domain 110 is formed
separately from the cell disruptive domain 108 and the cell
impermeable domain 110 is attached to the sensing region 16
simultaneously with the sensing membrane 88 via the mechanical
fastener 132.
[0243] In attaching the membrane 130 to the sensing region 16, it
is desirable to avoid over-stretching of the membranes, which may
create cracks or fissures that allows cells to penetrate to the
sensing membrane. It is also desirable to avoid lack of tension in
the membrane, which may create excess spacing under the membrane
(for example, bubbles), which distorts device function.
[0244] The preferred embodiments provide an advantageous method of
attachment, wherein the membrane 130 is applied to the sensing
region 16 with appropriate tension to optimize sensor performance;
namely, by minimizing tearing or excess spacing. Preferably, the
membranes are hydrated prior to attachment, which minimizes or
avoids distortion of the membrane. Preferably, the hydrated sensing
membrane 88 is first placed over the sensing region 16 (after being
released from the liner 98). Next, the cell impermeable domain 110
is placed over the sensing membrane 88, ensuring no wrinkles or
bubbles exist (in embodiments wherein the cell disruptive domain is
attached to the sensing region under the mechanical fastener 132,
the cell disruptive domain would also be disposed over the cell
impermeable domain). Finally the mechanical fastener is applied
over the sensing region 16 and preferably into a groove surrounding
the sensing region 16. In general, the mechanical fastener 132 (for
example, a metal or plastic O-ring in FIGS. 9A and 9B) is designed
provide sufficient tension to the combined membrane 130 in order to
maintain tautness and provide a seal that prevents cellular
ingrowth. Advantageously, this combined membrane 130 may be
unattached, and a new membrane 130 attached as necessary, making
the system and method reusable and cost-effective.
[0245] FIGS. 9C to 9H are exploded and collapsed perspective views
of alternative membrane attachment embodiments. In each embodiment,
the mechanical fastener 132 is provided for attaching the membrane
130 to the sensing region 16 substantially as described above. It
is noted that although chemical attachment techniques (such as an
adhesive or solvent) may be used to enhance the mechanical
attachment, it is believed that the preferred mechanical
attachments provide appropriate tension to enable a sufficient seal
of the membrane 130 on the sensing region 16 without chemically
altering the membrane 130. Additionally, mechanical attachment of
the membrane makes the system and method reusable and
cost-effective as described above.
[0246] FIGS. 9C and 9D are perspective views of one alternative
embodiment, wherein the mechanical fastener 132 is a metal or
plastic disc 132b adapted to be press- or snap-fit into the device
body 12. Namely, the disc 132b has a plurality of legs, which are
designed to securely fit into holes in the sensor body 12. The disc
132b has a central aperture sized and arranged to provide access to
the sensing region 16 of the device so that analytes may pass
therethrough. The membrane 130 may be any appropriate
configuration, such as described in more detail above. The membrane
attachment of FIGS. 9C and 9D is advantageous for the reasons
described above and further may enable a lower profile of the
sensing region 16 as compared to the embodiment of FIGS. 9A and 9B,
for example. While not wishing to be bound by theory, it is
believed that design optimization (for example, reduction of size,
mass, and/or profile) of the device enables a more discrete and
secure implantation than a larger device, and is believed to reduce
macro-motion of the device induced by the patient (for example,
fiddling) and micro-motion caused by movement of the device (for
example, which produces chronic inflammation) within the
subcutaneous pocket, and thereby improves overall device
performance.
[0247] FIGS. 9E and 9F are perspective views of another alternative
embodiment, wherein the mechanical fastener 132 is a metal or
plastic, ring or donut 132c adapted to be press- or snap-fit into
the device body 12. In this embodiment, the donut 132c is designed
to substantially fill the aperture surrounding the sensing region
16. While this embodiment is similar to that of FIGS. 9A and 9B,
the sizing of the donut to substantially fill the aperture provides
an optimized seal of the membrane edges and device body, and
reduces opportunity for cellular ingrowth.
[0248] FIGS. 9G and 9H are perspective views of yet another
alternative embodiment, wherein the mechanical fastener 132 is a
metal or plastic clip 132d adapted to be press- or snap-fit into
sides of the device body 12. This embodiment provides advantages of
a low profile body with a substantially flat, smooth upper surface
of the device body. It is believed the a smoother upper surface
provides less opportunity for inflammation proximal the sensing
region in vivo such a described in co-pending U.S. patent
application Ser. No. 10/646,333, which is incorporated herein by
reference in its entirety.
[0249] Long- and Short-Term Anchoring
[0250] The final step in the assembly of the implantable sensor
includes attaching the outermost layers, which serve as the
device-tissue interface and may play a critical role in device
stabilization in vivo. The preferred embodiments may be designed
with short- and/or long-term anchoring systems and methods in order
to ensure stabilization of the device in vivo. As discussed above
and in more detail below, stabilization of the device in the
subcutaneous tissue is believed to impact the performance of the
sensor short- and long-term. In one preferred embodiment, the
sensor comprises a short-term anchoring component configured to
anchor the sensor to the tissue and thereby minimize motion-related
damage at the device-tissue interface, which is believed to cause
local inflammation and poor wound healing during the initial tissue
ingrowth phase. Additionally, the preferred embodiments comprise a
long-term anchoring component on the sensor to ensure long-term
stabilization of the sensor in the subcutaneous pocket. Although
both long- and short-term anchoring are preferred, some embodiments
may utilize only one or the other, for example, when the sensor is
sufficiently miniaturized such that the sensor body substantially
"floats" within the subcutaneous space, or when a precise pocket
forming or implantation technique is utilized, at least one of
short- and long-term anchoring may not be required for sufficient
sensor performance.
[0251] FIG. 10A is an exploded perspective view of the machined
sensor geometry of FIG. 6 and the tissue-facing components of the
sensor. FIG. 10B is a perspective view of the assembled analyte
sensor, including the tissue-facing components attached thereto.
FIG. 10C is a perspective view of the non-sensing side of the
assembled analyte sensor, showing a short-term anchoring device in
one embodiment.
[0252] Referring now to FIG. 10A, the tissue-facing components,
including the biointerface membrane 106, which is described in more
detail with reference to FIG. 10, a short-term anchoring component
134, and a long-term anchoring component 136 are shown. Each of
these components is securely attached to the sensor body as
described in more detail below.
[0253] A variety of attachment methods are contemplated for
attaching the biointerface membrane 106 to the sensor 12, including
mechanical attachment (such as under clip 132) and chemical
attachment (such as laser welding, ultrasonic welding, solvent
welding, or the like). In the preferred embodiments, the
biointerface membrane 106, namely the cell disruptive domain 108,
is adhered to the sensor body using an adhesive, such as a silicone
adhesive, which may be particularly advantageous when the
biointerface membrane 106 is formed from a silicone material, for
example. It is noted that the silicone adhesive is preferably
applied on the circumference of the biointerface membrane 108 to
avoid blockage of the interconnected cavities 114 of the cell
disruptive domain 108. In one preferred embodiment, the
biointerface membrane 106 shown in FIG. 10A represents only a cell
disruptive domain 108 because the cell impermeable domain 110 is
mechanically attached under clip 132 (FIG. 9).
[0254] In the illustrated embodiment, the short-term anchoring
component includes a suture strip 134, which is used by a surgeon
to immobilize the sensor against the fascia or other substantially
tough tissue after insertion, such as described in more detail with
reference to FIGS. 12 and 13, below. Alternatively, other
short-term anchoring components that may be used include prongs,
spines, barbs, wings, hooks, helical surface topography, gradually
changing diameter, or the like, which may be disposed on the sensor
body. For example, when an oblong or cylindrical type sensor is
implanted within the subcutaneous tissue, it may tend to slip along
the pocket that was formed during implantation, particularly if
some additional space exists within the pocket. This slippage can
lead to increased inflammatory response and/or movement of the
sensor prior to or during tissue ingrowth. Accordingly, a
short-term anchoring component can aid in immobilizing the sensor
in place, particularly prior to formation of a mature foreign body
capsule.
[0255] Generally, short-term anchoring provides a system and method
for immobilizing the sensor within the soft tissue during the acute
wound-healing phase immediately following the implantation surgery.
While not wishing to be bound by theory, it is believed that the
short-term anchoring component prevents the sensor from movement
within any remaining space in a subcutaneous pocket or space
immediately under the incision, thereby minimizing tissue trauma
and its associated inflammation and foreign body response. By
minimizing tissue trauma, a healthy vascularized tissue bed is more
likely to heal within the biointerface membrane, which is believed
to optimize analyte transport to the sensing region, such as
described in more detail with reference to FIG. 8.
[0256] In the illustrated embodiment, the long-term anchoring
component 136 is an anchoring material. The term "anchoring
material," as used herein, is a broad term and is used in its
ordinary sense, including, without limitation, biocompatible
material or surface that is non-smooth, and particularly comprises
an architecture that supports tissue ingrowth in order to
facilitate anchoring of the material into bodily tissue in vivo.
Some examples of anchoring materials include polyester, velours,
polypropylene cloth, expanded polytetrafluoroethylene, and porous
silicone, for example. However, anchoring material may be built
into the sensor body, for example, by texturizing the non-sensing
region 18 of the analyte sensor 12. In one embodiment, the entire
surface of the sensor is covered with an anchoring material to
provide for strong attachment to the tissues. In another
embodiment, only the sensing side of the sensor incorporates
anchoring material, with the other sides of the sensor lacking
fibers or porous anchoring structures and instead presenting a very
smooth, non-reactive biomaterial surface to prevent attachment to
tissue and to support the formation of a thicker capsule.
[0257] Other configurations may also be suitable for use in certain
embodiments, including configurations with different degrees of
surface coverage. For example, from less than about 5, 10, 15, 20,
25, 30, 35, 40, 45, or 50% to more than about 55, 60, 65, 70, 75,
80, 85, 90, or 95% of the surface of the device may be covered with
anchoring material. The anchoring material may cover one surface,
two surfaces, three surfaces, four surfaces, five surfaces, or six
surfaces. The anchoring material may cover only a portion of one or
more sides, for example, strips, dots, weaves, fibers, meshes, and
other configurations or shapes of anchoring material may cover one
or more sides. It may be noted that the optimum amount of anchoring
material that may be used for any particular sensor is dependent
upon one or more of the following parameters: implantation site
(for example, location in the host), surface area, shape, size,
geometry, mass, density, volume, surface area-to-volume, surface
area-to-density, and surface area-to-mass. For example, a device
with a greater density as compared to a device with a lesser
density may require more anchoring material.
[0258] Generally, long-term anchoring provides a system and method
for immobilizing the sensor long-term in vivo which is believed to
reduce or eliminate the effects of chronic movement on the sensor
(for example, macro- or micro-motion). Immobilization of the sensor
long-term in vivo sustains sensor performance, for example,
sensitivity of the sensor to the analyte, at least in part by
minimizing the inflammation and associated foreign body response
that is known to block analytes from freely diffusing to the
sensing region, for example.
[0259] Implantation--Sizing Tool
[0260] The preferred embodiments employ implantation techniques
that exploit the knowledge gained by the inventors in
experimentation with implantation techniques. In one study,
nineteen glucose sensors were implanted in humans, wherein the
sensors were designed with a cylindrical configuration with the
sensing region on one end thereof. It was observed that of the
nineteen patients participating in the study, acceptable efficacy
was observed for only about half. The study is described in more
detail with reference to co-pending U.S. Provisional Patent
Application 60/460,825. In summary, surgical methods employed in
the clinical study entailed making a 1-inch incision, then forming
a pocket lateral to the incision by blunt dissection. After
placement of the device, a suture was placed by pulling the
connective tissue together at the end of the device proximal to the
incision. During the first several weeks after implantation of the
sensor in the human test subjects, photographs were taken of the
wound site and the position of the device was determined tactilely.
It was observed that eighteen of the nineteen sensors migrated in a
retrograde fashion from the placement site towards the incision
site. Thirteen of these devices moved a significant distance,
namely, a distance of 1 cm or more (device movement of 0.5 cm or
less is not considered to be significant, based on the resolution
of the test measurements utilized). While not wishing to be bound
by theory, it is believed that the sutures did not hold in the
softer, fatty subcutaneous tissues as healing began and wound
contracture formed. This permitted the devices to move into the
virtual space remaining after the formation of the pocket, and in
some cases even permitted the device to migrate into the space
immediately under the incision.
[0261] Although the sensors of the above-described experiment
included layers of porous anchoring materials and biointerface
materials that are designed to be ingrown with tissues, especially
blood vessels, during the wound-healing period; it is well known in
the literature that devices that do not become anchored well in
tissues may become encapsulated by a connective tissue scar much
more aggressively than devices that are well-anchored in tissues.
It is believed that the gross movement of devices that was observed
in the clinical study may have in some cases prevented proper
ingrowth of tissues. Delayed ingrowth of tissues or lack of
ingrowth of tissues may affect device function in a variety of
ways, including lack of glucose sensitivity, delayed start-up of
glucose tracking, and compromised function after start-up,
including low sensitivity and long time lags. In some cases, empty
space left behind after sensor movement has filled with scar tissue
as determined by histological examination of explanted sensors.
[0262] Additionally, it is believed that certain behavioral issues
lead to incomplete healing. Complicating factors included
"fiddling" behaviors (namely, feeling and moving the sensor under
the skin), and the like. These complicating factors were present to
some extent in all of the patients, but were less frequent in
patients with working sensors. Thus, the inventors attribute at
least some sensor performance issue to patient-related
movement.
[0263] Taken together, the inventors identified likely contributing
factors to these performance issues, including: the location of the
sensing region in this experiment on the end of the device may have
led to problems with healing if the sensor moves in such a way that
even a small gap is produced at the end of the sensor; the high
profile of the cylindrical geometry is believed to have caused the
sensor to have a high profile, which makes it easier to bump the
sensor and can lead to the patient touching and feeling the sensor
("fiddling"); and the sensing region can be disrupted easily if
pressure is placed on the opposite end of the device because the
sensor may act as a lever, for example, and rotational energy can
be applied to the sensor, which can also cause disruption of the
sensor-tissue interface.
[0264] Accordingly, the preferred embodiments employ sensor
geometry, short- and/or long-term anchoring, and implantation
techniques that are designed to overcome to the shortcomings
observed in the previously described experiment. For example,
implantation techniques are disclosed that are believed to reduce
the likelihood of device migration, and thereby provide optimized
immobilization and healthy wound healing in vivo. Anchoring
components are disclosed that are believed to substantially
immobilize the device, and thereby provide short- and long-term
stability with minimized trauma. Sensor geometry is disclosed that
reduces fiddling and improves stability due to its low profile and
curved surface(s). All of the above advantages improve analyte
transport through the sensing region and thereby improve long-term
sensor performance.
[0265] FIGS. 11A to 11C are views of a pocket sizing tool in one
preferred embodiment, which has been designed to aid a surgeon in
forming a precisely sized pocket. As described in more detail
above, it has been found that when extra space is formed within
soft tissue during implantation (for example, an oversized pocket),
post-surgical device migration may result, which is believed to
result in suboptimal healing of the tissue into the device.
Suboptimal wound healing is believed to lead to increased thickness
of foreign body capsule formation and increased barrier cell layer
formation, which results in poor analyte transport and sensor
performance in vivo. Accordingly, a pocket-sizing tool is provided,
which may be utilized by the surgeon in precisely forming a
subcutaneous pocket.
[0266] FIG. 11A is a perspective view of the sizing tool 138 in one
preferred embodiment, including a head 140 and a handle 142. FIG.
11B is a side view of the sizing tool 138 showing the offset
placement of the handle 142 on the head 140 in one embodiment. FIG.
11C is a top view of the sizing tool 138 showing a curvature
substantially similar to that of the sensor body (FIG. 6).
[0267] Referring now to the head 140 of the sizing tool, the
preferred embodiments configure the head geometry to be slightly
smaller than that of the analyte sensor geometry to be inserted in
the host. In some preferred embodiments, the head 140a is
dimensioned to be slightly smaller than the dimensions of the
analyte sensor 12 such that the sensor inserted into the pocket is
in compression within the tissue, thereby immobilizing the sensor
body within the subcutaneous pocket. In a preferred embodiment, the
x, y, and z dimensions of the head 140 of the sizing tool are about
0.8 times the dimensions of the analyte sensor. It is appreciated
by one skilled in the art that the dimensions and geometry of the
head 140 of the sizing tool 138 will vary depending on the
dimensions and geometry of the implantable analyte sensor.
[0268] Referring now to the handle 142 of the sizing tool, in this
preferred embodiment, the handle is offset from the center of the
head. This design is based on a variety of contributing factors.
Firstly, the healing within and adjacent to sensing region 16 is
more critical and sensitive than the non-sensing region 18 because
analyte transport optimally requires a healthy and
well-vascularized tissue bed without barrier cell formation
adjacent to the sensing region. Therefore, the offset handle may be
provided for an incision that is spaced from the location wherein
the sensing region of the sensor is to be located after insertion,
which is believed to minimize post-surgical trauma surrounding the
sensing region. It is further noted that one preferred embodiment
provides a short-term anchoring component 134 at a location offset
from and diametrically opposed to the sensing region 16 (FIG. 10C).
In this way, the sensor is adapted to be inserted with the
short-term anchoring component 134 exposed to the surgeon for
suturing to the fascia and diametrically disposed from the sensing
region to minimize trauma associated with the incision and suturing
adjacent the sensing region 16.
[0269] In another preferred embodiment, the handle 142 is centrally
located on the head, which may be advantageous in for a variety of
implantation techniques. For example, when the analyte sensor is
implanted within the fatty tissue of a host, equal sizing on each
size of the incision may provide for ease and accuracy of precise
subcutaneous pocket formation. However, it is noted that the
length, width, thickness, and orientation of the handle may be
optimized for various implantation sites and may be adapted for
patient size, for example. In general, the sizing tool 138 of the
preferred embodiments may be any design that aids a surgeon or
doctor in forming a pocket in the soft tissue with minimized tissue
trauma adjacent to the sensing region 16 and a precisely sized
subcutaneous pocket in a host. It is believed that these factors
immobilize the sensor and provide an opportunity for a well healed
vascularized tissue bed adjacent to the sensing region of the
sensor.
[0270] Implantation--Technique
[0271] The implantable analyte sensor of the preferred embodiments
may be implanted in variety of locations, including: subcutaneous,
intramuscular, intraperotoneal, intrafascial, axillary region, soft
tissue of a body, or the like. Although the preferred embodiments
illustrate implantation within the subcutaneous space of the
abdominal region, the systems and methods described herein are
limited neither to abdominal nor to subcutaneous implantation. One
skilled in the art appreciates that these systems and methods may
be implemented and/or modified for other implantation sites and may
be dependent upon the type, configuration, and dimensions of the
analyte sensor.
[0272] FIG. 12A is a perspective view of the abdominal region 144
of a human, showing the incision 146 and sizing tool 138 that may
be used for surgical implantation. In this embodiment, a skin
incision 146 is made through to the lower plane of the abdominal
subcutaneous fatty layer, preferably avoiding disrupting the muscle
fascia 148 (FIG. 13A). Preferably, a surgeon blunt dissects a
pocket 150 to the precise size of the analyte sensor 12 using the
sizing tool 138 to confirm pocket size. Namely, the surgeon may use
a tool or aid for visualization of the pocket during formation and
incrementally dissects the pocket using the sizing tool 138
frequently to test the length and width of the pocket. As described
in more detail above, it is believed that excess space in the
pocket surrounding a sensor 12 provides an opportunity for
mobilization and migration of the sensor, which results in wounding
and build-up of the thickness of the foreign body capsule and may
encourage barrier cell layer formation adjacent to the sensing
region 16. It is noted that because analyte transport at the
sensing region is critical for sensor performance, the implantation
systems and methods are designed to aid the surgeon in forming a
precise pocket with minimal tissue trauma.
[0273] In one preferred embodiment, the pocket is formed adjacent
to the fascia 148 to provide a tough tissue for short term
anchoring (for example, suturing). In one aspect of this preferred
embodiment, two sutures (FIG. 13A) are placed on the fascia within
the pocket directly under the incision, cranial and caudal to where
the sensor 12 will be placed. These sutures are preferably placed
in a location 152 on the fascia 148 parallel to and as close as
possible to the location of the sensor insertion, optionally using
the sizing tool to estimate proper suture placement. These sutures
are preferably non-resorbable, however resorbable sutures may also
be advantageous in some alternative embodiments. Suturing technique
will be described in more detail with reference to FIG. 13A, below.
Alternatively, other implantation and anchoring techniques are
contemplated that may not benefit from suturing, or may not require
suturing to a tough tissue.
[0274] In another preferred embodiment, the pocket is formed within
the fatty tissue or other soft tissue of the host. These
embodiments minimize invasive dissection as compared with the
above-described embodiment. Namely, by forming the pocket in more
shallow location of the host than the fascia, less dissection is
required, which is believed to induce less tissue trauma.
Additionally, it may be easier for the doctor or surgeon to
visually confirm pocket formation (for example, sizing and
location). It is noted that these embodiments may or may not
include a short-term anchoring mechanism, such as described in more
detail with reference to FIG. 10A.
[0275] FIG. 12B is a perspective view of a portion of the abdominal
region of a human, showing sensor insertion into the precisely
formed pocket in one preferred embodiment. In this embodiment, the
sensor 12 is placed in a lower plane of the subcutaneous fatty
layer as close as possible to the muscle fascia 148 without
disrupting the fascia. Additionally, the sensor 12 is placed with
the long axis perpendicular to and slightly to one side of the
midline 154 of the host. It is noted that in this embodiment, the
sensing region 16 is located off-center and is placed within the
patient with the sensing region facing the fascia 148 and more
distal to the mid-line 154 than the suture strip 134. Thus, in the
illustrated embodiment, the sensor 12 is inserted into the pocket
so that approximately one-third of the sensor is facing medial to
the incision 146 and the other approximately two-thirds is facing
lateral to the incision 146. While not wishing to be bound by
theory, it is believed that because there exists less connective
tissue in this region lateral from the midline 154 than more
proximal to the midline, the sensing region may experience less
trauma associated with the connective tissue dissection. It is
noted, however, that the incision may be centrally or otherwise
located in alternative embodiments, which may advantageously
simplify the implantation procedure, for example.
[0276] FIG. 13A is a schematic view of the sensor after insertion
into the pocket with the sensing region positioned adjacent to the
fascia. Short-term anchoring may be preferable in this embodiment,
in order to immobilize the sensor within the pocket and adjacent to
the fascia for providing optimal opportunity for healthy
vascularized wound healing. However, alternative preferred
embodiments may not include short-term anchoring and/or placement
of the sensor adjacent to the fascia, which is described in more
detail above.
[0277] In the embodiment of FIG. 13A, the sensor body 12 is sutured
to the fascia 148 using the short-term anchoring component, namely
the suture strip 134. Preferably, the suturing technique maintains
the sensor body 12 in compression against the fascia 148 to ensure
minimal movement of the sensing region 16 against the tissue 148.
In one preferred method of suturing the sensor 12 to the fascia
148, the pre-positioned sutures, which are located on the fascia in
a position 152 in line with each side of the suture strip, such as
described with reference to FIGS. 12A and 12B, one of which is
shown in the side view of FIG. 13A. Two additional sutures are
sutured to the suture strip 134, on each side of the suture strip
in a position 156 in line with the pre-positioned sutures 152. The
sutures are then tied together, preferably tying diametrically
opposed sutures to each other, for example the pre-positioned
suture on one side with the suture-strip suture on the opposite
side, and vice versa. In this way, the sutures and the sensor body
are held in compression to ensure immobilization of the sensor 12
within the pocket 150. Alternative methods of suturing the sensor
body to the fascia or other tough tissue are appreciated by one
skilled in the art and are considered within the scope of the
preferred embodiments.
[0278] After short-term anchoring is complete, the incision is
closed; preferably ensuring the subcutaneous tissue is held snugly
around the sensor with resorbable sutures, for example, to minimize
vacant space adjacent to the sensor. The skin incision may be
closed using standard wound closure techniques.
[0279] In one embodiment, the entire surface of the sensor is
covered with an anchoring material to provide for strong attachment
to the tissues. In another embodiment, only the sensor head side of
the sensor incorporates anchoring material, with the other sides of
the sensor lacking fibers or porous anchoring structures and
instead presenting a very smooth, non-reactive biomaterial surface
to prevent attachment to tissue and to support the formation of a
thicker capsule. The anchoring material may be polyester,
polypropylene cloth, polytetrafluoroethylene felts, expanded
polytetrafluoroethylene, porous silicone, or the like.
[0280] FIG. 13B is a side view of an analyte sensor with long-term
anchoring component in the form of an anchoring material 136 on
both sides of the sensor, such as described in more detail with
reference to FIG. 10. In this embodiment, the analyte sensor is
implanted subcutaneously and is ingrown with fibrous, vascularized
tissue after implantation.
[0281] In preferred embodiments, the sensor of the described
geometry is implanted at the interface between two kinds of tissue,
and is preferably anchored to the more robust tissue type. While
the sensor geometries of preferred embodiments are particularly
preferred for use at tissue interfaces, such sensors are also
suitable for use when implanted into a single type of tissue, for
example, muscle tissue or adipose tissue. In one alternative
embodiment, the sensor may be suspended, with or without sutures,
in a single tissue type, or be placed between two tissue types, and
anchoring material covering substantially the entire surface of the
device may be sufficient to prevent device migration, such as
described elsewhere herein.
[0282] FIG. 13A illustrates the surface of the sensor 12 in
mechanical contact with the overlying tissue 158, as well as the
underlying muscle fascia 148, due to the ingrowth of the fibrous
tissue and vasculature. In this embodiment, any surface of the
sensor 12 covered with anchoring material 136 is typically ingrown
with fibrous, vascularized tissue 160, which aids in anchoring the
sensor and mitigating motion artifact. It may be noted however,
that in some cases, forces applied laterally to this tissue may be
translated to the sensor, and likewise to the fascia side of the
sensor, causing potential disruption of the interface with the
fascia. Therefore, although the radial profile of the side of the
sensor incorporating the sensing region assists in preventing
forces in the distal subcutaneous tissue from exerting forces on
the sensor head side, which is attached to the muscle fascia by an
anchoring material, complete coverage of the device with anchoring
material may not be preferred in certain embodiments.
[0283] It may be noted that other factors of the preferred
embodiments aid in immobilizing the sensor within the host, for
example the sensor geometry, such as described in more detail with
reference to FIG. 6. Additionally, the inventors have designed the
sensor with a "low profile" to reduce the possibility of "fiddling"
behavior that was seen in the above-described experiment. "Low
profile" is loosely defined as an overall configuration, including
dimensions, shape, and aspect ratio, that is deliberately
inconspicuous when implanted.
[0284] Additionally, the preferred embodiments describe a
relatively thin, rectangular or oval sensor wherein the sensing
region is positioned on one of the large sides of the sensor (for
example, rather than at the tip). When implanted, the sensor is
preferably oriented such that the sensing region is adjacent to the
muscle fascia underlying the subcutaneous space; however other
locations are also possible.
[0285] Taken together, the preferred sensor substantially does not
protrude through the host's skin (which may be somewhat dependent
upon the host's body fat) and is less amenable to accidental
bumping or movement, and less available for patient "fiddling." It
is a thin, oblong shape, not cylindrical, so rotational forces are
not as likely to affect the sensor-tissue interface. With the
sensing region oriented down towards the fascia, and nearer to the
center of the sensor, downward pressure on either end is not
transferred as shear force to the sensing region.
[0286] In general, miniaturization of the sensor (for example, size
and mass) prevents motion of the device and may reduce the time for
sensor start-up by minimizing the disruption of the fragile new
capillaries necessary to supply analytes (for example, oxygen and
glucose) to the sensor. Miniaturization is also believed to reduce
the vascular compression/postural effects that have been found to
block the transport of oxygen when the host is positioned in such a
way as to block proper transport of analytes to the sensing region,
for example.
[0287] Methods and devices that are suitable for use in conjunction
with aspects of the preferred embodiments are disclosed in
copending U.S. patent application Ser. No. 10/789,359 filed Feb.
26, 2004 and entitled, "INTEGRATED DELIVERY DEVICE FOR CONTINUOUS
GLUCOSE SENSOR"; U.S. application Ser. No. 10/685,636 filed Oct.
28, 2003 and entitled, "SILICONE COMPOSITION FOR BIOCOMPATIBLE
MEMBRANE"; U.S. application Ser. No. 10/648,849 filed Aug. 22, 2003
and entitled, "SYSTEMS AND METHODS FOR REPLACING SIGNAL ARTIFACTS
IN A GLUCOSE SENSOR DATA STREAM"; U.S. application Ser. No.
10/646,333 filed Aug. 22, 2003 entitled, "OPTIMIZED SENSOR GEOMETRY
FOR AN IMPLANTABLE GLUCOSE SENSOR"; U.S. application Ser. No.
10/647,065 filed Aug. 22, 2003 entitled, "POROUS MEMBRANES FOR USE
WITH IMPLANTABLE DEVICES"; U.S. application Ser. No. 10/633,367
filed Aug. 1, 2003 entitled, "SYSTEM AND METHODS FOR PROCESSING
ANALYTE SENSOR DATA"; U.S. application Ser. No. 09/916,386 filed
Jul. 27, 2001 and entitled "MEMBRANE FOR USE WITH IMPLANTABLE
DEVICES"; U.S. application Ser. No. 09/916,711 filed Jul. 27, 2001
and entitled "SENSOR HEAD FOR USE WITH IMPLANTABLE DEVICE"; U.S.
application Ser. No. 09/447,227 filed Nov. 22, 1999 and entitled
"DEVICE AND METHOD FOR DETERMINING ANALYTE LEVELS"; U.S.
application Ser. No. 10/153,356 filed May 22, 2002 and entitled
"TECHNIQUES TO IMPROVE POLYURETHANE MEMBRANES FOR IMPLANTABLE
GLUCOSE SENSORS"; U.S. application Ser. No. 09/489,588 filed Jan.
21, 2000 and entitled "DEVICE AND METHOD FOR DETERMINING ANALYTE
LEVELS"; U.S. application Ser. No. 09/636,369 filed Aug. 11, 2000
and entitled "SYSTEMS AND METHODS FOR REMOTE MONITORING AND
MODULATION OF MEDICAL DEVICES"; and U.S. application Ser. No.
09/916,858 filed Jul. 27, 2001 and entitled "DEVICE AND METHOD FOR
DETERMINING ANALYTE LEVELS," as well as issued patents including
U.S. Pat. No. 6,001,067 issued Dec. 14, 1999 and entitled "DEVICE
AND METHOD FOR DETERMINING ANALYTE LEVELS"; U.S. Pat. No. 4,994,167
issued Feb. 19, 1991 and entitled "BIOLOGICAL FLUID MEASURING
DEVICE"; and U.S. Pat. No. 4,757,022 filed Jul. 12, 1988 and
entitled "BIOLOGICAL FLUID MEASURING DEVICE." The foregoing patent
applications and patents are incorporated herein by reference in
their entireties.
[0288] The above description discloses several methods and
materials of the present invention. This invention is susceptible
to modifications in the methods and materials, as well as
alterations in the fabrication methods and equipment. Such
modifications will become apparent to those skilled in the art from
a consideration of this disclosure or practice of the invention
disclosed herein. Consequently, it is not intended that this
invention be limited to the specific embodiments disclosed herein,
but that it cover all modifications and alternatives coming within
the true scope and spirit of the invention as embodied in the
attached claims. All patents, applications, and other references
cited herein, are hereby incorporated by reference in their
entirety.
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