U.S. patent application number 15/967338 was filed with the patent office on 2018-08-30 for analyte sensor.
The applicant listed for this patent is DexCom, Inc.. Invention is credited to Robert J. Boock, James H. Brauker, Mark C. Brister, Monica A. Rixman, Mark C. Shults, Peter C. Simpson.
Application Number | 20180242894 15/967338 |
Document ID | / |
Family ID | 37695272 |
Filed Date | 2018-08-30 |
United States Patent
Application |
20180242894 |
Kind Code |
A1 |
Brauker; James H. ; et
al. |
August 30, 2018 |
ANALYTE SENSOR
Abstract
Biointerface membranes are provided which can be utilized with
implantable devices, such as devices for the detection of analyte
concentrations in a biological sample. More particularly, methods
for monitoring glucose levels in a biological fluid sample using an
implantable analyte detection device incorporating such membranes
are provided.
Inventors: |
Brauker; James H.;
(Coldwater, MI) ; Boock; Robert J.; (Carlsbad,
CA) ; Rixman; Monica A.; (Medford, MA) ;
Simpson; Peter C.; (Cardiff, CA) ; Brister; Mark
C.; (Encinitas, CA) ; Shults; Mark C.;
(Madison, WI) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
DexCom, Inc. |
San Diego |
CA |
US |
|
|
Family ID: |
37695272 |
Appl. No.: |
15/967338 |
Filed: |
April 30, 2018 |
Related U.S. Patent Documents
|
|
|
|
|
|
Application
Number |
Filing Date |
Patent Number |
|
|
11503367 |
Aug 10, 2006 |
9986942 |
|
|
15967338 |
|
|
|
|
11439630 |
May 23, 2006 |
10022078 |
|
|
11503367 |
|
|
|
|
11077715 |
Mar 10, 2005 |
7497827 |
|
|
11439630 |
|
|
|
|
60587787 |
Jul 13, 2004 |
|
|
|
60587800 |
Jul 13, 2004 |
|
|
|
60614683 |
Sep 30, 2004 |
|
|
|
60614764 |
Sep 30, 2004 |
|
|
|
60683923 |
May 23, 2005 |
|
|
|
Current U.S.
Class: |
1/1 |
Current CPC
Class: |
Y02A 50/58 20180101;
A61B 2562/08 20130101; A61B 5/14865 20130101; A61B 5/14532
20130101; A61B 5/1473 20130101; G01N 33/54366 20130101; C12Q 1/006
20130101; A61B 5/6848 20130101; A61B 5/411 20130101; G01N 33/66
20130101; A61B 5/418 20130101; Y02A 50/30 20180101 |
International
Class: |
A61B 5/1486 20060101
A61B005/1486; G01N 33/66 20060101 G01N033/66; A61B 5/145 20060101
A61B005/145; C12Q 1/00 20060101 C12Q001/00; A61B 5/00 20060101
A61B005/00; G01N 33/543 20060101 G01N033/543 |
Claims
1. An analyte sensing device adapted for implantation into a host's
tissue, comprising: a sensor configured to measure an analyte in a
host, wherein the sensor comprises a biointerface configured to
promote at least one function selected from the group consisting of
increasing fluid bulk surrounding at least a portion of the sensor
in vivo, increasing bulk fluid flow surrounding at least a portion
of the sensor in vivo, and increasing diffusion rates surrounding
at least a portion of the sensor in vivo.
2. The device of claim 1, wherein the biointerface comprises a
spacer.
3. The device of claim 2, wherein the spacer comprises a mesh.
4. The device of claim 2, wherein the spacer comprises a
hydrogel.
5. The device of claim 4, wherein the hydrogel comprises from about
20 wt. % to about 99 wt. % water.
6. The device of claim 4, wherein the hydrogel comprises from about
80 wt. % to about 99 wt. % water.
7. The device of claim 2, wherein the spacer comprises a shedding
layer.
8. The device of claim 2, wherein the spacer is a fibrous
structure.
9. The device of claim 2, wherein the spacer is a porous polymer
membrane.
10. The device of claim 2, wherein the spacer comprises a material
selected from the group consisting of polysulfone,
polytetrafluoroethylene, polyvinylidene difluoride,
polyacrylonitrile, silicone, polytetrafluoroethylene, expanded
polytetrafluoroethylene, polyethylene-co-tetrafluoroethylene,
polyolefin, polyester, polycarbonate, biostable
polytetrafluoroethylene, polyurethane, polypropylene,
polyvinylchloride, polyvinylidene fluoride, polyvinyl alcohol,
polybutylene terephthalate, polymethylmethacrylate, polyether ether
ketone, polyamides, cellulosic polymer, poly(ethylene oxide),
poly(propylene oxide), hydrogel polymer, poly(2-hydroxyethyl
methacrylate), hydroxyethyl methacrylate, high density
polyethylene, acrylic copolymer, nylon, polyvinyl difluoride,
polyanhydride, poly(l-lysine), poly (L-lactic acid),
hydroxyethylmethacrylate, homopolymers thereof, copolymers thereof,
di-block copolymers thereof, tri-block copolymers thereof,
alternating copolymers thereof, random copolymers thereof, graft
copolymers thereof, terpolymers thereof, and blends thereof.
11. The device of claim 2, wherein the spacer comprises a material
selected from the group consisting of metal, ceramic,
hydroxyapeptite, alumina, zirconia, carbon fiber, aluminum, calcium
phosphate, titanium, titanium alloy, nintinol, stainless steel,
CoCr alloy, and combinations thereof.
12. The device of claim 2, wherein the spacer has an average
nominal pore size of from about 0.6 .mu.m to about 20 .mu.m.
13. The device of claim 2, wherein at least 50% of the pores of the
spacer have an average size of from about 0.6 .mu.m to about 20
.mu.m.
14. The device of claim 1, wherein the biointerface is configured
to provide a fluid pocket.
15. The device of claim 1, wherein the biointerface comprises a
roughened surface.
16. The device of claim 15, wherein the roughened surface is a
vasodilating surface.
17. The device of claim 1, wherein the biointerface comprises an
irregular surface.
18. The device of claim 1, wherein the biointerface comprises a
nanoporous material, a swellable material, or a collapsible
material.
19. The device of claim 1, wherein the biointerface comprises an
irritating superstructure.
20. The device of claim 19, wherein the irritating superstructure
comprises a coiled silver wire.
Description
INCORPORATION BY REFERENCE TO RELATED APPLICATIONS
[0001] Any and all priority claims identified in the Application
Data Sheet, or any correction thereto, are hereby incorporated by
reference under 37 CFR 1.57. This application is a continuation of
U.S. application Ser. No. 11/503,367, filed Aug. 10, 2006, which is
a continuation-in-part of U.S. application Ser. No. 11/439,630,
filed May 23, 2006, which is a continuation-in-part of U.S.
application Ser. No. 11/077,715, filed Mar. 10, 2005, now U.S. Pat.
No. 7,497,827, which claims the benefit of U.S. Provisional
Application No. 60/587,787 filed Jul. 13, 2004; U.S. Provisional
Application No. 60/587,800 filed Jul. 13, 2004; U.S. Provisional
Application No. 60/614,683 filed Sep. 30, 2004; and U.S.
Provisional Application No. 60/614,764 filed Sep. 30, 2004. U.S.
application Ser. No. 11/439,630 claims the benefit of U.S.
Provisional Application No. 60/683,923, filed May 23, 2005. Each of
the aforementioned applications is incorporated by reference herein
in its entirety, and each is hereby expressly made a part of this
specification.
FIELD OF THE INVENTION
[0002] Biointerface membranes are provided which can be utilized
with implantable devices, such as devices for the detection of
analyte concentrations in a biological sample. More particularly,
methods for monitoring glucose levels in a biological fluid sample
using an implantable analyte detection device incorporating such
membranes are provided.
BACKGROUND OF THE INVENTION
[0003] One of the most heavily investigated analyte sensing devices
is the implantable glucose device for detecting glucose levels in
hosts with diabetes. Despite the increasing number of individuals
diagnosed with diabetes and recent advances in the field of
implantable glucose monitoring devices, currently used devices are
unable to provide data safely and reliably for certain periods of
time. There are two commonly used types of subcutaneously
implantable glucose sensing devices. These types include those that
are implanted transcutaneously and those that are wholly
implanted.
SUMMARY OF THE INVENTION
[0004] Accordingly, in a first aspect, an analyte sensing device
adapted for implantation into a host's tissue is provided,
comprising a sensor configured to measure an analyte in a host,
wherein the sensor comprises a biointerface configured to promote
at least one function selected from the group consisting of
increasing fluid bulk surrounding at least a portion of the sensor
in vivo, increasing bulk fluid flow surrounding at least a portion
of the sensor in vivo, and increasing diffusion rates surrounding
at least a portion of the sensor in vivo.
[0005] In an embodiment of the first aspect, the biointerface
comprises a spacer.
[0006] In an embodiment of the first aspect, the spacer comprises a
mesh.
[0007] In an embodiment of the first aspect, the spacer comprises a
hydrogel.
[0008] In an embodiment of the first aspect, the hydrogel comprises
from about 20 wt. % to about 99 wt. % water.
[0009] In an embodiment of the first aspect, the hydrogel comprises
from about 80 wt. % to about 99 wt. % water.
[0010] In an embodiment of the first aspect, the spacer comprises a
shedding layer.
[0011] In an embodiment of the first aspect, the spacer is a
fibrous structure.
[0012] In an embodiment of the first aspect, the spacer is a porous
polymer membrane.
[0013] In an embodiment of the first aspect, the spacer comprises a
material selected from the group consisting of polysulfone,
polytetrafluoroethylene, polyvinylidene difluoride,
polyacrylonitrile, silicone, polytetrafluoroethylene, expanded
polytetrafluoroethylene, polyethylene-co-tetrafluoroethylene,
polyolefin, polyester, polycarbonate, biostable
polytetrafluoroethylene, polyurethane, polypropylene,
polyvinylchloride, polyvinylidene fluoride, polyvinyl alcohol,
polybutylene terephthalate, polymethylmethacrylate, polyether ether
ketone, polyamides, cellulosic polymer, poly(ethylene oxide),
poly(propylene oxide), hydrogel polymer, poly(2-hydroxyethyl
methacrylate), hydroxyethyl methacrylate, high density
polyethylene, acrylic copolymer, nylon, polyvinyl difluoride,
polyanhydride, poly(l-lysine), poly (L-lactic acid),
hydroxyethylmethacrylate, homopolymers thereof, copolymers thereof,
di-block copolymers thereof, tri-block copolymers thereof,
alternating copolymers thereof, random copolymers thereof, graft
copolymers thereof, terpolymers thereof, and blends thereof.
[0014] In an embodiment of the first aspect, the spacer comprises a
material selected from the group consisting of metal, ceramic,
hydroxyapeptite, alumina, zirconia, carbon fiber, aluminum, calcium
phosphate, titanium, titanium alloy, nintinol, stainless steel,
CoCr alloy, and combinations thereof.
[0015] In an embodiment of the first aspect, the spacer has an
average nominal pore size of from about 0.6 .mu.m to about 20
.mu.m.
[0016] In an embodiment of the first aspect, at least 50% of the
pores of the spacer have an average size of from about 0.6 .mu.m to
about 20 .mu.m.
[0017] In an embodiment of the first aspect, the biointerface is
configured to provide a fluid pocket.
[0018] In an embodiment of the first aspect, the biointerface
comprises a roughened surface.
[0019] In an embodiment of the first aspect, the roughened surface
is a vasodilating surface.
[0020] In an embodiment of the first aspect, the biointerface
comprises an irregular surface.
[0021] In an embodiment of the first aspect, the biointerface
comprises a nanoporous material, a swellable material, or a
collapsible material.
[0022] In an embodiment of the first aspect, the biointerface
comprises an irritating superstructure.
[0023] In an embodiment of the first aspect, the irritating
superstructure comprises a coiled silver wire.
[0024] In an embodiment of the first aspect, the biointerface
comprises a biodegradable material.
[0025] In an embodiment of the first aspect, the biodegradable
material is a biodegradable polymer.
[0026] In an embodiment of the first aspect, the biodegradable
polymer comprises an irritating polymer.
[0027] In an embodiment of the first aspect, the spacer comprises a
self-assembling material.
[0028] In an embodiment of the first aspect, the self-assembling
material comprises a self-assembling peptide.
[0029] In an embodiment of the first aspect, the biointerface
comprises a bioactive agent.
[0030] In an embodiment of the first aspect, the bioactive agent is
selected from the group consisting of anti-barrier cell agent, an
anti-infective agent, a necrosing agent, an inflammatory agent, a
growth factor, an angiogenic factor, an adjuvant, an antiplatelet
agent, an anticoagulant, an ACE inhibitor, a cytotoxic agent, a
vascularization compound, an anti-sense molecule, an enzyme, a
metal, a hydrophilic biodegradable polymer, a glycolic acid-based
polymer, a lactic acid-based polymer, polyethylene oxide, silver,
and combinations thereof.
[0031] In an embodiment of the first aspect, the sensor is
configured to measure a signal that is indicative of a quantity of
the analyte within a fluid surrounding at least a portion of the
sensor.
[0032] In an embodiment of the first aspect, the fluid surrounding
at least a portion of the sensor comprises wound fluid.
[0033] In an embodiment of the first aspect, the device further
comprises electronics operably connected to the sensor and adapted
for detecting a signal from the sensor, wherein the signal is
indicative of a quantity of analyte within the host.
[0034] In an embodiment of the first aspect, the device further
comprises a housing adapted for placement adjacent to the host's
skin, wherein at least a portion of the electronics are disposed in
the housing.
[0035] In an embodiment of the first aspect, the sensor is adapted
for short-term implantation.
[0036] In an embodiment of the first aspect, the sensor is a
transcutaneous sensor.
[0037] In a second aspect, an analyte sensing device adapted for
implantation into a host's tissue is provided, comprising a sensor
for measuring an analyte in the host, wherein the sensor comprises
a biointerface configured to irritate a surrounding in vivo
environment.
[0038] In an embodiment of the second aspect, the biointerface
comprises a shedding layer.
[0039] In an embodiment of the second aspect, the biointerface
comprises a roughened surface.
[0040] In an embodiment of the second aspect, the biointerface
comprises an irritating superstructure.
[0041] In an embodiment of the second aspect, the irritating
superstructure comprises a coiled silver wire.
[0042] In an embodiment of the second aspect, the biointerface
comprises an irregular surface.
[0043] In an embodiment of the second aspect, the biointerface
comprises a biodegradable material.
[0044] In an embodiment of the second aspect, the biodegradable
material is a biodegradable polymer.
[0045] In an embodiment of the second aspect, the biodegradable
polymer comprises an irritating polymer.
[0046] In an embodiment of the second aspect, the biointerface
comprises a bioactive agent.
[0047] In an embodiment of the second aspect, the bioactive agent
is selected from the group consisting of an anti-barrier cell
agent, an anti-infective agent, a necrosing agent, an inflammatory
agent, a growth factor, an angiogenic factor, an adjuvant, an
antiplatelet agent, an anticoagulant, an ACE inhibitor, a cytotoxic
agent, a vascularization compound, an anti-sense molecule, an
enzyme, a metal, a hydrophilic biodegradable polymer, a glycolic
acid-based polymer, a lactic acid-based polymer, polyethylene
oxide, silver, and combinations thereof.
[0048] In an embodiment of the second aspect, the sensor is
configured to measure a signal that is indicative of a quantity of
the analyte within a fluid surrounding at least a portion of the
sensor.
[0049] In an embodiment of the second aspect, the fluid surrounding
at least a portion of the sensor comprises wound fluid.
[0050] In an embodiment of the second aspect, the device further
comprises electronics operably connected to the sensor and adapted
for detecting a signal from the sensor, wherein the signal is
indicative of a quantity of the analyte within the host.
[0051] In an embodiment of the second aspect, the device further
comprises a housing adapted for placement adjacent to the host's
skin, wherein at least a portion of the electronics are disposed in
the housing.
[0052] In an embodiment of the second aspect, the sensor is adapted
for short-term implantation
[0053] In an embodiment of the second aspect, the sensor is a
transcutaneous sensor.
[0054] In a third aspect, an analyte sensing device adapted for
implantation into a host's tissue is provided, comprising a sensor
for measuring an analyte in a host, wherein the sensor comprises a
biointerface configured to suppress wound healing around at least a
portion of the sensor in vivo.
[0055] In an embodiment of the third aspect, the biointerface
comprises a scavenging agent.
[0056] In an embodiment of the third aspect, the biointerface
comprises a bioactive agent.
[0057] In an embodiment of the third aspect, the bioactive agent is
selected from the group consisting of an anti-inflammatory agent,
an anti-infective agent, an anesthetic, a growth factor, an
angiogenic factor, an immunosuppressive agent, an antiplatelet
agent, an anticoagulant, a scavenging agent, an anti-histamine, and
combinations thereof.
[0058] In an embodiment of the third aspect, the bioactive agent
comprises an anti-histamine.
[0059] In an embodiment of the third aspect, the biointerface
comprises an architecture configured to suppress wounding.
[0060] In an embodiment of the third aspect, the biointerface
comprises an anti-inflammatory architecture.
[0061] In an embodiment of the third aspect, the biointerface
comprises a pro-inflammatory architecture.
[0062] In an embodiment of the third aspect, the biointerface
comprises an artificial protective coating.
[0063] In an embodiment of the third aspect, the artificial
protective coating comprises a substance selected from the group
consisting of albumin, fibrin, collagen, endothelial cells, wound
closure chemicals, blood products, platelet-rich plasma, growth
factors, and combinations thereof.
[0064] In an embodiment of the third aspect, the sensor is
configured to measure a signal that is indicative of a quantity of
the analyte within a fluid surrounding at least a portion of the
sensor.
[0065] In an embodiment of the third aspect, the fluid surrounding
at least a portion of the sensor comprises wound fluid.
[0066] In an embodiment of the third aspect, the device further
comprises electronics operably connected to the sensor and adapted
for detecting a signal from the sensor, wherein the signal is
indicative of a quantity of the analyte within the host.
[0067] In an embodiment of the third aspect, the device further
comprises a housing adapted for placement adjacent to the host's
skin, wherein at least a portion of the electronics are disposed in
the housing.
[0068] In an embodiment of the third aspect, the sensor is adapted
for short-term implantation
[0069] In an embodiment of the third aspect, the sensor is a
transcutaneous sensor.
[0070] In a fourth aspect, a method for detecting an analyte in a
host is provided, comprising providing an analyte sensing device
adapted for transcutaneous insertion into the host, the device
comprising a sensor for measuring the analyte in the host, wherein
the sensor is configured to reduce noise in vivo; inserting the
sensor through the host's skin and into the host; waiting a first
period of time, during which first period of time the sensor
remains in the host, wherein the first period of time is sufficient
for at least partial wound healing to occur; initiating a sensor
function; and detecting a signal from the sensor, wherein the
signal is indicative of a concentration of an analyte in the
host.
[0071] In an embodiment of the fourth aspect, the first time period
is at least about 1 hour.
[0072] In an embodiment of the fourth aspect, the first time period
is at least about 24 hours.
[0073] In an embodiment of the fourth aspect, the first period of
time is from about 1 hour to about 48 hours.
[0074] In an embodiment of the fourth aspect, the method further
comprises a step of waiting a second period of time during which
the sensor remains in the host, wherein the step of waiting a
second period of time is conducted after the step of initiating a
sensor function and before the step of detecting a signal from the
sensor.
[0075] In an embodiment of the fourth aspect, the second period of
time is at least about 1 hour.
[0076] In an embodiment of the fourth aspect, the second period of
time is at least about 24 hours.
[0077] In an embodiment of the fourth aspect, the second period of
time is from about 1 hour to about 48 hours.
BRIEF DESCRIPTION OF THE DRAWINGS
[0078] FIG. 1A is a graph of intermittent, sedentary noise in a
non-diabetic host wearing a STS glucose sensor. The upper line
shows the sensor signal. The lower line shows the noise within the
sensor signal.
[0079] FIG. 1B is a graph illustrating nighttime noise in a
non-diabetic host wearing a STS glucose sensor built without
enzyme. The black line shows the sensor signal from the sensor
without enzyme.
[0080] FIG. 1C is a graph comparing glucose measurements from blood
samples collected from the lower abdomen (diamonds, dashed line)
and the fingertip (squares, solid line) using a lancet, in a normal
host that has high levels of nighttime noise. Measurements were
made with a hand-held glucose monitor.
[0081] FIG. 1D is a graph comparing signals from samples collected
from the lower abdomen (diamonds, dashed line) and the fingertip
(squares, solid line) using a lancet, in a normal host that has low
levels of nighttime noise. Measurements were made with a hand-held
glucose monitor.
[0082] FIG. 1E is a photograph of an approximately 3-inch portion
of the abdomen (where samples were collected) of the host of FIG.
1C.
[0083] FIG. 1F is a photo of the index and middle fingers (where
samples were collected) of the host of FIG. 1C.
[0084] FIG. 2A is an illustration of classical three-layered
foreign body response to a conventional synthetic membrane
implanted under the skin.
[0085] FIG. 2B is a side schematic view of adipose cell contact
with an inserted transcutaneous sensor or an implanted sensor.
[0086] FIG. 2C is a side schematic view of a biointerface membrane
preventing adipose cell contact with an inserted transcutaneous
sensor or an implanted sensor.
[0087] FIG. 3A is an expanded view of an exemplary embodiment of a
continuous analyte sensor.
[0088] FIG. 3B is a cross-sectional view through the sensor of FIG.
3A on line B-B.
[0089] FIG. 4A is a side schematic view of a transcutaneous analyte
sensor in one embodiment.
[0090] FIG. 4B is a side schematic view of a transcutaneous analyte
sensor in an alternative embodiment.
[0091] FIG. 4C is a side schematic view of a wholly implantable
analyte sensor in one embodiment.
[0092] FIG. 4D is a side schematic view of a wholly implantable
analyte sensor in an alternative embodiment.
[0093] FIG. 4E is a side schematic view of a wholly implantable
analyte sensor in another alternative embodiment.
[0094] FIG. 4F is a side view of one embodiment of an implanted
sensor inductively coupled to an electronics unit within a
functionally useful distance on the host's skin.
[0095] FIG. 4G is a side view of one embodiment of an implanted
sensor inductively coupled to an electronics unit implanted in the
host's tissue at a functionally useful distance.
[0096] FIG. 5A is a cross-sectional schematic view of a membrane of
a preferred embodiment that facilitates vascularization of the
first domain without barrier cell layer formation.
[0097] FIG. 5B is a cross-sectional schematic view of the membrane
of FIG. 5A showing contractile forces caused by the fibrous tissue
of the FBR.
[0098] FIG. 6 is a flow chart that illustrates the process of
forming a biointerface-coated small structured sensor in one
embodiment.
[0099] FIG. 7 is a flow chart that illustrates the process of
forming a biointerface-coated sensor in an alternative
embodiment.
[0100] FIG. 8 is a flow chart that illustrates the process of
forming a biointerface-coated sensor in another alternative
embodiment.
[0101] FIG. 9 is a flow chart that illustrates the process of
forming a biointerface-wrapped sensor in one embodiment.
[0102] FIG. 10 is a flow chart that illustrates the process of
forming a sensing biointerface in one embodiment.
[0103] FIG. 11A is a scanning electron micrograph showing a
cross-sectional view of a cut porous silicone tube. The scale line
equals 500 .mu.m.
[0104] FIG. 11B is a scanning electron micrograph of a sugar mold
formed on a sensor, prior to silicone application. The scale line
equals 100 .mu.m.
DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENT
[0105] The following description and examples illustrate a
preferred embodiment of the present 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 preferred embodiment
should not be deemed to limit the scope of the present
invention.
Definitions
[0106] In order to facilitate an understanding of the preferred
embodiment, a number of terms are defined below.
[0107] The term "biointerface" as used herein is a broad term, and
is to be given its ordinary and customary meaning to a person of
ordinary skill in the art (and is not to be limited to a special or
customized meaning), and refers without limitation to any structure
or substance that interfaces between host (tissue or body fluid)
and an implantable device.
[0108] The term "biointerface membrane" as used herein is a broad
term, and is to be given its ordinary and customary meaning to a
person of ordinary skill in the art (and is not to be limited to a
special or customized meaning), and refers without limitation to a
membrane that functions as an interface between host (tissue or
body fluid) and an implantable device.
[0109] The term "interface" as used herein is a broad term, and is
to be given its ordinary and customary meaning to a person of
ordinary skill in the art (and is not to be limited to a special or
customized meaning), and refers without limitation to 1) a common
boundary, such as the surface, place, or point where two things
touch each other or meet, or 2) a point of interaction, including
the place, situation, or way in which two things act together or
affect each other, or the point of connection between things.
[0110] The term "barrier cell layer" as used herein is a broad
term, and is to be given its ordinary and customary meaning to a
person of ordinary skill in the art (and is not to be limited to a
special or customized meaning), and refers without limitation to a
part of a foreign body response that forms a cohesive monolayer of
cells (for example, macrophages and foreign body giant cells) that
substantially block the transport of molecules and other substances
to the implantable device.
[0111] The term "cell processes" as used herein is a broad term,
and is to be given its ordinary and customary meaning to a person
of ordinary skill in the art (and is not to be limited to a special
or customized meaning), and refers without limitation to
pseudopodia of a cell.
[0112] The term "cellular attachment" as used herein is a broad
term, and is to be given its ordinary and customary meaning to a
person of ordinary skill in the art (and is not to be limited to a
special or customized meaning), and refers without limitation to
adhesion of cells and/or cell processes to a material at the
molecular level, and/or attachment of cells and/or cell processes
to microporous material surfaces or macroporous material surfaces.
One example of a material used in the prior art that encourages
cellular attachment to its porous surfaces is the BIOPORE.TM. cell
culture support marketed by Millipore (Bedford, Mass.), and as
described in Brauker et al., U.S. Pat. No. 5,741,330.
[0113] The term "solid portions" as used herein is a broad term,
and is to be given its ordinary and customary meaning to a person
of ordinary skill in the art (and is not to be limited to a special
or customized meaning), and refers without limitation to portions
of a membrane's material having a mechanical structure that
demarcates cavities, voids, pores, or other non-solid portions.
[0114] The term "co-continuous" as used herein is a broad term, and
is to be given its ordinary and customary meaning to a person of
ordinary skill in the art (and is not to be limited to a special or
customized meaning), and refers without limitation to a solid
portion or cavity or pore wherein an unbroken curved line in three
dimensions can be drawn between two sides of a membrane.
[0115] The term "biostable" as used herein is a broad term, and is
to be given its ordinary and customary meaning to a person of
ordinary skill in the art (and is not to be limited to a special or
customized meaning), and refers without limitation to materials
that are relatively resistant to degradation by processes that are
encountered in vivo.
[0116] The terms "bioresorbable" or "bioabsorbable" as used herein
are broad terms, and are to be given their ordinary and customary
meaning to a person of ordinary skill in the art (and are not to be
limited to a special or customized meaning), and refer without
limitation to materials that can be absorbed, or lose substance, in
a biological system.
[0117] The terms "nonbioresorbable" or "nonbioabsorbable" as used
herein are broad terms, and are to be given their ordinary and
customary meaning to a person of ordinary skill in the art (and are
not to be limited to a special or customized meaning), and refer
without limitation to materials that are not substantially
absorbed, or do not substantially lose substance, in a biological
system.
[0118] The term "analyte" as used herein is a broad term, and is to
be given its ordinary and customary meaning to a person of ordinary
skill in the art (and is not to be limited to a special or
customized meaning), and refers without limitation 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- 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, hemoglobin A, hemoglobin S,
hemoglobin C, hemoglobin D, hemoglobin E, hemoglobin F, 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 -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, ); 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), 5-hydroxyindoleacetic acid (FHIAA), and
histamine.
[0119] The term "host" as used herein is a broad term, and is to be
given its ordinary and customary meaning to a person of ordinary
skill in the art (and is not to be limited to a special or
customized meaning), and refers without limitation to mammals,
preferably humans.
[0120] The phrase "continuous analyte sensing" as used herein is a
broad term, and is to be given its ordinary and customary meaning
to a person of ordinary skill in the art (and is not to be limited
to a special or customized meaning), and refers without limitation
to the period in which monitoring of analyte concentration is
continuously, continually, and/or intermittently (but regularly)
performed, for example, from about every 5 seconds or less to about
10 minutes or more, preferably from about 10, 15, 20, 25, 30, 35,
40, 45, 50, 55, or 60 second to about 1.25, 1.50, 1.75, 2.00, 2.25,
2.50, 2.75, 3.00, 3.25, 3.50, 3.75, 4.00, 4.25, 4.50, 4.75, 5.00,
5.25, 5.50, 5.75, 6.00, 6.25, 6.50, 6.75, 7.00, 7.25, 7.50, 7.75,
8.00, 8.25, 8.50, 8.75, 9.00, 9.25, 9.50 or 9.75 minutes.
[0121] The terms "analyte measuring device," "sensor," "sensing
region," and "sensing mechanism" as used herein are broad terms,
and are to be given their ordinary and customary meaning to a
person of ordinary skill in the art (and are not to be limited to a
special or customized meaning), and refer without limitation to an
area of an analyte-monitoring device that enables the detection of
a particular analyte. For example, the sensing region can comprise
a non-conductive body, a working electrode, a reference electrode,
and a counter electrode (optional), forming an electrochemically
reactive surface at one location on the body and an electronic
connection at another location on the body, and a sensing membrane
affixed to the body and covering the electrochemically reactive
surface. During general operation of the device, a biological
sample, for example, blood or interstitial fluid, or a component
thereof contacts, either directly or after passage through one or
more membranes, an enzyme, for example, glucose oxidase. The
reaction of the biological sample or component thereof results in
the formation of reaction products that permit a determination of
the analyte level, for example, glucose, in the biological sample.
In some embodiments, the sensing membrane further comprises an
enzyme domain, for example, an enzyme layer, and an electrolyte
phase, for example, a free-flowing liquid phase comprising an
electrolyte-containing fluid described further below. The terms are
broad enough to include the entire device, or only the sensing
portion thereof (or something in between).
[0122] The term "electrochemically reactive surface" as used herein
is a broad term, and is to be given its ordinary and customary
meaning to a person of ordinary skill in the art (and is not to be
limited to a special or customized meaning), and refers without
limitation to the surface of an electrode where an electrochemical
reaction takes place. In a working electrode, hydrogen peroxide
produced by an enzyme-catalyzed reaction of an analyte being
detected reacts can create a measurable electronic current. For
example, in the detection of glucose, glucose oxidase produces
H.sub.2O.sub.2 peroxide as a byproduct. The H.sub.2O.sub.2 reacts
with the surface of the working electrode to produce two protons
(2H.sup.+), two electrons (2e.sup.-) and one molecule of oxygen
(O.sub.2), which produces the electronic current being detected. In
a counter electrode, a reducible species, for example, O.sub.2 is
reduced at the electrode surface so as to balance the current
generated by the working electrode.
[0123] The term "sensing membrane" as used herein is a broad term,
and is to be given its ordinary and customary meaning to a person
of ordinary skill in the art (and is not to be limited to a special
or customized meaning), and refers without limitation to a
permeable or semi-permeable membrane that can comprise one or more
domains and that is constructed of materials having a thickness of
a few microns or more, and that are permeable to reactants and/or
co-reactants employed in determining the analyte of interest. As an
example, a sensing membrane can comprise an immobilized glucose
oxidase enzyme, which catalyzes an electrochemical reaction with
glucose and oxygen to permit measurement of a concentration of
glucose.
[0124] The term "proximal as used herein is a broad term, and is to
be given its ordinary and customary meaning to a person of ordinary
skill in the art (and is not to be limited to a special or
customized meaning), and refers without limitation to a region near
to a point of reference, such as an origin or a point of
attachment.
[0125] The term "distal" as used herein is a broad term, and is to
be given its ordinary and customary meaning to a person of ordinary
skill in the art (and is not to be limited to a special or
customized meaning), and refers without limitation to a region
spaced relatively far from a point of reference, such as an origin
or a point of attachment.
[0126] The terms "operably connected," "operably linked" and
"operatively coupled" as used herein are broad terms, and are to be
given their ordinary and customary meaning to a person of ordinary
skill in the art (and are not to be limited to a special or
customized meaning), and refer without limitation to one or more
components linked to another component(s) in a manner that
facilitates transmission of signals between the components. For
example, one or more electrodes can be used to detect an analyte in
a sample and convert that information into a signal; the signal can
then be transmitted to an electronic circuit. In this example, the
electrode is "operably linked" to the electronic circuit.
[0127] The term "adhere" and "attach" as used herein are broad
terms, and are to be given their ordinary and customary meaning to
a person of ordinary skill in the art (and are not to be limited to
a special or customized meaning), and refer without limitation to
hold, bind, or stick, for example, by gluing, bonding, grasping,
interpenetrating, or fusing.
[0128] The term "bioactive agent" as used herein is a broad term,
and is to be given its ordinary and customary meaning to a person
of ordinary skill in the art (and is not to be limited to a special
or customized meaning), and refers without limitation to any
substance that has an effect on or elicits a response from living
tissue.
[0129] The term "bioerodible" or "biodegradable" as used herein are
a broad terms, and are to be given their ordinary and customary
meaning to a person of ordinary skill in the art (and are not to be
limited to a special or customized meaning), and refer without
limitation to materials that are enzymatically degraded or
chemically degraded in vivo into simpler components. One example of
a biodegradable material includes a biodegradable polymer that is
broken down into simpler components by the body.
[0130] The terms "small diameter sensor," "small structured
sensor," and "micro-sensor," as used herein are broad terms, and
are to be given their ordinary and customary meaning to a person of
ordinary skill in the art (and are not to be limited to a special
or customized meaning), and refer without limitation to sensing
mechanisms that are less than about 2 mm in at least one dimension,
and more preferably less than about 1 mm in at least one dimension.
In some embodiments, the sensing mechanism (sensor) is less than
about 0.95, 0.9, 0.85, 0.8, 0.75, 0.7, 0.65, 0.6, 0.5, 0.4, 0.3,
0.2, or 0.1 mm. In some embodiments, the sensing mechanism is a
needle-type sensor, wherein the diameter is less than about 1 mm,
see, for example, U.S. Pat. No. 6,613,379 to Ward et al. and
co-pending U.S. patent application Ser. No. 11/077,715, filed on
Mar. 10, 2005 and entitled, "TRANSCUTANEOUS ANALYTE SENSOR," both
of which are incorporated herein by reference in their entirety. In
some alternative embodiments, the sensing mechanism includes
electrodes deposited on a planar substrate, wherein the thickness
of the implantable portion is less than about 1 mm, see, for
example U.S. Pat. No. 6,175,752 to Say et al. and U.S. Pat. No.
5,779,665 to Mastrototaro et al., both of which are incorporated
herein by reference in their entirety.
[0131] The term "electrospinning" as used herein is a broad term,
and is to be given its ordinary and customary meaning to a person
of ordinary skill in the art (and is not to be limited to a special
or customized meaning), and refers without limitation to a process
by which fibers are drawn out from a viscous polymer solution or
melt by applying an electric field to a droplet of the solution
(most often at a metallic needle tip). The electric field draws
this droplet into a structure called a Taylor cone. If the
viscosity and surface tension of the solution are appropriately
tuned, varicose breakup (electrospray) is avoided and a stable jet
is formed. A bending instability results in a whipping process
which stretches and elongates this fiber until it has a diameter of
micrometers (or nanometers).
[0132] The terms "interferants," "interferents" and "interfering
species," as used herein are broad terms, and are to be given their
ordinary and customary meaning to a person of ordinary skill in the
art (and are not to be limited to a special or customized meaning),
and refer without limitation to effects and/or species that
interfere with the measurement of an analyte of interest in a
sensor to produce a signal that does not accurately represent the
analyte measurement. In one example of an electrochemical sensor,
interfering species are compounds with oxidation potentials that
overlap with the oxidation potential of the analyte to be
measured.
[0133] The term "drift," as used herein is a broad term, and is to
be given its ordinary and customary meaning to a person of ordinary
skill in the art (and is not to be limited to a special or
customized meaning), and refers without limitation to a progressive
increase or decrease in signal over time that is unrelated to
changes in host systemic analyte concentrations, such as host
postprandial glucose concentrations, for example. While not wishing
to be bound by theory, it is believed that drift can be the result
of a local decrease in glucose transport to the sensor, due to
cellular invasion, which surrounds the sensor and forms a FBC, for
example. It is also believed that an insufficient amount of
interstitial fluid is surrounding the sensor, which results in
reduced oxygen and/or glucose transport to the sensor, for example.
An increase in local interstitial fluid can slow or reduce drift
and thus improve sensor performance.
[0134] The term "sensing region" as used herein is a broad term,
and is to be given its ordinary and customary meaning to a person
of ordinary skill in the art (and is not to be limited to a special
or customized meaning), and refers without limitation to 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 (optional), and/or a counter electrode (cathode) passing
through and secured within the body forming electrochemically
reactive surfaces on the body and an electronic connective means at
another location on the body, and a multi-domain membrane affixed
to the body and covering the electrochemically reactive
surface.
[0135] The term "domain" as used herein is a broad term, and is to
be given its ordinary and customary meaning to a person of ordinary
skill in the art (and is not to be limited to a special or
customized meaning), and refers without limitation to a region of
the membrane system that can be a layer, a uniform or non-uniform
gradient (for example, an anisotropic region of a membrane), or a
portion of a membrane.
[0136] The term "membrane system," as used herein is a broad term,
and is to be given its ordinary and customary meaning to a person
of ordinary skill in the art (and is not to be limited to a special
or customized meaning), and refers without limitation to 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 is permeable to oxygen and is
optionally permeable to, e.g., glucose or another analyte. In one
example, the membrane system comprises an immobilized glucose
oxidase enzyme, which enables a reaction to occur between glucose
and oxygen whereby a concentration of glucose can be measured.
[0137] The terms "processor module" and "microprocessor," as used
herein are broad terms, and are to be given their ordinary and
customary meaning to a person of ordinary skill in the art (and are
not to be limited to a special or customized meaning), and refer
without limitation to a computer system, state machine, processor,
or the like designed to perform arithmetic or logic operations
using logic circuitry that responds to and processes the basic
instructions that drive a computer.
[0138] The term "STS" or short-term sensor as used herein is a
broad term, and is to be given its ordinary and customary meaning
to a person of ordinary skill in the art (and is not to be limited
to a special or customized meaning), and refers without limitation
to sensors used during a short period of time (e.g., short-term),
such as 1-3 days, 1-7 days, or longer. In some embodiments, the
sensor is used during a short period of time, such as, for 1 day or
less, 2, 3, 4, 5, 6, 7, 8, 9, 10, 11, 12, 13, 24, or 15 days. In
some embodiments, the sensor is used for a short period of time,
such as prior to tissue ingrowth or FBC formation. In some
embodiments, a STS is transcutaneous.
[0139] The term "bulk fluid flow," as used herein is a broad term,
and is to be given its ordinary and customary meaning to a person
of ordinary skill in the art (and is not to be limited to a special
or customized meaning), and refers without limitation to the
movement of fluid(s) within an area or space, or in or out of the
area or space. In one embodiment, the fluid moves in and/or out of
a fluid pocket surrounding the sensor. In another embodiment, the
fluid moves within the fluid pocket. In yet another embodiment, the
fluid moves by convection (e.g., the circulatory motion that occurs
in a fluid at a non-uniform temperature owing to the variation of
its density and the action of gravity).
[0140] The term "fluid influx," as used herein is a broad term, and
is to be given its ordinary and customary meaning to a person of
ordinary skill in the art (and is not to be limited to a special or
customized meaning), and refers without limitation to the movement
of fluid(s) into the locality of an implanted sensor.
[0141] The term "fluid efflux," as used herein is a broad term, and
is to be given its ordinary and customary meaning to a person of
ordinary skill in the art (and is not to be limited to a special or
customized meaning), and refers without limitation to the movement
of fluid(s) out of the locality of an implanted sensor.
[0142] The term "adipose" as used herein is a broad term, and is to
be given its ordinary and customary meaning to a person of ordinary
skill in the art (and is not to be limited to a special or
customized meaning), and refers without limitation to fat under the
skin and surrounding major organs. For example, "adipose tissue" is
fat tissue. In another example, an "adipocyte" is a fat cell.
[0143] The term "edema" as used herein is a broad term, and is to
be given its ordinary and customary meaning to a person of ordinary
skill in the art (and is not to be limited to a special or
customized meaning), and refers without limitation to an abnormal
infiltration and excess accumulation of serous fluid in connective
tissue or in a serous cavity. In one example, edematous fluid is
the fluid an edema.
[0144] The term "comprising" as used herein is a broad term, and is
to be given its ordinary and customary meaning to a person of
ordinary skill in the art (and is not to be limited to a special or
customized meaning), and without limitation to is synonymous with
"including," "containing," or "characterized by," and is inclusive
or open-ended and does not exclude additional, unrecited elements
or method steps.
[0145] The term "shedding layer" as used herein is a broad term,
and is to be given its ordinary and customary meaning to a person
of ordinary skill in the art (and is not to be limited to a special
or customized meaning), and refers without limitation to a layer of
material (e.g., incorporated into a biointerface) that leaches or
releases molecules or components into the surrounding area. One
example of a shedding layer includes, a coating of a biodegradable
material (e.g., polyvinylalcohol or polyethylene oxide) that is
eroded by tissue surrounding the sensor. In another example, the
shedding layer includes a polymer hydrogel that degrades and is
engulfed by circulating macrophages, which can be stimulated to
release inflammatory factors.
[0146] The term "noise," as used herein is a broad term, and is to
be given its ordinary and customary meaning to a person of ordinary
skill in the art (and is not to be limited to a special or
customized meaning), and refers without limitation to a signal
detected by the sensor that is substantially non-analyte related
(e.g., non-glucose related) and can result in less accurate sensor
performance. One type of noise has been observed during the few
hours (e.g., about 2 to about 36 hours) after sensor insertion.
After the first 24-36 hours, the noise often disappears, but in
some hosts, the noise can last for about three to four days.
[0147] The term "nanoporous," as used herein is a broad term, and
is to be given its ordinary and customary meaning to a person of
ordinary skill in the art (and is not to be limited to a special or
customized meaning), and refers without limitation to materials
consist of a regular organic or inorganic framework supporting a
regular, porous structure having pores roughly in the nanometer
range (e.g., between 1.times.10.sup.-7 and 0.2.times.10.sup.-9
m).
[0148] All references cited herein, including but not limited to
published and unpublished applications, patents, and literature
references, are incorporated herein by reference in their entirety
and are hereby made a part of this specification. To the extent
publications and patents or patent applications incorporated by
reference contradict the disclosure contained in the specification,
the specification is intended to supersede and/or take precedence
over any such contradictory material.
[0149] All numbers expressing quantities of ingredients, reaction
conditions, and so forth used in the specification are to be
understood as being modified in all instances by the term "about."
Accordingly, unless indicated to the contrary, the numerical
parameters set forth herein are approximations that can vary
depending upon the desired properties sought to be obtained. At the
very least, and not as an attempt to limit the application of the
doctrine of equivalents to the scope of any claims in any
application claiming priority to the present application, each
numerical parameter should be construed in light of the number of
significant digits and ordinary rounding approaches.
Overview
Noise
[0150] Generally, implantable sensors measure a signal (e.g.,
counts) related to an analyte of interest in a host. For example,
an electrochemical sensor can measure glucose, creatinine, or urea
in a host, such as an animal, especially a human. Generally, the
signal is converted mathematically to a numeric value indicative of
analyte status, such as analyte concentration. It is not unusual
for a sensor to experience a certain level of noise. "Noise," as
used herein, is a broad term and is used in its ordinary sense,
including, without limitation, a signal detected by the sensor that
is substantially non-analyte related (e.g., non-glucose related)
and can result in reduced sensor performance. Noise can be caused
by a variety of factors, such as interfering species, macro- or
micro-motion, ischemia, pH changes, temperature changes, pressure,
stress, or even unknown sources of mechanical, electrical and/or
biochemical noise for example. Since noise can obscure analyte
data, reduction of noise is desirable.
[0151] There are a variety of ways noise can be recognized and/or
analyzed. In preferred embodiments, the sensor data stream is
monitored, signal artifacts are detected and data processing is
based at least in part on whether or not a signal artifact has been
detected, such as described in U.S. Publication No.
US-2005-0043598-A1.
[0152] It was observed that some inserted sensors functioned more
poorly during the first few hours or days after insertion than they
did later. This was exemplified by noise and/or a suppression of
the signal during the first about 2-36 hours or more after
insertion. These anomalies often resolved spontaneously, after
which the sensors became less noisy, had improved sensitivity, and
were more accurate than during the early period. Moreover, the
noise predominated when hosts were sleeping or sedentary for a
period of time.
[0153] FIG. 1A illustrates this phenomenon of noise associated with
the above-described intermittent sedentary activity during the
first few days of insertion of a STS glucose sensor containing
active enzyme (in a non-diabetic host). The X-axis represents time;
the left Y-axis represents sensor signal in counts (e.g., signal to
be converted into glucose level in mg/dL) and the right Y-axis
represents noise within the sensor signal in counts (determined
algorithmically according to copending U.S. patent application Ser.
No. 11/498,410, filed Aug. 2, 2006 and entitled "SYSTEMS AND
METHODS FOR REPLACING SIGNAL ARTIFACTS IN A GLUCOSE SENSOR DATA
STREAM" herein incorporated by reference in its entirety). An
enzymatic glucose sensor was built, including enzyme, as described
in U.S. Publication No. US-2005-0020187-A1. During the day, the
sensor signal (upper line) varied and substantially correlated with
glucose concentration. But, when the host went to sleep at about
midnight, noise (lower line) began to occur. Between midnight and 6
AM, when the host was asleep, there was a lot of noise, as
evidenced by the large number of high peaks in the noise plot
(lower line). When the host awoke and began moving around, at about
6 AM, the noise dissipated and signal substantially represented
glucose concentration again.
[0154] Studies using enzymatic-type glucose sensors built without
enzyme were tested in non-diabetic individuals. These sensors
(without enzyme) do not react with or measure glucose and therefore
provide a signal due to non-glucose effects (e.g., baseline,
interferants, and noise). These studies demonstrated that the noise
observed during sedentary periods was caused by something other
than glucose concentration. FIG. 1B shows one example of the
experimental results, in a non-diabetic host wearing a STS glucose
sensor built without enzyme. When the host was asleep, the
no-enzyme sensor showed large, sustained positive signals that
resembled glucose peaks, but could not represent actual glucose
concentration because the sensor lacked enzyme. In the morning,
when the host awoke and moved around, the no-enzyme signal rapidly
corrected, becoming measurably reduced and smoother. From these
results, the inventors believe that a reactant was diffusing to the
electrodes and producing the unexpected positive signal.
[0155] Additional, in vitro experiments were conducted to determine
if a sensor (e.g., electrode) component might have leached into the
area surrounding the sensor. These in vitro experiments provided
evidence that the non-glucose signals (observed during host
sedentary periods) were not produced by contaminants of the sensor
itself, or products of the chemical reaction at the electrodes,
because the noise and non-glucose peaks did not occur in vitro.
[0156] While not wishing to be bound by theory, it is believed that
intermittent, sedentary noise is caused by an interferant that is
likely produced by local cellular activity (e.g., associated with
wound healing) at the site of sensor insertion. Physiologic
activity at a wound site is complex and involves the interaction of
a variety of body processes. In order to fully understand the cause
of intermittent, sedentary noise (as well as solutions), we must
understand wound healing, fluid transport within the body (e.g.,
lymph transport) and tissue response to implanted materials (e.g.,
foreign body response). Each of these processes is discussed in
greater detail below.
Wound Healing
[0157] When a foreign body is inserted into a host, it creates a
wound, by breaking the skin and some of the underlying tissue,
thereby initiating the wound-healing cascade of events. A wound is
also produced, when a sensor, such as an implantable glucose
sensor, is implanted into the subcutaneous tissue. For short-term
use sensors, as described elsewhere herein, wounding occurs at
least from the penetration of the sharp needle or device, which can
be used to deliver the sensor. The wound can be relatively
extensive, including bruising and/or bleeding, or it can be
relatively benign, with little tissue damage and little or
virtually no bleeding. Wound healing is initiated immediately upon
wounding and is directed by a series of signaling cascades. Wound
healing has four main phases: 1) hemostasis, 2) inflammation, 3)
granulation, and 4) remodeling, which are discussed in more detail
below.
[0158] The "hemostasis" phase begins during the first few seconds
and minutes after wounding and entails a cascade of molecular
events that lead to cessation of bleeding, and the formation of a
fibrin scaffold that will be used as a support for cellular
responses that follow. During hemostasis, blood platelets are
activated by exposure to extravascular collagen and release soluble
mediators (growth factors and cAMP) and adhesive glycoproteins that
cause the platelets to aggregate and form a fibrin clot.
Neutrophils and monocytes are attracted to the wound by
platelet-derived growth factor (PDGR) and transforming growth
factor beta (TGF-.beta.), to clean the wound of infectious
material, foreign matter and devitalized tissue. Vascular
endothelial growth factor (VEGF or VPF), transforming growth factor
alpha (TGF-.alpha.) and basic fibroblast growth factor (bFGF),
which are also secreted by activated platelets, activate
endothelial cells that begin angiogenesis. "Angiogenesis" is a
physiological process involving the growth of new blood vessels
from pre-existing vessels. Platelet secreted PDGF also activates
and recruits fibroblasts to produce extracellular matrix
components.
[0159] The "inflammation" stage begins within the first 24 hours
after injury and can last for several weeks in normal wounds and
significantly longer in chronic nonhealing wounds. This occurs
within several hours after implantation, and is the stage that most
closely correlates with the anomalous behavior of the short-term
sensor (STS) Inflammation involves the influx of polymorphonuclear
cells and the formation of an edematous fluid pocket surrounding
the implant. The vascular epithelium becomes highly permeable to
cells and fluid so that invading cells (neutrophils, monocytes, and
macrophages) can get to the wound site. Mast cells in the wound
site release enzymes, histamine, and active amines can cause
swelling, redness, heat, and pain depending on the severity of the
wound. In most needle track wounds, the extent of the reaction is
not sufficient to cause noticeable welling, redness, heat, or pain.
Neutrophils, monocytes and macrophages release proinflammatory
cytokines (IL-1, IL-6, IL-8 and TNF-.alpha.) and cleanse the wound
by engulfing bacteria, debris and devitalized tissue. These cells
are highly active phagocytic cells with high metabolic
requirements, and in an early wound they are proliferating
exponentially, creating a need for oxygen, glucose and other
molecules. Fibroblasts and epithelial cells are recruited and
activated by PDGF, TGF-.beta., TGF-.alpha., insulin-like growth
factor 1 (IGF-1) and FGF, in preparation for the next phase of
wound healing.
[0160] The "granulation" phase occurs after several days, involving
the full participation of a large number of macrophages, and the
initiation of fibrosis and vascularization. During the
proliferative phase of wound healing, fibroblasts proliferate and
deposit granulation tissue components (various types of collagen,
elastin, and proteoglycans). Angiogenesis also takes place at this
time. Angiogenesis is stimulated by local low oxygen tension.
Oxygen promotes angiogenesis by binding hypoxia-inducible factor
(HIF) within capillary endothelial cells. When oxygen is low around
capillary endothelial cells, HIF levels inside the cells increase
and stimulate the production of VEGF, which stimulates
angiogenesis. Low pH, high lactate levels, bFGF, and TGF-.beta.
also stimulate angiogenesis. Epithelial cells also proliferate and
form a new epidermis over the wound.
[0161] The "remodeling" phase occurs after several weeks and is not
relevant to sensors used for short periods of time, such as about 1
to 3 days, or up to about 7 days or more, or up to about 2 weeks.
In the case of long-term wholly implantable sensors, this process
is involved in remodeling tissue around the wholly implantable
sensor.
[0162] The rate of these responses can vary dramatically in a host
population, especially among diabetics, who are known to suffer
from vascular and wound-healing disorders. Moreover, there is wide
variability in the amount, texture, morphology, color, and
vascularity of subcutaneous tissue. Therefore it is to be expected
that the rate of progress of the wound-healing response, and the
quality of the response can vary dramatically among hosts.
[0163] Dramatic differences in wounding and noise exist among
individuals. Some people wound easily (e.g., bruise more easily or
have more bleeding) while others do not. Some people exhibit more
noise (e.g., are noisier) in their sensor signal than others. In
one example, a glucose tracking study was performed with two
non-diabetic volunteer hosts. Samples were collected from the
fingertip and the lower abdomen, (e.g., where some short-term
sensors are usually implanted). Concurrent blood samples were
collected from both the fingertip and abdomen, using a lancet
device. The collected blood samples were measured with a hand-held
glucose meter.
[0164] FIG. 1C illustrates the difference in responses of finger
and abdominal tissue to oral sugar consumption, in a first
non-diabetic volunteer host (Host 1). The solid line (with squares)
shows glucose concentration at the fingertip. The dashed line (with
diamonds) shows glucose concentration at the lower abdomen. When
Host 1 ingested about 100 gm of oral sucrose, there was a dramatic
and rapid increase in glucose signal from the fingertip samples.
Host 1's abdominal signal exhibited a slower and reduced rise, when
compared with the fingertip samples.
[0165] FIG. 1D illustrates the difference in responses of finger
and abdominal tissue to oral sugar consumption, in a second
volunteer non-diabetic host (Host 2). The solid line (with squares)
shows glucose concentration at the fingertip. The dashed line (with
diamonds) shows glucose concentration at the lower abdomen. When
Host 2 was challenged with sucrose consumption, he exhibited little
difference between his fingertip and abdominal samples. These data
suggest that sensors implanted in different individuals can behave
differently.
[0166] Different individuals experience relatively different
amounts of intermittent, sedentary noise. For example, Host 1, when
wearing a short-term sensor, typically was known to experience high
levels of nighttime noise, whereas Host 2 experienced very little
noise at any time while wearing an exemplary STS.
[0167] In addition, the amount of wounding varies between
individuals as well as between body sites of a single individual.
For example, the next day, Host 1's lower abdomen exhibited
extensive bruising (e.g., approximately 20 hours after completing
the study). Note the many bruises 250, 252 in FIG. 1E. However,
Host 1's fingertips had very little observable wounding the next
day (FIG. 1F). In contrast, Host 2 (not shown) sustained little
visible wounding the next day (from the lancet), at either the
lower abdomen or fingertips.
[0168] When a sensor is first inserted into the subcutaneous
tissue, it comes into contact with a wide variety of possible
tissue conformations. Subcutaneous tissue in different hosts can be
relatively fat free in cases of very athletic people, or can be
mostly composed of fat as in the majority of people. The fat comes
in a wide array of textures from very white, puffy fat to very
dense, fibrous fat. Some fat is very yellow and dense in
appearance; some is very clear, puffy, and white in appearance,
while in other cases it is more red or brown in appearance. The fat
can be several inches thick or only 1 cm thick. It can be very
vascular or relatively nonvascular. Many diabetes hosts have some
subcutaneous scar tissue due to years of insulin pump use or
insulin injection. At times, sensors can come to rest in such a
scarred area. The subcutaneous tissue can even vary greatly from
one location to another in the abdomen of a given host. Moreover,
by chance, the sensor can come to rest near a more densely
vascularized area of a given host or in a less vascularized
area.
[0169] FIG. 2B is a side schematic view of adipose cell contact
with an inserted transcutaneous sensor or an implanted sensor 34.
In this case, the sensor is firmly inserted into a small space with
adipose cells pressing up against the surface. Close association of
the adipose cells with the sensor can also occur, for example
wherein the surface of the sensor is hydrophobic. For example, the
adipose cells 200 can physically block the surface of the
sensor.
[0170] Typically adipose cells are about 120 microns in diameter
and are typically fed by tiny capillaries 205. When the sensor is
pressed against the fat tissue, as shown in FIG. 2B, very few
capillaries can actually come near the surface of the sensor. This
can be analogous to covering the surface of the sensor with an
impermeable material such as plastic wrap, for example. Even if
there were a few small holes in the plastic wrap, the sensor's
function would likely be compromised. Additionally, the surrounding
tissue has a low metabolic rate and therefore does not require high
amounts of glucose and oxygen. While not wishing to be bound by
theory, it is believed that, during this early period, the sensor's
signal can be noisy and the signal can be suppressed due to close
association of the sensor surface with the adipose cells and
decreased availability of oxygen and glucose both for
physical-mechanical reasons and physiological reasons.
[0171] Because of the host-to-host variability, the location
variability in a given host, and the random possibility of hitting
a favorable or unfavorable spot in a host, every time an
implantable device (e.g., a sensor) is inserted into a host it has
the chance of responding differently than it did in another host or
at another time or place in the same host. For example, another
host can insert a needle or device on day 1 and have no bleeding or
bruising, but when she inserts another needle or device on day 3
she can have bleeding with an associated bruise. The wound healing
response in a bloody wound will be expected to be considerably
different than in a less traumatized wound. As another example,
another host can have produced considerable trauma on insertion of
a needle/device, without visible bleeding or bruising.
[0172] In the case of a less traumatic wound, we believe the
inflammatory phase of the wound response would be delayed for some
length of time. In the case of a more traumatized wound, we believe
it would be accelerated. For example, a fluid pocket can take hours
to form in the less traumatic wound whereas it could take much less
time in the case of the more traumatic wound.
[0173] In the case of a less traumatic wound, when an implantable
device, such as a glucose sensor, is initially inserted, relatively
little tissue damage occurs. The device finds itself firmly
inserted into a small space with adipose tissue pressing up against
the surface. Because the surface of the sensor (e.g., a STS sensor
as described herein) is mainly very hydrophobic, it can associate
very closely with the adipose tissue. Because no edema (e.g., wound
fluid) is forming or is forming slowly, there will be very little
fluid around the sensor for glucose transport. Accordingly, adipose
cells can physically block the surface of the sensor. When the
sensor is pressed against the adipose tissue, it is believed that
that very few capillaries come near the surface of the sensor.
Additionally, the surrounding tissue has a low metabolic rate and
therefore does not require high amounts of glucose and oxygen.
While not wishing to be bound by hypothesis, it is believed that
during this period (prior to the formation of an edematous pocket
and the influx of cells and glucose) the sensor signal can be noisy
and suppressed due to close association of the sensor surface with
the adipose cells and lack of availability of oxygen and glucose
both for physical-mechanical reasons and physiological reasons.
While not wishing to be bound by theory, it is believed that the
short-term sensor measures wound fluid surrounding the sensor.
Thus, if the rate of edema collection (e.g., collection of wound
fluid into a fluid pocket) can be increased then early noise can be
alleviated or reduced.
Lymph System and Fluid Transport
[0174] The circulatory and lymph systems are the body's means of
moving fluids, cells, protein, lipids, and the like throughout the
body in an organized fashion. The two systems parallel each other,
throughout the body. The circulatory system is a closed system that
relies on a pump (the heart) for control of bulk flow. In contrast,
the lymph system is an open system with no central pump. The lymph
system relies upon pressure differentials, local muscle
contraction, among other things, for fluid movement. Gravity and
inactivity can have dramatic effects on lymph movement throughout
the body, and consequently on noise and sensor function.
[0175] Lymph forms when dissolved proteins and solutes filter out
of the circulatory system into the surrounding tissues, because of
local differences in luminal hydrostatic and osmotic pressure. The
fluid within the extracellular spaces is called interstitial fluid.
A portion of the interstitial fluid flows back into the circulatory
system, while the remaining fluid is collected into the lymph
capillaries through valve-like openings between the endothelial
cells of the lymph capillaries.
[0176] Lymph is generally a clear and transparent semifluid medium.
It is known in the art that normal cellular metabolism produces
waste species that are removed from the local environment by the
lymphatics. Lymph contains a "lymphatic load" of protein, water,
lymphocytes, cellular components, metabolic waste and particles,
and fat. The lymphatics return the lymph to the circulatory system
at the thoracic duct. It is known that lymph has almost the same
composition as the original interstitial fluid.
[0177] In contrast to the circulatory system, the lymph system is
an open system with no central pump. Lymph capillaries take in
fluid through "open junctions," until they are filled to capacity.
When the pressure inside the capillary is greater than that of the
surrounding interstitial tissue, the open junctions close. The
lymph moves freely toward larger, downstream portions of the lymph
system, where pressure is lower. As the lymph moves forward, it is
picked up by "lymph collectors," which have valves that prevent
fluid back-flow. Larger portions of the lymph system segmentally
contract, to push the lymph forward, from one segment to the next.
Breathing movements and skeletal muscle contractions also push the
lymph forward. Eventually, the lymph is returned to the circulatory
system via the thoracic duct.
[0178] Lymph capillaries are delicate and easily flattened. When
lymph capillaries are flattened, fluid cannot enter them.
Consequently, lymph flow is impeded by a local collapse of the
lymph capillaries. Gravity and local pinching of lymph capillaries
affect the movement of lymph. For example, it is well known in the
medical community that a tourniquet placed on the upper arm can
impede lymph flow out of the arm. It is also known that during
sleep lymph pools on the side of the body on which a person is
lying. In another example, sitting can pinch some of the lower
lymphatics, causing lymph to pool in the legs over an extended
period of time.
[0179] As discussed with reference to FIG. 1B, above, the inventors
have found that, soon after insertion of a sensor, noise (e.g.,
signal) not associated with glucose concentration can occur
intermittently during sedentary activities, such as sleeping,
watching television or reading a book. The inventors have
demonstrated experimentally that early intermittent, sedentary
noise is, at least in part, the result of unknown interferants that
affect the sensor during periods of sustained inactivity.
[0180] While not wishing to be bound by theory, it is believed that
a local build up of electroactive interferants, such as
electroactive metabolites from cellular metabolism and wound
healing, interfere with sensor function and cause early
intermittent, sedentary noise. Local lymph pooling, when parts of
the body are compressed or when the body is inactive can cause, in
part, this local build up of interferants (e.g., electroactive
metabolites). Interferants can include but are not limited to
compounds with electroactive acidic, amine or sulfhydryl groups,
urea, lactic acid, phosphates, citrates, peroxides, amino acids
(e.g., L-arginine), amino acid precursors or break-down products,
nitric oxide (NO), NO-donors, NO-precursors or other electroactive
species or metabolites produced during cell metabolism and/or wound
healing, for example.
Foreign Body Response
[0181] Devices and probes that are transcutaneously inserted or
implanted into subcutaneous tissue conventionally elicit a foreign
body response (FBR), which includes invasion of inflammatory cells
that ultimately forms a foreign body capsule (FBC), as part of the
body's response to the introduction of a foreign material.
Specifically, insertion or implantation of a device, for example, a
glucose sensing device, can result in an acute inflammatory
reaction resolving to chronic inflammation with concurrent building
of fibrotic tissue, such as is described in detail above.
Eventually, over a period of two to three weeks, a mature FBC,
including primarily contractile fibrous tissue forms around the
device. See Shanker and Greisler, Inflammation and Biomaterials in
Greco RS, ed., "Implantation Biology: The Host Response and
Biomedical Devices" pp 68-80, CRC Press (1994). The FBC surrounding
conventional implanted devices has been shown to hinder or block
the transport of analytes across the device-tissue interface. Thus,
continuous extended life analyte transport (e.g., beyond the first
few days) in vivo has been conventionally believed to be unreliable
or impossible.
[0182] FIG. 2A is a schematic drawing that illustrates a classical
FBR to a conventional cell-impermeable synthetic membrane 10
implanted under the skin. There are three main layers of a FBR. The
innermost FBR layer 12, adjacent to the device, is composed
generally of macrophages and foreign body giant cells 14 (herein
referred to as the "barrier cell layer"). These cells form a
monolayer of closely opposed cells over the entire surface of a
microscopically smooth membrane, a macroscopically smooth (but
microscopically rough) membrane, or a microporous (i.e., average
pore size of less than about 1 .mu.m) membrane. A membrane can be
adhesive or non-adhesive to cells; however, its relatively smooth
surface causes the downward tissue contracture 21 (discussed below)
to translate directly to the cells at the device-tissue interface
26. The intermediate FBR layer 16 (herein referred to as the
"fibrous zone"), lying distal to the first layer with respect to
the device, is a wide zone (about 30 to 100 .mu.m) composed
primarily of fibroblasts 18, fibrous matrixes, and contractile
fibrous tissue 20. The organization of the fibrous zone, and
particularly the contractile fibrous tissue 20, contributes to the
formation of the monolayer of closely opposed cells due to the
contractile forces 21 around the surface of the foreign body (for
example, membrane 10). The outermost FBR layer 22 is loose
connective granular tissue containing new blood vessels 24 (herein
referred to as the "vascular zone"). Over time, this FBR tissue
becomes muscular in nature and contracts around the foreign body so
that the foreign body remains tightly encapsulated. Accordingly,
the downward forces 21 press against the tissue-device interface
26, and without any counteracting forces, aid in the formation of a
barrier cell layer 14 that blocks and/or refracts the transport of
analytes 23 (for example, glucose) across the tissue-device
interface 26.
[0183] A consistent feature, of the innermost layers 12, 16, is
that they are devoid of blood vessels. This has led to widely
supported speculation that poor transport of molecules across the
device-tissue interface 26 is due to a lack of vascularization near
the interface. See Scharp et al., World J. Surg., 8:221-229 (1984);
and Colton et al., J. Biomech. Eng., 113:152-170 (1991). Previous
efforts to overcome this problem have been aimed at increasing
local vascularization at the device-tissue interface, but have
achieved only limited success.
[0184] Although local vascularization can aid in sustenance of
local tissue over time, the presence of a barrier cell layer 14
prevents the passage of molecules that cannot diffuse through the
layer. For example, when applied to an implantable
glucose-measuring device, it is unlikely that glucose would enter
the cell via glucose transporters on one side of the cell and exit
on the other side. Instead, it is likely that any glucose that
enters the cell is phosphorylated and remains within the cell. The
only cells known to facilitate transport of glucose from one side
of the cell to another are endothelial cells. Consequently, little
glucose reaches the implant's membrane through the barrier cell
layer. The known art purports to increase the local vascularization
in order to increase solute availability. See Brauker et al., U.S.
Pat. No. 5,741,330. However, it has been observed by the inventors
that once the monolayer of cells (barrier cell layer) is
established adjacent to a membrane, increasing angiogenesis is not
sufficient to increase transport of molecules such as glucose and
oxygen across the device-tissue interface 26. In fact, the barrier
cell layer blocks and/or reflects the analytes 23 from transport
across the device-tissue interface 26.
[0185] Referring now to short-term sensors, or the short-term
function of long-term sensors, it is believed that certain aspects
of the FBR in the first few days can play a role in noise. It has
been observed that some sensors function more poorly during the
first few hours after insertion than they do later. This is
exemplified by noise and/or a suppression of the signal during the
first few hours (e.g., about 2 to about 36 hours) after insertion.
These anomalies often resolve spontaneously after which the sensors
become less noisy, have improved sensitivity, and are more accurate
than during the early period. It has been observed that some
transcutaneous sensors and wholly implantable sensors are subject
to noise for a period of time after application to the host (i.e.,
inserted transcutaneously or wholly implanted below the skin).
"Noise," as used herein, is a broad term and is used in its
ordinary sense, including, without limitation, a signal detected by
the sensor that is unrelated to analyte concentration and can
result in less accurate sensor performance. One type of noise has
been observed during the few hours (e.g., about 2 to about 36
hours) after sensor insertion. After the first few hours to 36
hours, the noise often disappears, but in some hosts, the noise can
last longer.
[0186] Referring now to long-term function of a sensor, after a few
days to two or more weeks of implantation, many prior art devices
typically lose their function. In some applications, cellular
attack or migration of cells to the sensor can cause reduced
sensitivity and/or function of the device, particularly after the
first day of implantation. See also, for example, U.S. Pat. No.
5,791,344 and Gross et al. and "Performance Evaluation of the
MiniMed Continuous Monitoring System During Host home Use,"
Diabetes Technology and Therapeutics, (2000) 2(1):49-56, which have
reported a glucose oxidase-based device, approved for use in humans
by the Food and Drug Administration, that functions well for
several days following implantation but loses function quickly
after the several days (e.g., a few days up to about 14 days).
[0187] It is believed that this lack of device function is most
likely due to cells, such as polymorphonuclear cells and monocytes,
which migrate to the sensor site during the first few days after
implantation. These cells consume local glucose and oxygen. If
there is an overabundance of such cells, they can deplete glucose
and/or oxygen before it is able to reach the device enzyme layer,
thereby reducing the sensitivity of the device or rendering it
non-functional. Further inhibition of device function can be due to
inflammatory cells, for example, macrophages, that associate with
the implantable device (for example, align at an interface) and
physically block the transport of glucose into the device (for
example, by formation of a barrier cell layer). Additionally, these
inflammatory cells can biodegrade many artificial biomaterials
(some of which were, until recently, considered non-biodegradable).
When activated by a foreign body, tissue macrophages degranulate,
releasing hypochlorite (bleach) and other oxidative species.
Hypochlorite and other oxidative species are known to break down a
variety of polymers.
[0188] In some circumstances, for example in long-term sensors, it
is believed that the foreign body response is the dominant event
surrounding extended implantation of an implanted device, and can
be managed or manipulated to support rather than hinder or block
analyte transport. In another aspect, in order to extend the
lifetime of the sensor, preferred embodiments employ materials that
promote vascularized tissue ingrowth, for example within a porous
biointerface membrane. For example tissue in-growth into a porous
biointerface material surrounding a long-term sensor can promote
sensor function over extended periods of time (e.g., weeks, months,
or years). It has been observed that in-growth and formation of a
tissue bed can take up to about 3 weeks or more. Tissue ingrowth
and tissue bed formation is believed to be part of the foreign body
response. As will be discussed herein, the foreign body response
can be manipulated by the use of porous biointerface materials that
surround the sensor and promote ingrowth of tissue and
microvasculature over time. Long-term use sensors (LTS), for use
over a period of weeks, months or even years, have also been
produced. LTS can be wholly implantable, and placed within the
host's soft tissue below the skin, for example.
[0189] Accordingly, a long-term sensor including a biointerface,
including but not limited to, for example, porous biointerface
materials including a solid portion and interconnected cavities,
all of which are described in more detail elsewhere herein, can be
employed to improve sensor function in the long-term (e.g., after
tissue ingrowth).
Reduction of Intermittent, Sedentary Noise
[0190] As discussed above, noise can occur during the first few
hours or days after sensor implantation, during periods of
inactivity. While not wishing to be bound by theory, the inventors
believe noise that occurs during these early intermittent sedentary
time periods can be caused by a local increase in interferants
(e.g., electroactive metabolites) that disrupt sensor function,
resulting in apparent glucose signals that are generally unrelated
to the host's glucose concentration. Accordingly, the noise
intensity and/or number of intermittent, sedentary noise
occurrences can be reduced or eliminated by reducing the local
concentration of interferants produced during normal cellular
metabolism and/or wound healing.
[0191] In some circumstances, the inventors believe that
intermittent, sedentary noise can be addressed either by affecting
wounding and/or the wound healing process. For example, in some
circumstances a wounding response initiated when the sensor (e.g.,
a glucose sensor) is implanted can lead to in insubstantial
transport of interferents away from the sensor during sedentary
periods, which can result in increased intermittent, sedentary
noise. Thus, it interferent concentration is reduced, such as by
increasing fluid bulk, bulk fluid flow, or diffusion rates (e.g.,
with vasodilation agents or inflammatory agents), prolonging
wounding (e.g., with irritating structures or agents) or promoting
wound healing's inflammation stage, then noise can be reduced.
[0192] The present invention provides, among other things, devices,
and methods for reducing or eliminating noise caused by
intermittent interferant build-up in the area surrounding an
inserted sensor during the first few hours or days
post-implantation. As will be discussed in greater detail below,
these devices and methods contemplate, among other things,
increasing bulk fluid flow in and/or out of the sensor locality,
increased fluid bulk, production of increased or continued wounding
of the insertion site, suppression, and/or prevention of wounding
during and after sensor insertion, and combinations thereof. Those
knowledgeable in the art will recognize that the various structures
and bioactive agents disclosed herein can be employed in a
plurality of combinations, depending upon the desired effect and
the noise reduction strategy selected.
Increasing Fluid Bulk or Bulk Fluid Flow
[0193] Analyte sensors for in vivo use over various lengths of time
have been developed. For example, sensors to be used for a short
period of time, such as about 1 to about 14 days, have been
produced. Herein, this sensor will be referred to as a short-term
sensor (STS). A STS can be a transcutaneous device, in that a
portion of the device can be inserted through the host's skin and
into the underlying soft tissue while a portion of the device
remains on the surface of the host's skin. In one aspect, in order
to overcome the problems associated with noise, such as
intermittent, sedentary noise, or other sensor function in the
short-term (e.g., short-term sensors or short-term function of
long-term sensors), preferred embodiments employ materials that
promote formation of a fluid pocket around the sensor, for example
architectures such as porous biointerface membrane, matrices or
other membrane/mechanical structures that create a space between
the sensor and the surrounding tissue.
[0194] The concentration of interferants (e.g., electroactive
metabolites) surrounding the sensor can be reduced by, among other
things, increasing fluid bulk (e.g., a fluid pocket), an increased
bulk fluid flow and/or an increased diffusion rate around at least
a portion of the sensor, such as the sensing portion of the sensor.
One embodiment of the present invention provides a device with
reduced intermittent sedentary noise having an architecture that
allows and/or promotes increased fluid bulk and/or increased bulk
fluid flow in the area surrounding at least a portion of an
implanted sensor in vivo.
[0195] A variety of structures can be incorporated into the sensor
to allow and/or promote increased (e.g., to stimulate or to
promote) fluid bulk, bulk fluid flow, and/or diffusion rate. These
structures can include but are not limited to spacers, meshes,
shedding layers, roughened surfaces, machineable materials,
nanoporous materials, shape-memory materials, porous memory
materials, self-assembly materials, collapsible materials,
biodegradable materials, combinations thereof, and the like.
Structures that promote increased fluid bulk and/or increased bulk
fluid flow can also include but are not limited to structures that
promote fluid influx or efflux (e.g., fluid influx-promoting
architecture, fluid efflux-promoting architecture), that promote
vasodilation (e.g., vasodilating architecture), that promote
inflammation (e.g., inflammatory architecture), that promote wound
healing or perpetuate wounding (e.g., wound-healing architecture
and wounding architecture, respectively), that promote angiogenesis
(e.g., angiogenic architecture), that suppress inflammation (e.g.,
an anti-inflammatory architecture) or combinations thereof.
[0196] In one embodiment, a porous material that results in
increased fluid bulk, bulk fluid flow and/or diffusion rate, as
well as formation of close vascular structures, is a porous polymer
membrane, such as but not limited to polytetrafluoroethylene
(PTFE), polysulfone, polyvinylidene difluoride, polyacrylonitrile,
silicone, polytetrafluoroethylene, expanded
polytetrafluoroethylene, polyethylene-co-tetrafluoroethylene,
polyolefin, polyester, polycarbonate, biostable
polytetrafluoroethylene, homopolymers, copolymers, terpolymers of
polyurethanes, polypropylene (PP), polyvinylchloride (PVC),
polyvinylidene fluoride (PVDF), polyvinyl alcohol (PVA),
polybutylene terephthalate (PBT), polymethylmethacrylate (PMMA),
polyether ether ketone (PEEK), polyamides, polyurethanes,
cellulosic polymers, poly(ethylene oxide), poly(propylene oxide)
and copolymers and blends thereof, polysulfones and block
copolymers thereof including, for example, di-block, tri-block,
alternating, random and graft copolymers, as well as metals,
ceramics, cellulose, hydrogel polymers, poly (2-hydroxyethyl
methacrylate, pHEMA), hydroxyethyl methacrylate, (HEMA),
polyacrylonitrile-polyvinyl chloride (PAN-PVC), high density
polyethylene, acrylic copolymers, nylon, polyvinyl difluoride,
polyanhydrides, poly(l-lysine), poly (L-lactic acid), and
hydroxyethylmethacrylate, having an average nominal pore size of at
least about 0.6 to 20 .mu.m, using conventional methods for
determination of pore size in the trade. In one embodiment, at
least approximately 50% of the pores of the membrane have an
average size of approximately 0.6 to about 20 .mu.m, such as
described in U.S. Pat. No. 5,882,354. In this exemplary embodiment,
the structural elements, which provide the three-dimensional
conformation, can include fibers, strands, globules, cones or rods
of amorphous or uniform geometry that is smooth or rough. These
elements, hereafter referred to as "strands," have in general one
dimension larger than the other two and the smaller dimensions do
not exceed five microns.
[0197] In another further embodiment, the porous polymer membrane
material, as described above, consists of strands that define
"apertures" formed by a frame of the interconnected strands. The
apertures have an average size of no more than about 20 .mu.m in
any but the longest dimension. The apertures of the material form a
framework of interconnected apertures, defining "cavities" that are
no greater than an average of about 20 .mu.m in any but the longest
dimension. In another embodiment the porous polymer membrane
material has at least some apertures having a sufficient size to
allow at least some vascular structures to be created within the
cavities. At least some of these apertures, while allowing vascular
structures to form within the cavities, prevent connective tissue
from forming therein because of size restrictions.
[0198] In a further embodiment, the porous membrane has frames of
elongated strands of material that are less than 5 microns in all
but the longest dimension and the frames define apertures which
interconnect to form three-dimensional cavities which permit
substantially all inflammatory cells migrating into the cavities to
maintain a rounded morphology. Additionally, the porous material
promotes vascularization adjacent but not substantially into the
porous material upon implantation into a host. Exemplary materials
include but are not limited to polyethylene, polypropylene,
polytetrafluoroethylene (PTFE), cellulose acetate, cellulose
nitrate, polycarbonate, polyester, nylon, polysulfone, mixed esters
of cellulose, polyvinylidene difluoride, silicone,
polyacrylonitrile, and the like.
[0199] In some embodiments, a short-term sensor is provided with a
spacer adapted to provide a fluid pocket between the sensor and the
host's tissue. It is believed that this spacer, for example a
biointerface material, matrix, mesh, hydrogel and like structures
and the resultant fluid pocket provide for oxygen and/or glucose
transport to the sensor.
[0200] FIG. 2C is a side schematic view of a biointerface membrane
as the spacer preventing adipose cell contact with an inserted
transcutaneous sensor or an implanted sensor in one exemplary
embodiment. In this illustration, a porous biointerface membrane 68
surrounds the sensor 34, covering the sensing mechanism (e.g., at
least a working electrode 38) and is configured to fill with fluid
in vivo, thereby creating a fluid pocket surrounding the sensor.
Accordingly, the adipose cells surrounding the sensor are held a
distance away (such as the thickness of the porous biointerface
membrane, for example) from the sensor surface. Accordingly, as the
porous biointerface membrane fills with fluid (e.g., creates a
fluid pocket), oxygen and glucose are transported to the sensing
mechanism in quantities sufficient to maintain accurate sensor
function. Additionally, as discussed elsewhere herein, interferants
are diluted, suppressing or reducing interference with sensor
function.
[0201] Accordingly, a short-term sensor (or short-term function of
a long-term sensor) including a biointerface, including but not
limited to, for example, porous biointerface materials, mesh cages,
and the like, all of which are described in more detail elsewhere
herein, can be employed to improve sensor function in the
short-term (e.g., first few hours to days). Porous biointerface
membranes need not necessarily include interconnected cavities for
creating a fluid pocket in the short-term.
[0202] In certain embodiments, the device includes a physical
spacer between the sensor and the surrounding tissue. A spacer
allows for a liquid sheath to form around at least a portion of the
sensor, such as the area surrounding the electrodes, for example. A
fluid sheath can provide a fluid bulk that dilutes or buffers
interferants while promoting glucose and oxygen transport to the
sensor.
[0203] In some embodiments, the spacer is a mesh or optionally a
fibrous structure. Suitable mesh materials are known in the art and
include open-weave meshes fabricated of biocompatible materials
such as but not limited to PLA, PGA, PP, nylon and the like. Mesh
spacers can be applied directly to the sensing mechanism or over a
biointerface membrane, such as a porous biointerface membrane
disclosed elsewhere herein. Mesh spacers can act as a fluid influx-
or efflux-promoting structure and provides the advantage of
relatively more rapid fluid movement, mixing and/or diffusion
within the mesh to reduce local interferant concentrations and
increasing glucose and oxygen concentrations. The increased fluid
volume within the mesh can also promote increased fluid movement in
and out of the area, which brings in glucose and oxygen while
removing or diluting interferants.
[0204] Furthermore, a physical spacer can reduce the effect of
lymph pooling due to local compression (during sedentary activity)
by mechanically maintaining the fluid pocket. When the host is
sedentary (e.g., lies down to sleep) the area surrounding the
sensor can be compressed. For example, if the sensor is on the
right side of the host's abdomen and he lies down on that side for
a few hours, the lymphatics on the abdominal right side will be
pinched off. When the tissue is compressed/pinched, fluid will not
be able to move into the pinched lymphatic capillaries and
interferants (from local tissue metabolism) can build up and cause
noise. When the host gets up, the compression/pinching is relieved
and the interferants can be removed via the lymphatics. Since a
spacer can maintain the fluid bulk around the sensor during local
compression, the effect of interferant concentration increases can
be suppressed or reduced, thereby reducing noise and promoting
optimal sensor function.
[0205] In one exemplary embodiment, the sensor is wrapped with a
single layer of open weave polypropylene (PP) biocompatible mesh.
When the sensor is inserted, the mesh holds the surrounding tissue
away from the sensor surface and allows an influx of extracellular
fluid to enter the spaces within the mesh, thereby creating a fluid
pocket around the sensor. Within the fluid pocket, fluid can mix
substantially rapidly as extracellular fluid enters and leaves the
fluid pocket or due to host movement. Interferants are carried by
the fluid and therefore can be mixed and/or diluted. Since the host
can wear the sensor for a plurality of days, sedentary periods will
inevitably occur. During these periods interferants can accumulate.
However, the increased fluid volume provided by the mesh can
substantially buffer accumulated interferants until the sedentary
period ends. When the sedentary period is over, any accumulated
interferants can be diluted or carried away by an influx or efflux
of fluid.
[0206] In an alternative embodiment, a mesh can be applied to a
sensor either symmetrically or asymmetrically. For example, the
mesh can be tightly wrapped around the sensor. In another example,
a strip of mesh can be applied to only one side of the sensor. In
yet another example, the mesh can form a flat envelope about a few
millimeters to about a centimeter wide, with the sensor sandwiched
within the envelope. In some embodiments, the mesh can cover only a
portion of the sensor, such as the portion containing the
electrochemically reactive surface(s). In other embodiments, the
mesh can cover the entire sensor.
[0207] In another alternative embodiment, noise can be reduced by
inclusion of a hydrogel on the surface of at least a portion of the
sensor, such as the sensing region. A hydrogel is a network of
super absorbent (they can contain 20%-99% or weight % water,
preferably 80% to over 99% weight % water) natural or synthetic
polymer chains. Hydrogels are sometimes found as a colloidal gel in
which water is the dispersion medium. Since hydrogels are
nonporous, fluid and interferants within the hydrogel move by
diffusion. Accordingly, the movement of molecules within hydrogels
is relatively slower than that possible within mesh-based fluid
pockets as described above. Optionally, the hydrogel can be
biodegradable. A biodegradable hydrogel can provide a fluid pocket
that gradually diminishes and is eventually eliminated by the
surrounding tissue.
[0208] In a further embodiment, a hydrogel includes a flexible,
water-swellable, film (as disclosed elsewhere herein) having a "dry
film" thickness of from about 0.05 micron or less to about 20
microns or more, more preferably from about 0.05, 0.1, 0.15, 0.2,
0.25, 0.3, 0.35, 0.4, 0.45, 0.5, 1, 1.5, 2, 2.5, 3, or 3.5 to about
4, 5, 6, 7, 8, 9, 10, 11, 12, 13, 14, 15, 16, 17, 18, 19, or 19.5
microns, and more preferably from about 2, 2.5 or 3 microns to
about 3.5, 4, 4.5, or 5 microns. "Dry film" thickness refers to the
thickness of a cured film cast from a coating formulation by
standard coating techniques. The hydrogel material can be applied
to the entire sensor or a portion of it, using any method known in
the art, such as but not limited to dipping, painting, spraying,
wrapping, and the like.
[0209] In certain embodiments, scavenging agents (e.g., bioactive
agents that can scavenge, bind-up or substantially inactivate
interferants) can be incorporated into the hydrogel or other aspect
of the device (e.g., membrane system). Scavenging agents can
suppress prolonged wounding and inflammation by removing irritating
substances from the locality of the sensor.
[0210] One exemplary scavenging agent embodiment incorporates an
H.sub.2O.sub.2-degrading enzyme, such as but not limited to
glutathione peroxidase (GSH peroxidase), heme-containing
peroxidases, eosinophil peroxidase, thyroid peroxidase or
horseradish peroxidase (HRP) into the hydrogel to degrade the
available H.sub.2O.sub.2 and produce oxygen. The scavenging agent
can act within the hydrogel or can be released into the local
environment to act outside the hydrogel.
[0211] In a further embodiment, a mesh and a hydrogel can be used
in combination to provide greater mechanical support (to hold the
surrounding tissue away from the sensor) while slowing down the
diffusion rate within the mesh-hydrogel layer. For example, a PP
mesh can be applied to the sensor followed by spraying a dry
hydrogel material onto the PP-wrapped sensor. Alternatively, the
hydrogel can be dried within the mesh before application to the
sensor. Upon sensor implantation, the hydrogel can absorb fluid
from the surrounding tissue, expand and fill the mesh pores. In a
further example, the hydrogel can be biodegradable. In this
example, the hydrogel can initially slow fluid movement. But as the
hydrogel is biodegraded, the pores of the mesh are opened up and
fluid movement can speed up or increase.
[0212] A variety of alternative materials can be used to create
architectures that create a fluid pocket. For example, shape-memory
materials can be used as an alternative to a mesh, to form a fluid
pocket around the sensor. Shape-memory materials are metals or
polymers that "remember" their geometries. Shape-memory metals
(e.g., memory metals or smart wire) include copper-zinc-aluminum,
copper-aluminum-nickel, and nickel-titanium (NiTi) alloys.
Shape-memory polymers include materials such as polynorbornene,
segmented poly(epsilon-caprolactone) polyurethanes, poly(ethylene
glycol)-poly(epsilon-caprolactone) diblock copolymers, and the
like, for example. A shape-memory material can be deformed from its
"original" conformation and regains its original geometry by itself
in response to a force, such as temperature or pressure.
[0213] In one embodiment, a porous memory material that has been
collapsed into a flat, nonporous sheet can be applied to the
exterior of the sensor as a flat film. After insertion into the
body, increased temperature or moisture exposure can stimulate the
memory material to transform to a 3-dimensional, porous
architecture that promotes fluid pocket formation, for example.
[0214] In an alternative embodiment, nanoporous materials, which
act as molecular sieves, can be used to exclude interferants
surrounding the sensor. In another alternative embodiment, a
swellable material (e.g., a material having an initial volume that
absorbs fluid, such as water, when it contacts the fluid to become
a second volume that is greater than the initial volume) or
collapsible material (e.g., a material having an initial volume
that collapse to a second volume that is smaller than the initial
volume) can produce or maintain a fluid pocket.
[0215] In yet another embodiment, materials with differing
characteristics can be applied in combination, such as alternating
bands or layers, to suppress uniform capsule formation. For
example, alternating bands of collapsible and non-collapsible
swellable material can be applied around a portion of the sensor.
Upon implantation, both materials swell with fluid from the
surrounding tissue. However, only the segments of collapsible
material can deform. Since the material surrounding the sensor will
be irregular, it can disrupt formation of a continuous cell layer,
thereby reducing noise and extending sensor life.
Wound Irritation
[0216] Another aspect of the present invention employs wound
irritation either by physical structure or chemical irritants, to
stimulate and/or prolong the wound healing process. Preferably, an
irritating architecture stimulates adjacent cells to release
soluble mediators of wound healing and/or inflammation. The
released soluble mediators are believed to increase the rates of
hemostasis and inflammation (e.g., promoting fluid bulk increase or
an increase in bulk fluid flow) and resulting in dilution/removal
of irritants and noise reduction.
[0217] Accordingly, one embodiment of an irritating biointerface
includes a structure having a roughened surface, which can rub or
poke adjacent cells in vivo. The sensor surface can be roughened by
coating the sensor with a machineable material that is or can be
etched to form ridges, bristles, spikes, grids, grooves, circles,
spirals, dots, bumps, pits or the like, for example. The material
can be any convenient, biocompatible material, such as machined
porous structures that are overlaid on the sensor, such as but not
limited to machineable metal matrix composites, bone substrates
such as hydroxyapatite, coral hydroxyapatite and .beta.-tricalcium
phosphate (TCP), porous titanium (Ti) mixtures made by sintering of
elemental powders, bioglasses (calcium and silicon-based porous
glass), ceramics and the like. The material can be "machined" by
any convenient means, such as but not limited to scraping, etching,
lathing or lasering, for example.
[0218] Micro-motion of the sensor can increase the irritating
effect of a roughened surface. Micro-motion is an inherent property
of any implanted device, such as an implanted glucose sensor.
Micro-motion of the device (e.g., minute movements of the device
within the host) is caused by host movements, ranging from
breathing and small local muscle movements to gross motor
movements, such as walking, running or even getting up and sitting
down. External forces, such as external pressure application, can
also cause micro-motion. Micro-motion includes movement of the
sensor back and forth, rotation, twisting and/or turning.
Accordingly, as the sensor is moved by micro-motion, the sensor's
rough surface can rub more vigorously against the surrounding
tissue, causing increased or extended wounding, resulting in
additional stimulation of the wound healing process and increases
in fluid bulk, bulk fluid flow and/or fluid pocket formation, with
a concomitant reduction in noise.
[0219] In another embodiment, an irritating architecture is formed
from self-assembly materials. Self-assembly biomaterials comprise
specific polypeptides that are designed a priori to self-assemble
into targeted nano- and microscopic structures. Intramolecular
self-assembling molecules are often complex polymers with the
ability to assemble from the random coil conformation into a
well-defined stable structure (secondary and tertiary structure). A
variety of self-assembly materials known in the art can find use in
the present embodiment. For example, PuraMatrix.TM. (3DM Inc.,
Cambridge, Mass., USA) can be used to create synthetic
self-assembling peptide nanofiber scaffolds and defined 3-D
microenvironments.
[0220] In an exemplary embodiment of an irritating biointerface, an
irritating superstructure is applied to the working electrode or
the completed sensor. A "superstructure," as used herein is a broad
term and used in its ordinary sense, including, without limitation,
to refer to any structure built on something else, such as but not
limited to the overlying portion of a structure. An irritating
superstructure can include any substantial structure that prevents
cell attachment and is irritating to the surrounding tissue in
vivo. In one example, an irritating superstructure can include
large spaces, such as at least about 50 .mu.m wide and at least
about 50 .mu.m deep. Cells surrounding the sensor can be prevented
from attachment in the spaces within the superstructure, allowing
fluid to fill these spaces. In some exemplary embodiments, an
irritating superstructure takes advantage of sensor micromotion, to
prevent cell attachment and stimulate fluid pocket formation.
[0221] In one exemplary embodiment, an irritating superstructure is
comprised of ridges at least about 0.25 to 0.50 .mu.m in diameter
and about 50 .mu.m high, and separated by at least about 0.25 to
0.50 .mu.m. In another exemplary embodiment, an exposed silver
wire, at least about 0.25 to 0.50 .mu.m in diameter, is applied to
the sensor exterior to form grooves about 50 .mu.m wide and about
50 .mu.m deep. Since silver is pro-inflammatory and stimulates
fluid influx from the surrounding tissues, the combination of an
irritating superstructure and a chemical irritant could promote an
increased rate of fluid influx or prolong irritation and fluid
influx. In yet another exemplary embodiment, with reference to the
embodiment shown in FIG. 3A, the configuration (e.g., diameter) of
the reference electrode 30 can be changed (e.g., increased in size
and/or coil spacing) such that the reference electrode, itself,
becomes an irritating superstructure, with or without a coating 32
as disclosed elsewhere herein.
[0222] Inflammation and fluid pocket formation can also be induced
by inclusion of irritating chemicals or agents that promote fluid
influx or efflux, vasodilating agents, inflammatory agents,
wounding agents, some wound-healing agents and the like. In some
embodiments, irritation and fluid pocket forming agents can include
but are not limited to enzymes, cytotoxic or necrosing agents
(e.g., pactataxyl, actinomycin, doxorubicin, daunorubicin,
epirubicin, bleomycin, plicamycin, mitomycin), cyclophosphamide,
chlorambucil, uramustine, melphalan, bryostatins, inflammatory
bacterial cell wall components, histamines, pro-inflammatory
factors and the like. Chemical systems/methods of irritation
include any materials that do not adversely affect the performance
or safety of the device such as pro-inflammatory agents. Generally,
pro-inflammatory agents are irritants or other substances that
induce chronic inflammation and chronic granular response at the
wound-site.
[0223] Chemical systems/methods of irritation can be applied to the
exterior of the sensor by any useful means known in the art, such
as by dipping, spraying or painting, for example. In one exemplary
embodiment, the completed sensor is dipped into a dilute solution
of histamine for about five seconds and dried at room temperature.
Upon insertion into a host, the histamine can be solublized and
stimulate an accelerated wound healing response, causing an influx
of fluid and inflammatory cell migration to the sensor within the
first few hours of sensor implantation, such as about 2 to 5 hours,
or longer.
[0224] In another exemplary embodiment, only the operative sensing
portion of the sensor is painted with a dilute necrosing agent
(e.g., compounds that stimulate tissue devitalization), such as
bleomycin) and dried. When the dried sensor is inserted into the
host, the necrosing agent can leach off the sensor and devitalize a
small amount of tissue around the sensing portion of the sensor.
Generally, wound healing rapidly ensues, resulting in
vasodilatation, fluid influx and an influx of macrophages and
polymorphonuclear leukocytes, which remove the devitalized tissue.
The space created by the removal of the devitalized tissue is
filled with fluid and acts as a substantial fluid pocket.
[0225] Chemical systems/methods of irritation can also be
incorporated into an exterior sensor structure, such as the
biointerface membrane (described below) or a shedding layer that
releases the irritating agent into the local environment. For
example, in some embodiments, a "shedding layer" releases (e.g.,
sheds or leaches) molecules into the local vicinity of the sensor
and can speed up osmotic fluid shifts. In some embodiments, a
shedding layer can provide a mild irritation and encourage a mild
inflammatory/foreign body response, thereby preventing cells from
stabilizing and building up an ordered, fibrous capsule and
promoting fluid pocket formation.
[0226] A shedding layer can be constructed of any convenient,
biocompatible material, include but not limited to hydrophilic,
degradable materials such as polyvinylalcohol (PVA), PGC,
Polyethylene oxide (PEO), polyethylene glycol-polyvinylpyrrolidone
(PEG-PVP) blends, PEG-sucrose blends, hydrogels such as
polyhydroxyethyl methacrylate (pHEMA), polymethyl methacrylate
(PMMA) or other polymers with quickly degrading ester linkages. In
certain embodiment, absorbable suture materials, which degrade to
compounds with acid residues, can be used. The acid residues are
chemical irritants that stimulate inflammation and wound healing.
In certain embodiments, these compounds include glycolic acid and
lactic acid based polymers, polyglactin, polydioxone, polydyconate,
poly(dioxanone), poly(trimethylene carbonate) copolymers, and poly
(-caprolactone) homopolymers and copolymers, and the like.
[0227] In other exemplary embodiments, the shedding layer can be a
layer of materials listed elsewhere herein for the first domain,
including copolymers or blends with hydrophilic polymers such as
polyvinylpyrrolidone (PVP), polyhydroxyethyl methacrylate,
polyvinylalcohol, polyacrylic acid, polyethers, such as
polyethylene glycol, and block copolymers thereof including, for
example, di-block, tri-block, alternating, random and graft
copolymers (block copolymers are discussed in U.S. Pat. Nos.
4,803,243 and 4,686,044, hereby incorporated by reference). In one
preferred embodiment, the shedding layer is comprised of
polyurethane and a hydrophilic polymer. For example, the
hydrophilic polymer can be polyvinylpyrrolidone. In one embodiment
of this aspect of the invention, the shedding layer is polyurethane
comprising not less than 5 weight percent polyvinylpyrrolidone and
not more than 45 weight percent polyvinylpyrrolidone. Preferably,
the shedding layer comprises not less than 20 weight percent
polyvinylpyrrolidone and not more than 35 weight percent
polyvinylpyrrolidone and, most preferably, polyurethane comprising
about 27 weight percent polyvinylpyrrolidone.
[0228] In other exemplary embodiments, the shedding layer can
include a silicone elastomer, such as a silicone elastomer and a
poly(ethylene oxide) and poly(propylene oxide) co-polymer blend, as
disclosed in copending U.S. patent application Ser. No. 11/404,417,
filed Apr. 14, 2006 and entitled "SILICONE BASED MEMBRANES FOR USE
IN IMPLANTABLE GLUCOSE SENSORS." In one embodiment, the silicone
elastomer is a dimethyl- and methylhydrogen-siloxane copolymer. In
one embodiment, the silicone elastomer comprises vinyl
substituents. In one embodiment, the silicone elastomer is an
elastomer produced by curing a MED-4840 mixture. In one embodiment,
the copolymer comprises hydroxy substituents. In one embodiment,
the co-polymer is a triblock poly(ethylene oxide)-poly(propylene
oxide)-poly(ethylene oxide) polymer. In one embodiment, the
co-polymer is a triblock poly(propylene oxide)-poly(ethylene
oxide)-poly(propylene oxide) polymer. In one embodiment, the
co-polymer is a PLURONIC.RTM. polymer. In one embodiment, the
co-polymer is PLURONIC.RTM. F-127. In one embodiment, at least a
portion of the co-polymer is cross-linked. In one embodiment, from
about 5% w/w to about 30% w/w of the membrane is the
co-polymer.
[0229] A shedding layer can take any shape or geometry, symmetrical
or asymmetrical, to promote fluid influx in a desired location of
the sensor, such as the sensor head or the electrochemically
reactive surfaces, for example. Shedding layers can be located on
one side of sensor or both sides. In another example, the shedding
layer can be applied to only a small portion of the sensor or the
entire sensor.
[0230] In one exemplary embodiment, a shedding layer comprising
polyethylene oxide (PEO) is applied to the exterior of the sensor,
where the tissue surrounding the sensor can directly access the
shedding layer. PEO leaches out of the shedding layer and is
ingested by local cells that release pro-inflammatory factors. The
pro-inflammatory factors diffuse through the surrounding tissue and
stimulate an inflammation response that includes an influx of
fluid. Accordingly, early noise can be reduced or eliminated and
sensor function can be improved.
[0231] In another exemplary embodiment, the shedding layer is
applied to the sensor in combination with an outer porous layer,
such as a mesh or a porous biointerface as disclosed elsewhere
herein. In one embodiment, local cells access the shedding layer
through the through pores of a porous silicone biointerface. In one
example, the shedding layer material is applied to the sensor prior
to application of the porous silicone. In another example, the
shedding layer material can be absorbed into the lower portion of
the porous silicone (e.g., the portion of the porous silicone that
will be proximal to the sensor after the porous silicone has been
applied to the sensor) prior to application of the porous silicone
to the sensor.
Vasodilatation
[0232] As discussed elsewhere herein, increased fluid bulk, bulk
fluid flow and/or diffusion rates can reduce local interferant
concentrations (e.g., electroactive species produced via cellular
metabolism in the local area) and promote glucose and oxygen influx
or transport, thereby reducing noise frequency or amplitude and
improving early sensor performance. In addition to the structural
and chemical systems/methods discussed above, increased fluid bulk,
fluid bulk flow and/or diffusion rates can be promoted by
vasodilation. Vasodilation occurs when the tight junctions of the
endothelial layer of the microvasculature open. This allows serum
and certain inflammatory cells to leave the circulatory system and
enter the extracellular matrix (ECM). A portion of the fluid in the
ECM can move back into the vasculature. Another portion of the ECM
fluid can leave the area via the lymphatics. Vasodilation promotes
"bulk fluid transport" (e.g., bulk fluid flow) in and out of the
local region and/or increase in fluid bulk around at least a
portion of the sensor. Increased fluid bulk and/or bulk fluid
transport ensures homeostasis with the local environment and the
blood system. Furthermore, rapid diffusion of solutes may be
facilitated by permeabilization of the blood vessels and increased
local temperature due to inflammation. Fluids leaving the local
extracellular spaces remove metabolites, such as the interferants
discussed herein. Preferably, as interferants are carried with the
moving fluid, noise is reduced and sensor function improved.
[0233] In some exemplary embodiments of the present invention,
bioactive agents that promote vasodilation are included in sensor
construction. In one example, the bioactive agents promote an
influx of fluid, causing an increase in fluid bulk. Noise is
reduced as the larger fluid volume reduces the interferant
concentration. In another example, the bioactive agents promote an
efflux of fluid out of the local area. Noise is reduced as the
leaving fluid carries interferants away with it.
[0234] A variety of bioactive agents can be found useful in
preferred embodiments. Exemplary bioactive agents include but are
not limited to blood-brain barrier disruptive agents and
vasodilating agents, such as mannitol, sodium thiosulfate,
VEGF/VPF, NO, NO-donors, leptin, bradykinin, histamines, blood
components, platelet rich plasma (PRP) and the like.
[0235] Bioactive agents can be added during manufacture of the
sensor by incorporating the desired bioactive agent in the
manufacturing material for one or more sensor layers or into an
exterior biomaterial, such as a porous silicone membrane. For
example, bioactive agents can be mixed with a solution during
membrane formation, which is subsequently applied onto the sensor
during manufacture. Alternatively, the completed sensor can be
dipped into or sprayed with a solution of a bioactive agent, for
example. The amount of bioactive agent can be controlled by varying
its concentration, varying the indwell time during dipping,
applying multiple layers until a desired thickness is reached, and
the like, as disclosed elsewhere herein.
[0236] VEGF is a bioactive agent that is known to be a vasodilator
and can promote fluid influx from the microvasculature. In one
embodiment, VEGF is sprayed onto the exterior of the completed
sensor. After insertion, VEGF is directly released into the local
environment when the VEGF-coated sensor is implanted into a host.
The released VEGF stimulates vasodilation around the implanted
sensor. In another embodiment, VEGF is mixed into the biointerface
material prior to sensor construction. After the sensor is
inserted, VEGF leaches form the biointerface, causing vasodilation
around the sensor. In an alternative embodiment, upstream or
downstream components of the VEGF signaling cascade can be
incorporated into the sensor, to affect vasodilatation around the
implanted sensor.
[0237] In one embodiment, biodegradable or bioerodible material can
be employed to release bioactive agents in a controlled manner. In
one exemplary embodiment, VEGF is incorporated into a biodegradable
material (e.g., shedding layer or hydrogel) that is applied to the
sensor exterior. Upon implantation, the surrounding tissue begins
to degrade the biodegradable material. As the material degrades,
VEGF is released into the local environment in a desired
rate-limiting manner. The rate of bioactive agent release (e.g.,
VEGF) can be manipulated by the selection of the biodegradable
material and the thickness of the biodegradable material layer.
Thus, constant bioactive agent release can be achieved for a
predetermined extended period of time and possibly promote
vasodilatation and fluid influx during that period of time.
[0238] In an alternative embodiment, the bioactive agent is
microencapsulated before application to the sensor. For example,
microencapsulated VEGF can be sprayed onto a completed sensor or
incorporated into a structure, such as an outer mesh layer or a
shedding layer. Microencapsulation can offer increased flexibility
in controlling bioactive agent release rate, time of release
occurrence and/or release duration.
[0239] In still another embodiment, vasodilation is achieved by
matrix metalloproteinases (MMP) incorporation into the sensor. MMPs
can degrade the proteins that keep blood vessel walls solid. This
proteolysis allows endothelial cells to escape into the
interstitial matrix and concomitantly fluid to enter and leave the
vasculature. Accordingly, MMPs can promote interferant
concentration reduction and intermittent, sedentary noise reduction
or elimination.
[0240] In another embodiment, angiogenic and/or preangiogenic
compounds or factors are included in the sensor to promote
vasodilation. Angiogenesis is the physiological process involving
the growth of new blood vessels from pre-existing vessels.
Formation of new vessels can reduce the frequency or magnitude of
intermittent, sedentary noise by increasing fluid flow, for
example. Angiogenic agents include, but are not limited to, Basic
Fibroblast Growth Factor (bFGF), (also known as Heparin Binding
Growth Factor-II and Fibroblast Growth Factor II), Acidic
Fibroblast Growth Factor (aFGF), (also known as Heparin Binding
Growth Factor-I and Fibroblast Growth Factor-I), Vascular
Endothelial Growth Factor (VEGF), Platelet Derived Endothelial Cell
Growth Factor BB (PDEGF-BB), Angiopoietin-1, Transforming Growth
Factor Beta (TGF-Beta), Transforming Growth Factor Alpha
(TGF-Alpha), Hepatocyte Growth Factor, Tumor Necrosis Factor-Alpha
(TNF-Alpha), Placental Growth Factor (PLGF), Angiogenin,
Interleukin-8 (IL-8), Hypoxia Inducible Factor-I (HIF-1),
Angiotensin-Converting Enzyme (ACE) Inhibitor Quinaprilat,
Angiotropin, Thrombospondin, Peptide KGHK, Low Oxygen Tension,
Lactic Acid, Insulin, Leptin, Copper Sulphate, Estradiol,
prostaglandins, cox inhibitors, endothelial cell binding agents
(for example, decorin or vimentin), glenipin, hydrogen peroxide,
nicotine, and Growth Hormone.
Wound Suppression
[0241] Wound suppression to reduce noise is an alternative aspect
of the preferred embodiment. Wound suppression includes any systems
or methods by which an amount of wounding that occurs upon sensor
insertion is reduced and/or eliminated. While not wishing to be
bound by theory, it is believed that if wounding is suppressed or
at least significantly reduced, the sensor will be surrounded by
substantially normal tissue (e.g., tissue that is substantially
similar to the tissue prior to sensor insertion). Substantially
normal tissue is believed to have a lower metabolism than wounded
tissue, producing fewer interferants and reducing early noise.
[0242] Wounds can be suppressed or minimized by adaptation of the
sensor's architecture to one that either suppresses wounding or
promotes rapid healing, such as an architecture that does not cause
substantial wounding (e.g., an architecture configured to prevent
wounding), an architecture that promotes wound healing, an
anti-inflammatory architecture, and the like. In one exemplary
embodiment, the sensor is configured to have a low profile, a
zero-footprint or a smooth surface. For example, the sensor can be
formed of substantially thin wires, such as wires about 50-150
.mu.m in diameter, for example. Preferably, the sensor is small
enough to fit within a very small gauge needle, such as a 30, 31,
32, 33, 34, or 35-gauge needle (or smaller) on the Stubs scale, for
example. In general, a smaller needle, the more reduces the amount
of wounding during insertion. For example, a very small needle can
reduce the amount of tissue disruption and thereby reduce the
subsequent wound healing response. In an alternative embodiment,
the sensor's surface is smoothed with a lubricious coating, to
reduce wounding upon sensor insertion.
[0243] Wounding can also be reduced by inclusion of
wound-suppressive agents that either reduce the amount of initial
wounding or suppress the wound healing process. While not wishing
to be bound by theory, it is believed that application of a
wound-suppressing agent, such as an anti-inflammatory, an
immunosuppressive agent, an anti-infective agent, or a scavenging
agent, to the sensor can create a locally quiescent environment and
suppress wound healing. In a quiescent environment, bodily
processes, such as the increased cellular metabolism associated
with wound healing, can minimally affect the sensor. If the tissue
surrounding the sensor is undisturbed, it can continue its normal
metabolism and promote sensor function.
[0244] It has been observed that anti-histamines can suppress or
eliminate early sedentary noise. Namely, it has been shown that
oral anti-histamines taken at nighttime can result in substantially
diminished early sedentary noise. While not wishing to be bound by
theory, it is believed that histamines, which are chemicals
released during wounding, produce electrochemical interference in
the sensor signal. Namely, histamine release is believed to promote
release of electrochemical interferants, which in certain
circumstances produce "noise" on the sensor signal.
[0245] Further, the inventors believe that during sedentary periods
(e.g., sleeping or long periods of sitting) host immobility can
cause local pooling of the interstitial and/or lymph fluids, which
results in a back up of lymph surrounding the sensor with a
corresponding build-up of electroactive species as a result of
normal cellular metabolism. Pooling of the wound fluid around the
sensor can suppress the usual movement of the wound fluids that
would enable accurate analyte measurement. The lack of fluid
movement results in modified sample fluid, including but not
limited to a local increase in electroactive species (e.g.,
histamines or other resulting products) during these periods of
intermittent sedentary noise. When the host moves or shifts body
position, the fluid is released and fluid flow is restored,
allowing an influx of oxygen and glucose and removal of
electroactive metabolic species (e.g., interfering species).
[0246] Accordingly, one embodiment of the present invention
provides for a sensor including an anti-histamine. Anti-histamines
are any drugs that serve to reduce or eliminate the effects
mediated by histamine. Some examples of conventional
anti-histamines suitable for incorporation into or onto the present
invention include, but are not limited to first-generation
H.sub.1-receptor antagonists: ethylenediamines (e.g., mepyramine
(pyrilamine), antazoline), ethanolamines (e.g., diphenhydramine,
carbinoxamine, doxylamine, clemastine, and dimenhydrinate),
alkylamines (pheniramine, chlorphenamine (chlorpheniramine),
dexchlorphenamine, brompheniramine, and triprolidine), piperazines
(cyclizine, hydroxyzine, and meclizine), and tricyclics
(promethazine, alimemazine (trimeprazine), cyproheptadine, and
azatadine).
[0247] Additionally, Second-generation H.sub.1-receptor antagonists
are newer antihistamine drugs that are much more selective for
peripheral H.sub.1 receptors in preference to the central nervous
system histaminergic and cholinergic receptors. This selectivity
significantly reduces the occurrence of adverse drug reactions
compared with first-generation agents, while still providing
effective relief of allergic conditions. Both systemic
(acrivastine, astemizole, cetirizine, loratadine, mizolastine) and
topical (azelastine, levocabastine, and olopatadine) could be
used.
[0248] In some alternative embodiments, other inhibitors of
histamine release, which appear to stabilize the mast cells to
suppress degranulation and mediator release, can be used (e.g.,
cromoglicate (cromolyn) and nedocromil).
[0249] Anti-histamine can be incorporated into the sensor by any
convenient system or technique known to those skilled in the art.
In one exemplary embodiment, anti-histamine is incorporated into a
biodegradable shedding layer. As the shedding layer is degraded,
the anti-histamine is released into the surrounding area, to
suppress histamine release and down-stream inflammation processes,
thereby suppressing interferant build up and improving sensor
function. In another exemplary embodiment, anti-histamine is
sprayed on the surface of the completed sensor and dried. Upon
insertion, the anti-histamine is solublized, suppresses histamine
production and downstream inflammation mediators, thereby reducing
noise.
[0250] Other agents that suppress the body's response to wounding
can also be incorporated into the sensor or the present invention.
In one embodiment, wounding can be suppressed by the inclusion of
anti-inflammatory agents. Generally, anti-inflammatory agents
reduce acute and/or chronic inflammation adjacent to the implant,
in order to decrease the formation of a FBC capsule to reduce or
prevent barrier cell layer formation. Suitable anti-inflammatory
agents include but are not limited to, for example, nonsteroidal
anti-inflammatory drugs (NSAIDs) such as acetometaphen,
aminosalicylic acid, aspirin, celecoxib, choline magnesium
trisalicylate, diclofenac potassium, diclofenac sodium, diflunisal,
etodolac, fenoprofen, flurbiprofen, ibuprofen, indomethacin,
interleukin (IL)-10, IL-6 mutein, anti-IL-6 iNOS inhibitors (for
example, L-NAME or L-NMDA), Interferon, ketoprofen, ketorolac,
leflunomide, melenamic acid, mycophenolic acid, mizoribine,
nabumetone, naproxen, naproxen sodium, oxaprozin, piroxicam,
rofecoxib, salsalate, sulindac, and tolmetin; and corticosteroids
such as cortisone, hydrocortisone, methylprednisolone, prednisone,
prednisolone, betamethesone, beclomethasone dipropionate,
budesonide, dexamethasone sodium phosphate, flunisolide,
fluticasone propionate, paclitaxel, tacrolimus, tranilast,
triamcinolone acetonide, betamethasone, fluocinolone, fluocinonide,
betamethasone dipropionate, betamethasone valerate, desonide,
desoximetasone, fluocinolone, triamcinolone, triamcinolone
acetonide, clobetasol propionate, and dexamethasone.
[0251] In one example, glucocorticoids stimulate the movement of
lipocortin-1 into the extracellular space, where it binds to
leukocyte membrane receptors and inhibits various inflammatory
events: such as epithelial adhesion, emigration, chemotaxis,
phagocytosis, respiratory burst and the release of various
inflammatory mediators (lysosomal enzymes, cytokines, tissue
plasminogen activator, chemokines etc.) from neutrophils,
macrophages and mastocytes.
[0252] In one exemplary embodiment, the sensor is coated with
dexamethasone or dexamethasone is incorporated into a protective
layer or film, such as a hydrophilic silicone protective film. In
vivo, the dexamethasone is released from the surface of the sensor
and interacts with the surrounding tissue, thereby reducing or
eliminating local inflammation and early noise.
[0253] In another embodiment, an immunosuppressive and/or
immunomodulatory agent is included in the sensor to suppress wound
healing and/or fluid pocket formation, thereby reducing noise.
Generally, immunosuppressive and/or immunomodulatory agents
interfere directly with several key mechanisms necessary for
involvement of different cellular elements in the inflammatory
response. Suitable immunosuppressive and/or immunomodulatory agents
include anti-proliferative, cell-cycle inhibitors, (for example,
paclitaxel, cytochalasin D, infiximab), taxol, actinomycin,
mitomycin, thospromote VEGF, estradiols, NO donors, QP-2,
tacrolimus, tranilast, actinomycin, everolimus, methothrexate,
mycophenolic acid, angiopeptin, vincristing, mitomycine, statins, C
MYC antisense, sirolimus (and analogs), RestenASE,
2-chloro-deoxyadenosine, PCNA Ribozyme, batimstat, prolyl
hydroxylase inhibitors, PPAR.gamma. ligands (for example
troglitazone, rosiglitazone, pioglitazone), halofuginone,
C-proteinase inhibitors, probucol, BCP671, EPC antibodies,
catchins, glycating agents, endothelin inhibitors (for example,
Ambrisentan, Tesosentan, Bosentan), Statins (for example,
Cerivasttin), E. coli heat-labile enterotoxin, and advanced
coatings. While not wishing to be bound by theory, it is believed
that inflammation suppression will promote a quiescent sensor
environment and a substantially normal local metabolism.
[0254] In another embodiment, the biointerface comprises a
pro-inflammatory architecture configured to promote substantially
rapid fluid influx (e.g., due to inflammation and the like) after
sensor insertion followed by an extended quiescent period (e.g.,
during which wound healing is suppressed). While not wishing to be
bound by theory, it is believed that, within a wound healing time
line, a brief period of inflammation, followed by wound healing
suppression can promote initial fluid pocket formation with
subsequent suppression of interferent concentration increase.
Initial fluid pocket formation can facilitate analyte (e.g.,
glucose) and oxygen transport from the surrounding tissues to the
working electrode. The subsequent suppression of interferent
build-up can reduce noise.
[0255] In yet another embodiment, an anti-infective agent is
incorporated into the sensor, to prevent a local infection that
would stimulate inflammation around the sensor. Accordingly, the
inflammation signal cascade and concomitant metabolic changes will
be suppressed, resulting in noise suppression. Generally,
anti-infective agents are substances capable of acting against
infection by inhibiting the spread of an infectious agent or by
killing the infectious agent outright, which can serve to reduce
immuno-response without inflammatory response at the implant site.
Anti-infective agents include, but are not limited to,
anthelmintics (mebendazole), antibiotics including aminoclycosides
(gentamicin, neomycin, tobramycin), antifungal antibiotics
(amphotericin b, fluconazole, griseofulvin, itraconazole,
ketoconazole, nystatin, micatin, tolnaftate), cephalosporins
(cefaclor, cefazolin, cefotaxime, ceftazidime, ceftriaxone,
cefuroxime, cephalexin), beta-lactam antibiotics (cefotetan,
meropenem), chloramphenicol, macrolides (azithromycin,
clarithromycin, erythromycin), penicillins (penicillin G sodium
salt, amoxicillin, ampicillin, dicloxacillin, nafcillin,
piperacillin, ticarcillin), tetracyclines (doxycycline,
minocycline, tetracycline), bacitracin; clindamycin; colistimethate
sodium; polymyxin b sulfate; vancomycin; antivirals including
acyclovir, amantadine, didanosine, efavirenz, foscarnet,
ganciclovir, indinavir, lamivudine, nelfinavir, ritonavir,
saquinavir, silver, stavudine, valacyclovir, valganciclovir,
zidovudine; quinolones (ciprofloxacin, levofloxacin); sulfonamides
(sulfadiazine, sulfisoxazole); sulfones (dapsone); furazolidone;
metronidazole; pentamidine; sulfanilamidum crystallinum;
gatifloxacin; and sulfamethoxazole/trimethoprim. While not wishing
to be bound by theory, it is believed that infection suppression
will promote a quiescent sensor environment and a substantially
normal local metabolism.
[0256] In another embodiment, interferant scavengers can be applied
to the sensor, to remove electroactive interferants. While not
wishing to be bound by theory, it is believed that removal of
electroactive interferants around the sensor can reduce early noise
and promote increased sensor sensitivity during the first few hours
or days of sensor use. Interferant scavengers can include enzymes,
such as superoxide dismutase (SOD), thioredoxin, glutathione
peroxidase and catalase, anti-oxidants, such as uric acid and
vitamin C, iron compounds, Heme compounds, and some heavy metals.
In one exemplary embodiment, a hydrogel containing SOD and
horseradish peroxidase (HRP) is coated on the surface of the
sensor. After sensor implantation, the SOD decomposes superoxide
radicals from the surrounding cells into O.sub.2 and
H.sub.2O.sub.2. The H.sub.2O.sub.2 is subsequently broken down into
water by HRP. Thus, electroactive interferants, such as superoxide
and hydrogen peroxide, can be removed and oxygen provided for the
glucose oxidase component of the sensor.
[0257] In another embodiment, at least a portion of the sensor is
coated with an artificial protective coating to reduce wounding.
The term "protective coating" as used herein is a broad term and is
used in its ordinary sense, including, without limitation, a
coating of proteins and other molecules, such as those found in
serous fluid. While not wishing to be bound by theory, it is
believed that after the first about 2 to about 36 hours of sensor
insertion, the host's biological processes provide a protective
coating surrounding the sensor that protects the sensor from these
endogenous interferants or other in vivo effects. In one exemplary
embodiment, at least a portion of the sensor is coated with an
artificial protective coating. The artificial protective coating
components can include but are not limited to albumin, fibrin,
collagen, endothelial cells, wound closure chemicals, blood
products, platelet-rich plasma, growth factors and the like. A
protective coating can be applied to the sensor in any convenient
way, such as but not limited to dipping the sensor into a mixture
of protective coating components, spraying or incorporating the
protective coating components into a biointerface membrane.
Advantageously, a protective film can prevent sensor degradation
associated with the local environment and promote integration of
the biointerface into the surrounding tissue.
[0258] In a further embodiment, a silicone coating or hydrophilic
shedding layer can be applied to the sensor. While not wishing to
be bound by theory, it is believed that a silicone bioprotective
coating or shedding layer can promote formation and maintenance of
a fluid pocket around the sensor, to enhance glucose and fluid
transport as well as clearance of interferants. A silicone
bioprotective coating can create a local environment with enhanced
vascular permeability and/or vascularization. Such a coating is
believed to speed up the inflammatory response to achieve a
substantially consistent wound environment more quickly than
without the coating. Furthermore, a silicone bioprotective coating
is believed to be able to subdue the inflammatory response to
reduce production of cellular byproducts that are believed to be
electrochemical interferants.
[0259] In one embodiment, a silicone bioprotective coating can
consist of one or more layer(s) formed from a composition that, in
addition to providing high oxygen solubility, allows for the
transport of glucose or other such water-soluble molecules (for
example, drugs). In one embodiment, these layers comprise a blend
of a silicone polymer with a hydrophilic polymer. By "hydrophilic
polymer," it is meant that the polymer has a substantially
hydrophilic domain in which aqueous substances can easily dissolve.
In one embodiment, the hydrophilic polymer has a molecular weight
of at least about 1000 g/mol, 5,000 g/mol, 8,000 g/mol, 10,000
g/mol, or 15,000 g/mol. In one embodiment, the hydrophilic polymer
comprises both a hydrophilic domain and a partially hydrophobic
domain (e.g., a copolymer). The hydrophobic domain(s) facilitate
the blending of the hydrophilic polymer with the hydrophobic
silicone polymer. In one embodiment, the hydrophobic domain is
itself a polymer (i.e., a polymeric hydrophobic domain). For
example, in one embodiment, the hydrophobic domain is not a simple
molecular head group but is rather polymeric. In various
embodiments, the molecular weight of any covalently continuous
hydrophobic domain within the hydrophilic polymer is at least about
500 g/mol, 700 g/mol, 1000 g/mol, 2000 g/mol, 5000 g/mol, or 8,000
g/mol. In various embodiments, the molecular weight of any
covalently continuous hydrophilic domain within the hydrophilic
polymer is at least about 500 g/mol, 700 g/mol, 1000 g/mol, 2000
g/mol, 5000 g/mol, or 8,000 g/mol. In various embodiments, the
layers comprise a blend of a silicone polymer with a hydrophilic
polymer as disclosed in copending U.S. patent application Ser. No.
11/404,417, filed Apr. 14, 2006 and entitled "SILICONE BASED
MEMBRANES FOR USE IN IMPLANTABLE GLUCOSE SENSORS."
[0260] Many of the above disclosed methods and structures for
forming a fluid pocket, diluting interferants, reducing noise and
the like can be used in combination to facilitate a desired effect
or outcome. For example, in one embodiment, a shedding layer
composed of a hydrophilic silicone film and a necrosing agent can
be applied in combination to at least a portion of the sensor. The
silicone film can suppress protein adherence to the sensor surface
while the necrosing agent can devitalize a small portion of tissue
adjacent to the sensor, stimulating formation of a fluid pocket
around the hydrophilic silicone film. Preferably, the increased
volume of fluid surrounding the sensor dilutes interferants while
the shedding layer provides a physical separation between the
sensor and the surrounding tissue.
[0261] In another exemplary embodiment, a mesh sprayed with
dexamethasone is wrapped around the exterior of the sensor. The
mesh can provide a physical spacer for a fluid pocket while the
dexamethasone inhibits inflammation. Preferably, fluid can fill the
mesh and the dexamethasone can promote normal tissue metabolism
around the sensor by inhibiting an influx of inflammatory cells.
Consequently, glucose and oxygen can travel freely between the
tissue and the sensor through the fluid filled mesh without a
buildup of interferants, even during periods of tissue compression,
thereby promoting sensor sensitivity and thereby reducing
noise.
Sensing Mechanism
[0262] In general, the analyte sensors 34 of the preferred
embodiments include a sensing mechanism 36 with a small structure
(e.g., small structured-, micro- or small diameter sensor, see FIG.
3A), for example, a needle-type sensor, in at least a portion
thereof. As used herein a "small structure" preferably refers to an
architecture with at least one dimension less than about 1 mm. The
small structured sensing mechanism can be wire-based, substrate
based, or any other architecture. In some alternative embodiments,
the term "small structure" can also refer to slightly larger
structures, such as those having their smallest dimension being
greater than about 1 mm, however, the architecture (e.g., mass or
size) is designed to minimize the foreign body response due to size
and/or mass. In the preferred embodiments, a biointerface membrane
is formed onto the sensing mechanism 36 as described in more detail
below.
[0263] FIG. 3A is an expanded view of an exemplary embodiment of a
continuous analyte sensor 34, also referred to as a transcutaneous
analyte sensor, or needle-type sensor, particularly illustrating
the sensing mechanism 36. Preferably, the sensing mechanism
comprises a small structure as defined herein and is adapted for
insertion under the host's skin, and the remaining body of the
sensor (e.g., electronics, etc) can reside ex vivo. In the
illustrated embodiment, the analyte sensor 34, includes two
electrodes, i.e., a working electrode 38 and at least one
additional electrode 30, which can function as a counter and/or
reference electrode, hereinafter referred to as the reference
electrode 30.
[0264] In some exemplary embodiments, each electrode is formed from
a fine wire with a diameter of from about 0.001 or less to about
0.010 inches or more, for example, and is formed from, e.g., a
plated insulator, a plated wire, or bulk electrically conductive
material. Although the illustrated electrode configuration and
associated text describe one preferred method of forming a
transcutaneous sensor, a variety of known transcutaneous sensor
configurations can be employed with the transcutaneous analyte
sensor system of the preferred embodiments, such as are described
in U.S. Pat. No. 6,695,860 to Ward et al., U.S. Pat. No. 6,565,509
to Say et al., U.S. Pat. No. 6,248,067 to Causey III, et al., and
U.S. Pat. No. 6,514,718 to Heller et al.
[0265] In preferred embodiments, the working electrode comprises a
wire formed from a conductive material, such as platinum,
platinum-iridium, palladium, graphite, gold, carbon, conductive
polymer, alloys, or the like. Although the electrodes can by formed
by a variety of manufacturing techniques (bulk metal processing,
deposition of metal onto a substrate, or the like), it can be
advantageous to form the electrodes from plated wire (e.g.,
platinum on steel wire) or bulk metal (e.g., platinum wire). It is
believed that electrodes formed from bulk metal wire provide
superior performance (e.g., in contrast to deposited electrodes),
including increased stability of assay, simplified
manufacturability, resistance to contamination (e.g., which can be
introduced in deposition processes), and improved surface reaction
(e.g., due to purity of material) without peeling or
delamination.
[0266] The working electrode 38 is configured to measure the
concentration of an analyte. In an enzymatic electrochemical sensor
for detecting glucose, for example, the working electrode measures
the hydrogen peroxide produced by an enzyme catalyzed reaction of
the analyte being detected and creates a measurable electronic
current. For example, in the detection of glucose wherein glucose
oxidase produces hydrogen peroxide as a byproduct, hydrogen
peroxide 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.
[0267] The working electrode 38 is covered with an insulating
material, for example, a non-conductive polymer. Dip-coating,
spray-coating, vapor-deposition, or other coating or deposition
techniques can be used to deposit the insulating material on the
working electrode. In one embodiment, the insulating material
comprises parylene, which can be an advantageous polymer coating
for its strength, lubricity, and electrical insulation properties.
Generally, parylene is produced by vapor deposition and
polymerization of para-xylylene (or its substituted derivatives).
However, any suitable insulating material can be used, for example,
fluorinated polymers, polyethyleneterephthalate, polyurethane,
polyimide, other nonconducting polymers, or the like. Glass or
ceramic materials can also be employed. Other materials suitable
for use include surface energy modified coating systems such as are
marketed under the trade names AMC18, AMC148, AMC141, and AMC321 by
Advanced Materials Components Express of Bellafonte, Pa. In some
alternative embodiments, however, the working electrode cannot
require a coating of insulator.
[0268] Preferably, the reference electrode 30, which can function
as a reference electrode alone, or as a dual reference and counter
electrode, is formed from silver, silver/silver chloride, or the
like. Preferably, the electrodes are juxtapositioned and/or twisted
with or around each other; however other configurations are also
possible. In one example, the reference electrode 30 is helically
wound around the working electrode 38 as illustrated in FIG. 3A.
The assembly of wires can then be optionally coated together with
an insulating material, similar to that described above, in order
to provide an insulating attachment (e.g., securing together of the
working and reference electrodes).
[0269] In embodiments wherein an outer insulator is disposed, a
portion of the coated assembly structure can be stripped or
otherwise removed, for example, by hand, excimer lasing, chemical
etching, laser ablation, grit-blasting (e.g., with sodium
bicarbonate or other suitable grit), or the like, to expose the
electroactive surfaces. Alternatively, a portion of the electrode
can be masked prior to depositing the insulator in order to
maintain an exposed electroactive surface area. In one exemplary
embodiment, grit blasting is implemented to expose the
electroactive surfaces, preferably utilizing a grit material that
is sufficiently hard to ablate the polymer material, while being
sufficiently soft so as to minimize or avoid damage to the
underlying metal electrode (e.g., a platinum electrode). Although a
variety of "grit" materials can be used (e.g., sand, talc, walnut
shell, ground plastic, sea salt, and the like), in some preferred
embodiments, sodium bicarbonate is an advantageous grit-material
because it is sufficiently hard to ablate, e.g., a parylene coating
without damaging, e.g., an underlying platinum conductor. One
additional advantage of sodium bicarbonate blasting includes its
polishing action on the metal as it strips the polymer layer,
thereby eliminating a cleaning step that might otherwise be
necessary.
[0270] In some embodiments, a radial window is formed through the
insulating material to expose a circumferential electroactive
surface of the working electrode. Additionally, sections of
electroactive surface of the reference electrode are exposed. For
example, the sections of electroactive surface can be masked during
deposition of an outer insulating layer or etched after deposition
of an outer insulating layer. In some applications, cellular attack
or migration of cells to the sensor can cause reduced sensitivity
and/or function of the device, particularly after the first day of
implantation. However, when the exposed electroactive surface is
distributed circumferentially about the sensor (e.g., as in a
radial window), the available surface area for reaction can be
sufficiently distributed so as to minimize the effect of local
cellular invasion of the sensor on the sensor signal.
Alternatively, a tangential exposed electroactive window can be
formed, for example, by stripping only one side of the coated
assembly structure. In other alternative embodiments, the window
can be provided at the tip of the coated assembly structure such
that the electroactive surfaces are exposed at the tip of the
sensor. Other methods and configurations for exposing electroactive
surfaces can also be employed.
[0271] Preferably, the above-exemplified sensor has an overall
diameter of not more than about 0.020 inches (about 0.51 mm), more
preferably not more than about 0.018 inches (about 0.46 mm), and
most preferably not more than about 0.016 inches (0.41 mm). In some
embodiments, the working electrode has a diameter of from about
0.001 inches or less to about 0.010 inches or more, preferably from
about 0.002 inches to about 0.008 inches, and more preferably from
about 0.004 inches to about 0.005 inches. The length of the window
can be from about 0.1 mm (about 0.004 inches) or less to about 2 mm
(about 0.078 inches) or more, and preferably from about 0.5 mm
(about 0.02 inches) to about 0.75 mm (0.03 inches). In such
embodiments, the exposed surface area of the working electrode is
preferably from about 0.000013 in.sup.2 (0.0000839 cm.sup.2) or
less to about 0.0025 in.sup.2 (0.016129 cm.sup.2) or more (assuming
a diameter of from about 0.001 inches to about 0.010 inches and a
length of from about 0.004 inches to about 0.078 inches). The
preferred exposed surface area of the working electrode is selected
to produce an analyte signal with a current in the picoAmp range,
such as is described in more detail elsewhere herein. However, a
current in the picoAmp range can be dependent upon a variety of
factors, for example the electronic circuitry design (e.g., sample
rate, current draw, A/D converter bit resolution, etc.), the
membrane system (e.g., permeability of the analyte through the
membrane system), and the exposed surface area of the working
electrode. Accordingly, the exposed electroactive working electrode
surface area can be selected to have a value greater than or less
than the above-described ranges taking into consideration
alterations in the membrane system and/or electronic circuitry. In
preferred embodiments of a glucose sensor, it can be advantageous
to minimize the surface area of the working electrode while
maximizing the diffusivity of glucose in order to optimize the
signal-to-noise ratio while maintaining sensor performance in both
high and low glucose concentration ranges.
[0272] In some alternative embodiments, the exposed surface area of
the working (and/or other) electrode can be increased by altering
the cross-section of the electrode itself. For example, in some
embodiments the cross-section of the working electrode can be
defined by a cross, star, cloverleaf, ribbed, dimpled, ridged,
irregular, or other non-circular configuration; thus, for any
predetermined length of electrode, a specific increased surface
area can be achieved (as compared to the area achieved by a
circular cross-section). Increasing the surface area of the working
electrode can be advantageous in providing an increased signal
responsive to the analyte concentration, which in turn can be
helpful in improving the signal-to-noise ratio, for example.
[0273] In some alternative embodiments, additional electrodes can
be included within the assembly, for example, a three-electrode
system (working, reference, and counter electrodes) and/or an
additional working electrode (e.g., an electrode which can be used
to generate oxygen, which is configured as a baseline subtracting
electrode, or which is configured for measuring additional
analytes). U.S. Pat. No. 7,081,195 and U.S. Publication No.
US-2005-0143635-A1 describe some systems and methods for
implementing and using additional working, counter, and/or
reference electrodes. In one implementation wherein the sensor
comprises two working electrodes, the two working electrodes are
juxtapositioned (e.g., extend parallel to each other), around which
the reference electrode is disposed (e.g., helically wound). In
some embodiments wherein two or more working electrodes are
provided, the working electrodes can be formed in a double-,
triple-, quad-, etc. helix configuration along the length of the
sensor (for example, surrounding a reference electrode, insulated
rod, or other support structure.) The resulting electrode system
can be configured with an appropriate membrane system, wherein the
first working electrode is configured to measure a first signal
comprising glucose and baseline and the additional working
electrode is configured to measure a baseline signal consisting of
baseline only (e.g., configured to be substantially similar to the
first working electrode without an enzyme disposed thereon.) In
this way, the baseline signal can be subtracted from the first
signal to produce a glucose-only signal that is substantially not
host to fluctuations in the baseline and/or interfering species on
the signal. Accordingly, the above-described dimensions can be
altered as desired. Although the preferred embodiments illustrate
one electrode configuration including one bulk metal wire helically
wound around another bulk metal wire, other electrode
configurations are also contemplated. In an alternative embodiment,
the working electrode comprises a tube with a reference electrode
disposed or coiled inside, including an insulator there between.
Alternatively, the reference electrode comprises a tube with a
working electrode disposed or coiled inside, including an insulator
there between. In another alternative embodiment, a polymer (e.g.,
insulating) rod is provided, wherein the electrodes are deposited
(e.g., electro-plated) thereon. In yet another alternative
embodiment, a metallic (e.g., steel) rod is provided, coated with
an insulating material, onto which the working and reference
electrodes are deposited. In yet another alternative embodiment,
one or more working electrodes are helically wound around a
reference electrode.
[0274] While the methods of preferred embodiments are especially
well suited for use with small structured-, micro- or small
diameter sensors, the methods can also be suitable for use with
larger diameter sensors, e.g., sensors of 1 mm to about 2 mm or
more in diameter.
[0275] In some alternative embodiments, the sensing mechanism
includes electrodes deposited on a planar substrate, wherein the
thickness of the implantable portion is less than about 1 mm, see,
for example U.S. Pat. No. 6,175,752 to Say et al. and U.S. Pat. No.
5,779,665 to Mastrototaro et al., both of which are incorporated
herein by reference in their entirety.
Sensing Membrane
[0276] Preferably, a sensing membrane 32 is disposed over the
electroactive surfaces of the sensor 34 and includes one or more
domains or layers (FIG. 3B, for example). In general, the sensing
membrane functions to control the flux of a biological fluid there
through and/or to protect sensitive regions of the sensor from
contamination by the biological fluid, for example. Some
conventional electrochemical enzyme-based analyte sensors generally
include a sensing membrane that controls the flux of the analyte
being measured, protects the electrodes from contamination of the
biological fluid, and/or provides an enzyme that catalyzes the
reaction of the analyte with a co-factor, for example. See, e.g.,
U.S. Publication No. 2005-0245799-A1 and U.S. Publication No.
US-2006-0020187-A1.
[0277] The sensing membranes of the preferred embodiments can
include any membrane configuration suitable for use with any
analyte sensor (such as described in more detail above). In
general, the sensing membranes of the preferred embodiments include
one or more domains, all or some of which can be adhered to or
deposited on the analyte sensor as is appreciated by one skilled in
the art. In one embodiment, the sensing membrane generally provides
one or more of the following functions: 1) protection of the
exposed electrode surface from the biological environment, 2)
diffusion resistance (limitation) of the analyte, 3) a catalyst for
enabling an enzymatic reaction, 4) limitation or blocking of
interfering species, and 5) hydrophilicity at the electrochemically
reactive surfaces of the sensor interface, such as described in the
above-referenced co-pending U.S. Patent Applications.
Electrode Domain
[0278] In some embodiments, the membrane system comprises an
optional electrode domain. The electrode domain is provided to
ensure that an electrochemical reaction occurs between the
electroactive surfaces of the working electrode and the reference
electrode, and thus the electrode domain is preferably situated
more proximal to the electroactive surfaces than the enzyme domain.
Preferably, the electrode domain includes a semipermeable coating
that maintains a layer of water at the electrochemically reactive
surfaces of the sensor, for example, a humectant in a binder
material can be employed as an electrode domain; this allows for
the full transport of ions in the aqueous environment. The
electrode domain can also assist in stabilizing the operation of
the sensor by overcoming electrode start-up and drifting problems
caused by inadequate electrolyte. The material that forms the
electrode domain can also protect against pH-mediated damage that
can result from the formation of a large pH gradient due to the
electrochemical activity of the electrodes.
[0279] In one embodiment, the electrode domain includes a flexible,
water-swellable, hydrogel film having a "dry film" thickness of
from about 0.05 micron or less to about 20 microns or more, more
preferably from about 0.05, 0.1, 0.15, 0.2, 0.25, 0.3, 0.35, 0.4,
0.45, 0.5, 1, 1.5, 2, 2.5, 3, or 3.5 to about 4, 5, 6, 7, 8, 9, 10,
11, 12, 13, 14, 15, 16, 17, 18, 19, or 19.5 microns, and more
preferably from about 2, 2.5 or 3 microns to about 3.5, 4, 4.5, or
5 microns. "Dry film" thickness refers to the thickness of a cured
film cast from a coating formulation by standard coating
techniques.
[0280] In certain embodiments, the electrode domain is formed of a
curable mixture of a urethane polymer and a hydrophilic polymer.
Particularly preferred coatings are formed of a polyurethane
polymer having carboxylate functional groups and non-ionic
hydrophilic polyether segments, wherein the polyurethane polymer is
crosslinked with a water soluble carbodiimide (e.g.,
1-ethyl-3-(3-dimethylaminopropyl)carbodiimide (EDC))) in the
presence of polyvinylpyrrolidone and cured at a moderate
temperature of about 50.degree. C.
[0281] Preferably, the electrode domain is deposited by spray or
dip-coating the electroactive surfaces of the sensor. More
preferably, the electrode domain is formed by dip-coating the
electroactive surfaces in an electrode solution and curing the
domain for a time of from about 15 to about 30 minutes at a
temperature of from about 40 to about 55.degree. C. (and can be
accomplished under vacuum (e.g., 20 to 30 mmHg)). In embodiments
wherein dip-coating is used to deposit the electrode domain, a
preferred insertion rate of from about 1 to about 3 inches per
minute, with a preferred dwell time of from about 0.5 to about 2
minutes, and a preferred withdrawal rate of from about 0.25 to
about 2 inches per minute provide a functional coating. However,
values outside of those set forth above can be acceptable or even
desirable in certain embodiments, for example, dependent upon
viscosity and surface tension as is appreciated by one skilled in
the art. In one embodiment, the electroactive surfaces of the
electrode system are dip-coated one time (one layer) and cured at
50.degree. C. under vacuum for 20 minutes.
[0282] Although an independent electrode domain is described
herein, in some embodiments, sufficient hydrophilicity can be
provided in the interference domain and/or enzyme domain (the
domain adjacent to the electroactive surfaces) so as to provide for
the full transport of ions in the aqueous environment (e.g. without
a distinct electrode domain).
Interference Domain
[0283] In some embodiments, an optional interference domain is
provided, which generally includes a polymer domain that restricts
the flow of one or more interferants. In some embodiments, the
interference domain functions as a molecular sieve that allows
analytes and other substances that are to be measured by the
electrodes to pass through, while preventing passage of other
substances, including interferants such as ascorbate and urea (see
U.S. Pat. No. 6,001,067 to Shults). Some known interferants for a
glucose-oxidase based electrochemical sensor include acetaminophen,
ascorbic acid, bilirubin, cholesterol, creatinine, dopamine,
ephedrine, ibuprofen, L-dopa, methyldopa, salicylate, tetracycline,
tolazamide, tolbutamide, triglycerides, and uric acid.
[0284] Several polymer types that can be utilized as a base
material for the interference domain include polyurethanes,
polymers having pendant ionic groups, and polymers having
controlled pore size, for example. In one embodiment, the
interference domain includes a thin, hydrophobic membrane that is
non-swellable and restricts diffusion of low molecular weight
species. The interference domain is permeable to relatively low
molecular weight substances, such as hydrogen peroxide, but
restricts the passage of higher molecular weight substances,
including glucose and ascorbic acid. Other systems and methods for
reducing or eliminating interference species that can be applied to
the membrane system of the preferred embodiments are described in
U.S. Pat. No. 7,074,307, U.S. Publication No. US-2005-0176136-A1,
U.S. Pat. No. 7,081,195 and U.S. Publication No.
US-2005-0143635-A1. In some alternative embodiments, a distinct
interference domain is not included.
[0285] In preferred embodiments, the interference domain is
deposited onto the electrode domain (or directly onto the
electroactive surfaces when a distinct electrode domain is not
included) for a domain thickness of from about 0.05 micron or less
to about 20 microns or more, more preferably from about 0.05, 0.1,
0.15, 0.2, 0.25, 0.3, 0.35, 0.4, 0.45, 0.5, 1, 1.5, 2, 2.5, 3, or
3.5 to about 4, 5, 6, 7, 8, 9, 10, 11, 12, 13, 14, 15, 16, 17, 18,
19, or 19.5 microns, and more preferably from about 2, 2.5 or 3
microns to about 3.5, 4, 4.5, or 5 microns. Thicker membranes can
also be useful, but thinner membranes are generally preferred
because they have a lower impact on the rate of diffusion of
hydrogen peroxide from the enzyme membrane to the electrodes.
Unfortunately, the thin thickness of the interference domains
conventionally used can introduce variability in the membrane
system processing. For example, if too much or too little
interference domain is incorporated within a membrane system, the
performance of the membrane can be adversely affected.
Enzyme Domain
[0286] In preferred embodiments, the membrane system further
includes an enzyme domain (e.g., FIG. 3B, 46) disposed more
distally from the electroactive surfaces than the interference
domain (or electrode domain when a distinct interference is not
included). In some embodiments, the enzyme domain is directly
deposited onto the electroactive surfaces (when neither an
electrode or interference domain is included; e.g., FIG. 3B, 44).
In the preferred embodiments, the enzyme domain provides an enzyme
to catalyze the reaction of the analyte and its co-reactant, as
described in more detail below. Preferably, the enzyme domain
includes glucose oxidase, however other oxidases, for example,
galactose oxidase or uricase oxidase, can also be used.
[0287] For an enzyme-based electrochemical glucose sensor to
perform well, the sensor's response is preferably limited by
neither enzyme activity nor co-reactant concentration. Because
enzymes, including glucose oxidase, are subject to deactivation as
a function of time even in ambient conditions, this behavior is
compensated for in forming the enzyme domain. Preferably, the
enzyme domain is constructed of aqueous dispersions of colloidal
polyurethane polymers including the enzyme. However, in alternative
embodiments the enzyme domain is constructed from an oxygen
enhancing material, for example, silicone or fluorocarbon, in order
to provide a supply of excess oxygen during transient ischemia.
Preferably, the enzyme is immobilized within the domain. See U.S.
Publication No. US-2005-0054909-A1.
[0288] In preferred embodiments, the enzyme domain is deposited
onto the interference domain for a domain thickness of from about
0.05 micron or less to about 20 microns or more, more preferably
from about 0.05, 0.1, 0.15, 0.2, 0.25, 0.3, 0.35, 0.4, 0.45, 0.5,
1, 1.5, 2, 2.5, 3, or 3.5 to about 4, 5, 6, 7, 8, 9, 10, 11, 12,
13, 14, 15, 16, 17, 18, 19, or 19.5 microns, and more preferably
from about 2, 2.5 or 3 microns to about 3.5, 4, 4.5, or 5 microns.
However in some embodiments, the enzyme domain is deposited onto
the electrode domain or directly onto the electroactive surfaces.
Preferably, the enzyme domain 46 is deposited by spray or dip
coating. More preferably, the enzyme domain is formed by
dip-coating the electrode domain into an enzyme domain solution and
curing the domain for from about 15 to about 30 minutes at a
temperature of from about 40 to about 55.degree. C. (and can be
accomplished under vacuum (e.g., 20 to 30 mmHg)). In embodiments
wherein dip-coating is used to deposit the enzyme domain at room
temperature, a preferred insertion rate of from about 1 inch per
minute to about 3 inches per minute, with a preferred dwell time of
from about 0.5 minutes to about 2 minutes, and a preferred
withdrawal rate of from about 0.25 inch per minute to about 2
inches per minute provide a functional coating. However, values
outside of those set forth above can be acceptable or even
desirable in certain embodiments, for example, dependent upon
viscosity and surface tension as is appreciated by one skilled in
the art. In one embodiment, the enzyme domain is formed by dip
coating two times (namely, forming two layers) in a coating
solution and curing at 50.degree. C. under vacuum for 20 minutes.
However, in some embodiments, the enzyme domain can be formed by
dip-coating and/or spray-coating one or more layers at a
predetermined concentration of the coating solution, insertion
rate, dwell time, withdrawal rate, and/or desired thickness.
Resistance Domain
[0289] In preferred embodiments, the membrane system includes a
resistance domain disposed more distal from the electroactive
surfaces than the enzyme domain (e.g., FIG. 3B, 48). Although the
following description is directed to a resistance domain for a
glucose sensor, the resistance domain can be modified for other
analytes and co-reactants as well.
[0290] 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 glucose sensor
employing oxygen as co-reactant is preferably supplied with oxygen
in non-rate-limiting excess in order for the sensor to respond
linearly to changes in glucose concentration, while not responding
to changes in oxygen concentration. 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 for glucose concentrations of up to about 40
mg/dL. However, in a clinical setting, a linear response to glucose
levels is desirable up to at least about 400 mg/dL.
[0291] The resistance domain includes a semi permeable membrane
that controls the flux of oxygen and glucose to the underlying
enzyme domain, 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 exhibits an oxygen to glucose permeability ratio
of from about 50:1 or less to about 400:1 or more, preferably about
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)).
[0292] In alternative embodiments, a lower ratio of
oxygen-to-glucose can be sufficient to provide excess oxygen by
using a high oxygen solubility domain (for example, a silicone or
fluorocarbon-based material or domain) to enhance the
supply/transport of oxygen to the enzyme domain. If more oxygen is
supplied to the enzyme, then more glucose can also be supplied to
the enzyme without creating an oxygen rate-limiting excess. In
alternative embodiments, the resistance domain is formed from a
silicone composition, such as is described in U.S. Publication No.
US-2005-0090607-A1.
[0293] In a preferred embodiment, the resistance domain includes a
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. A suitable hydrophobic polymer
component is a polyurethane, or polyetherurethaneurea. 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 about 4 to about 8 methylene units. Diisocyanates containing
cycloaliphatic moieties can also be useful in the preparation of
the polymer and copolymer components of the membranes of preferred
embodiments. The material that forms the basis of the hydrophobic
matrix of the resistance domain can be any of those known in the
art as appropriate for use as membranes in sensor devices and as
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 can be used to make non-polyurethane type membranes
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.
[0294] In a preferred embodiment, the hydrophilic polymer component
is polyethylene oxide as disclosed in copending U.S. patent
application Ser. No. 11/404,417, filed Apr. 14, 2006 and entitled
"SILICONE BASED MEMBRANES FOR USE IN IMPLANTABLE GLUCOSE SENSORS."
For example, one useful hydrophobic-hydrophilic copolymer component
is a polyurethane polymer that includes about 20% hydrophilic
polyethylene oxide. The polyethylene oxide portions of the
copolymer are 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 affects the water
pick-up and subsequent glucose permeability of the membrane.
[0295] In preferred embodiments, the resistance domain is deposited
onto the enzyme domain to yield a domain thickness of from about
0.05 micron or less to about 20 microns or more, more preferably
from about 0.05, 0.1, 0.15, 0.2, 0.25, 0.3, 0.35, 0.4, 0.45, 0.5,
1, 1.5, 2, 2.5, 3, or 3.5 to about 4, 5, 6, 7, 8, 9, 10, 11, 12,
13, 14, 15, 16, 17, 18, 19, or 19.5 microns, and more preferably
from about 2, 2.5 or 3 microns to about 3.5, 4, 4.5, or 5 microns.
Preferably, the resistance domain is deposited onto the enzyme
domain by spray coating or dip coating. In certain embodiments,
spray coating is the preferred deposition technique. The spraying
process atomizes and mists the solution, and therefore most or all
of the solvent is evaporated prior to the coating material settling
on the underlying domain, thereby minimizing contact of the solvent
with the enzyme. One additional advantage of spray-coating the
resistance domain as described in the preferred embodiments
includes formation of a membrane system that substantially blocks
or resists ascorbate (a known electrochemical interferant in
hydrogen peroxide-measuring glucose sensors). While not wishing to
be bound by theory, it is believed that during the process of
depositing the resistance domain as described in the preferred
embodiments, a structural morphology is formed, characterized in
that ascorbate does not substantially permeate there through.
[0296] In preferred embodiments, the resistance domain is deposited
on the enzyme domain by spray-coating a solution of from about 1
wt. % to about 5 wt. % polymer and from about 95 wt. % to about 99
wt. % solvent. In spraying a solution of resistance domain
material, including a solvent, onto the enzyme domain, it is
desirable to mitigate or substantially reduce any contact with
enzyme of any solvent in the spray solution that can deactivate the
underlying enzyme of the enzyme domain. Tetrahydrofuran (THF) is
one solvent that minimally or negligibly affects the enzyme of the
enzyme domain upon spraying. Other solvents can also be suitable
for use, as is appreciated by one skilled in the art.
[0297] Although a variety of spraying or deposition techniques can
be used, spraying the resistance domain material and rotating the
sensor at least one time by 180.degree. can provide adequate
coverage by the resistance domain. Spraying the resistance domain
material and rotating the sensor at least two times by 120 degrees
provides even greater coverage (one layer of 360.degree. coverage),
thereby ensuring resistivity to glucose, such as is described in
more detail above.
[0298] In preferred embodiments, the resistance domain is
spray-coated and subsequently cured for a time of from about 15 to
about 90 minutes at a temperature of from about 40 to about
60.degree. C. (and can be accomplished under vacuum (e.g., 20 to 30
mmHg)). A cure time of up to about 90 minutes or more can be
advantageous to ensure complete drying of the resistance domain.
While not wishing to be bound by theory, it is believed that
complete drying of the resistance domain aids in stabilizing the
sensitivity of the glucose sensor signal. It reduces drifting of
the signal sensitivity over time, and complete drying is believed
to stabilize performance of the glucose sensor signal in lower
oxygen environments.
[0299] In one embodiment, the resistance domain is formed by
spray-coating at least six layers (namely, rotating the sensor
seventeen times by 120.degree. for at least six layers of
360.degree. coverage) and curing at 50.degree. C. under vacuum for
60 minutes. However, the resistance domain can be formed by
dip-coating or spray-coating any layer or plurality of layers,
depending upon the concentration of the solution, insertion rate,
dwell time, withdrawal rate, and/or the desired thickness of the
resulting film.
[0300] Advantageously, sensors with the membrane system of the
preferred embodiments, including an electrode domain and/or
interference domain, an enzyme domain, and a resistance domain,
provide stable signal response to increasing glucose levels of from
about 40 to about 400 mg/dL, and sustained function (at least 90%
signal strength) even at low oxygen levels (for example, at about
0.6 mg/L O.sub.2). While not wishing to be bound by theory, it is
believed that the resistance domain provides sufficient
resistivity, or the enzyme domain provides sufficient enzyme, such
that oxygen limitations are seen at a much lower concentration of
oxygen as compared to prior art sensors.
[0301] In preferred embodiments, a sensor signal with a current in
the picoAmp range is preferred, which is described in more detail
elsewhere herein. However, the ability to produce a signal with a
current in the picoAmp range can be dependent upon a combination of
factors, including the electronic circuitry design (e.g., A/D
converter, bit resolution, and the like), the membrane system
(e.g., permeability of the analyte through the resistance domain,
enzyme concentration, and/or electrolyte availability to the
electrochemical reaction at the electrodes), and the exposed
surface area of the working electrode. For example, the resistance
domain can be designed to be more or less restrictive to the
analyte depending upon to the design of the electronic circuitry,
membrane system, and/or exposed electroactive surface area of the
working electrode.
[0302] Accordingly, in preferred embodiments, the membrane system
is designed with a sensitivity of from about 1 pA/mg/dL to about
100 pA/mg/dL, preferably from about 5 pA/mg/dL to 25 pA/mg/dL, and
more preferably from about 4 to about 7 pA/mg/dL. While not wishing
to be bound by any particular theory, it is believed that membrane
systems designed with a sensitivity in the preferred ranges permit
measurement of the analyte signal in low analyte and/or low oxygen
situations. Namely, conventional analyte sensors have shown reduced
measurement accuracy in low analyte ranges due to lower
availability of the analyte to the sensor and/or have shown
increased signal noise in high analyte ranges due to insufficient
oxygen necessary to react with the amount of analyte being
measured. While not wishing to be bound by theory, it is believed
that the membrane systems of the preferred embodiments, in
combination with the electronic circuitry design and exposed
electrochemical reactive surface area design, support measurement
of the analyte in the picoAmp range, which enables an improved
level of resolution and accuracy in both low and high analyte
ranges not seen in the prior art.
[0303] Although sensors of some embodiments described herein
include an optional interference domain in order to block or reduce
one or more interferants, sensors with the membrane system of the
preferred embodiments, including an electrode domain, an enzyme
domain, and a resistance domain, have been shown to inhibit
ascorbate without an additional interference domain. Namely, the
membrane system of the preferred embodiments, including an
electrode domain, an enzyme domain, and a resistance domain, has
been shown to be substantially non-responsive to ascorbate in
physiologically acceptable ranges. While not wishing to be bound by
theory, it is believed that the process of depositing the
resistance domain by spray coating, as described herein, results in
a structural morphology that is substantially resistance resistant
to ascorbate.
Interference-Free Membrane Systems
[0304] In general, it is believed that appropriate solvents and/or
deposition methods can be chosen for one or more of the domains of
the membrane system that form one or more transitional domains such
that interferants do not substantially permeate there through.
Thus, sensors can be built without distinct or deposited
interference domains, which are non-responsive to interferants.
While not wishing to be bound by theory, it is believed that a
simplified multilayer membrane system, more robust multilayer
manufacturing process, and reduced variability caused by the
thickness and associated oxygen and glucose sensitivity of the
deposited micron-thin interference domain can be provided.
Additionally, the optional polymer-based interference domain, which
usually inhibits hydrogen peroxide diffusion, is eliminated,
thereby enhancing the amount of hydrogen peroxide that passes
through the membrane system.
Oxygen Conduit
[0305] As described above, certain sensors depend upon an enzyme
within the membrane system through which the host's bodily fluid
passes and in which the analyte (for example, glucose) within the
bodily fluid reacts in the presence of a co-reactant (for example,
oxygen) to generate a product. The product is then measured using
electrochemical methods, and thus the output of an electrode system
functions as a measure of the analyte. For example, when the sensor
is a glucose oxidase based glucose sensor, the species measured at
the working electrode is H.sub.2O.sub.2. An enzyme, 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
[0306] Because for each glucose molecule reacted there is a
proportional change in the product, H.sub.2O.sub.2, one can monitor
the change in H.sub.2O.sub.2 to determine glucose concentration.
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 and
other reducible species at a counter electrode, for example. 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 Wiley and Sons, New York))
[0307] In vivo, glucose concentration is generally about one
hundred times or more that of the oxygen concentration.
Consequently, oxygen is a limiting reactant in the electrochemical
reaction, and when insufficient oxygen is provided to the sensor,
the sensor is unable to accurately measure glucose concentration.
Thus, depressed sensor function or inaccuracy is believed to be a
result of problems in availability of oxygen to the enzyme and/or
electroactive surface(s).
[0308] Accordingly, in an alternative embodiment, an oxygen conduit
(for example, a high oxygen solubility domain formed from silicone
or fluorochemicals) is provided that extends from the ex vivo
portion of the sensor to the in vivo portion of the sensor to
increase oxygen availability to the enzyme. The oxygen conduit can
be formed as a part of the coating (insulating) material or can be
a separate conduit associated with the assembly of wires that forms
the sensor.
[0309] FIG. 3B is a cross-sectional view through the sensor of FIG.
3A on line B-B, showing an exposed electroactive surface of at
least a working electrode 38 surrounded by a sensing membrane. In
general, the sensing membranes of the preferred embodiments include
a plurality of domains or layers, for example, an interference
domain 44, an enzyme domain 46, and a resistance domain 48, and can
include additional domains, such as an electrode domain, a cell
impermeable domain, and/or an oxygen domain (not shown), such as
described in more detail in the above-cited co-pending U.S. Patent
Applications. However, it is understood that a sensing membrane
modified for other sensors, for example, by including fewer or
additional domains is within the scope of the preferred
embodiments. In some embodiments, one or more domains of the
sensing membranes 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 fluoride (PVDF), polybutylene terephthalate (PBT),
polymethylmethacrylate (PMMA), polyether ether ketone (PEEK),
polyurethanes, cellulosic polymers, poly(ethylene oxide),
poly(propylene oxide) and copolymers and blends thereof,
polysulfones and block copolymers thereof including, for example,
di-block, tri-block, alternating, random and graft copolymers.
Co-pending U.S. patent application Ser. No. 10/838,912, which is
incorporated herein by reference in its entirety, describes
biointerface and sensing membrane configurations and materials that
can be applied to the preferred embodiments.
[0310] The sensing membrane can be deposited on the electroactive
surfaces of the electrode material using known thin or thick film
techniques (for example, spraying, electro-depositing, dipping, or
the like). The sensing membrane that surrounds the working
electrode does not have to be the same structure as the sensing
membrane that surrounds a reference electrode, etc. For example,
the enzyme domain deposited over the working electrode does not
necessarily need to be deposited over the reference and/or counter
electrodes.
[0311] In the illustrated embodiment, the sensor is an enzyme-based
electrochemical sensor, wherein the working electrode 38 measures
the hydrogen peroxide produced by the enzyme catalyzed reaction of
glucose being detected and creates a measurable electronic current
(for example, detection of glucose utilizing glucose oxidase
produces hydrogen 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),
such as described in more detail above and as is appreciated by one
skilled in the art. Preferably, one or more potentiostat is
employed to monitor the electrochemical reaction at the
electroactive surface of the working electrode(s). The potentiostat
applies a constant potential to the working electrode and its
associated reference electrode to determine the current produced at
the working electrode. 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. The output
signal is typically a raw data stream that is used to provide a
useful value of the measured analyte concentration in a host to the
host or doctor, for example.
[0312] Some alternative analyte sensors that can benefit from the
systems and methods of the preferred embodiments include U.S. Pat.
No. 5,711,861 to Ward et al., U.S. Pat. No. 6,642,015 to Vachon et
al., U.S. Pat. No. 6,654,625 to Say et al., U.S. Pat. No. 6,565,509
to Say et al., U.S. Pat. No. 6,514,718 to Heller, U.S. Pat. No.
6,465,066 to Essenpreis et al., U.S. Pat. No. 6,214,185 to
Offenbacher et al., U.S. Pat. No. 5,310,469 to Cunningham et al.,
and U.S. Pat. No. 5,683,562 to Shaffer et al., U.S. Pat. No.
6,579,690 to Bonnecaze et al., U.S. Pat. No. 6,484,046 to Say et
al., U.S. Pat. No. 6,512,939 to Colvin et al., U.S. Pat. No.
6,424,847 to Mastrototaro et al., U.S. Pat. No. 6,424,847 to
Mastrototaro et al, for example. All of the above patents are not
inclusive of all applicable analyte sensors; in general, it should
be understood that the disclosed embodiments are applicable to a
variety of analyte sensor configurations.
Exemplary Sensor Configurations
[0313] FIG. 4A is a side schematic view of a transcutaneous analyte
sensor 50 in one embodiment. The sensor 50 includes a mounting unit
52 adapted for mounting on the skin of a host, a small diameter
sensor 34 (as defined herein) adapted for transdermal insertion
through the skin of a host, and an electrical connection configured
to provide secure electrical contact between the sensor and the
electronics preferably housed within the mounting unit 52. In
general, the mounting unit 52 is designed to maintain the integrity
of the sensor in the host so as to reduce or eliminate translation
of motion between the mounting unit, the host, and/or the sensor.
See U.S. Publication No. US-2006-0020187-A1. Preferably, a
biointerface membrane is formed onto the sensing mechanism 34 as
described in more detail below.
[0314] FIG. 4B is a side schematic view of a transcutaneous analyte
sensor 54 in an alternative embodiment. The sensor 54 includes a
mounting unit 52 wherein the sensing mechanism 34 comprises a small
structure as defined herein and is tethered to the mounting unit 52
via a cable 56 (alternatively, a wireless connection can be
utilized). The mounting unit is adapted for mounting on the skin of
a host and is operably connected via a tether, or the like, to a
small structured sensor 34 adapted for transdermal insertion
through the skin of a host and measurement of the analyte therein;
see, for example, U.S. Pat. No. 6,558,330 to Causey III, et al.,
which is incorporated herein by reference in its entirety. In the
preferred embodiments, a biointerface membrane is formed onto the
sensing mechanism 34 as described in more detail below.
[0315] The short-term sensor of the preferred embodiments can be
inserted into a variety of locations on the host's body, such as
the abdomen, the thigh, the upper arm, and the neck or behind the
ear. Although the preferred embodiments illustrate insertion
through the abdominal region, the systems and methods described
herein are limited neither to the abdominal nor to the subcutaneous
insertions. One skilled in the art appreciates that these systems
and methods can be implemented and/or modified for other insertion
sites and can be dependent upon the type, configuration, and
dimensions of the analyte sensor.
[0316] In one embodiment, an analyte-sensing device adapted for
transcutaneous short-term insertion into the host is provided. For
example, the device includes a sensor, for measuring the analyte in
the host, a porous, biocompatible matrix covering at least a
portion of the sensor, and an applicator, for inserting the sensor
through the host's skin. In some embodiments, the sensor has
architecture with at least one dimension less than about 1 mm.
Examples of such a structure are shown in FIGS. 4A and 4B, as
described elsewhere herein. However, one skilled in the art will
recognize that alternative configurations are possible and can be
desirable, depending upon factors such as intended location of
insertion, for example. The sensor is inserted through the host's
skin and into the underlying tissue, such as soft tissue or fatty
tissue.
[0317] After insertion, fluid moves into the spacer, e.g., a
biocompatible matrix or membrane, creating a fluid-filled pocket
therein. This process can occur immediately or can take place over
a period of time, such as several minutes or hours post insertion.
A signal from the sensor is then detected, such as by the sensor
electronics unit located in the mounting unit on the surface of the
host's skin. In general, the sensor can be used continuously for a
short period of days, such as 1 to 14 days. After use, the sensor
is simply removed from the host's skin. In preferred embodiments,
the host can repeat the insertion and detection steps as many times
as desired. In some implementations, the sensor can be removed
after about 3 days, and then another sensor inserted, and so on.
Similarly in other implementations, the sensor is removed after
about 3, 5, 7, 10 or 14 days, followed by insertion of a new
sensor, and so on.
[0318] Some examples of transcutaneous analyte sensors are
described in co-pending U.S. patent application Ser. No.
11/360,250, filed Feb. 22, 2006 and entitled "ANALYTE SENSOR." In
general, transcutaneous analyte sensors comprise the sensor and a
mounting unit with electronics associated therewith.
[0319] In general, the mounting unit includes a base adapted for
mounting on the skin of a host, a sensor adapted for transdermal
insertion through the skin of a host, and one or more contacts
configured to provide secure electrical contact between the sensor
and the sensor electronics. The mounting unit is designed to
maintain the integrity of the sensor in the host so as to reduce or
eliminate translation of motion between the mounting unit, the
host, and/or the sensor.
[0320] The base can be formed from a variety of hard or soft
materials, and preferably comprises a low profile for minimizing
protrusion of the device from the host during use. In some
embodiments, the base is formed at least partially from a flexible
material, which is believed to provide numerous advantages over
conventional transcutaneous sensors, which, unfortunately, can
suffer from motion-related artifacts associated with the host's
movement when the host is using the device. For example, when a
transcutaneous analyte sensor is inserted into the host, various
movements of the sensor (for example, relative movement between the
in vivo portion and the ex vivo portion, movement of the skin,
and/or movement within the host (dermis or subcutaneous)) create
stresses on the device and can produce noise in the sensor signal.
It is believed that even small movements of the skin can translate
to discomfort and/or motion-related artifact, which can be reduced
or obviated by a flexible or articulated base. Thus, by providing
flexibility and/or articulation of the device against the host's
skin, better conformity of the sensor system to the regular use and
movements of the host can be achieved. Flexibility or articulation
is believed to increase adhesion (with the use of an adhesive pad)
of the mounting unit onto the skin, thereby decreasing
motion-related artifact that can otherwise translate from the
host's movements and reduced sensor performance.
[0321] In certain embodiments, the mounting unit is provided with
an adhesive pad, preferably disposed on the mounting unit's back
surface and preferably including a releasable backing layer. Thus,
removing the backing layer and pressing the base portion of the
mounting unit onto the host's skin adheres the mounting unit to the
host's skin. Additionally or alternatively, an adhesive pad can be
placed over some or all of the sensor system after sensor insertion
is complete to ensure adhesion, and optionally to ensure an
airtight seal or watertight seal around the wound exit-site (or
sensor insertion site). Appropriate adhesive pads can be chosen and
designed to stretch, elongate, conform to, and/or aerate the region
(e.g., host's skin).
[0322] In preferred embodiments, the adhesive pad is formed from
spun-laced, open- or closed-cell foam, and/or non-woven fibers, and
includes an adhesive disposed thereon, however a variety of
adhesive pads appropriate for adhesion to the host's skin can be
used, as is appreciated by one skilled in the art of medical
adhesive pads. In some embodiments, a double-sided adhesive pad is
used to adhere the mounting unit to the host's skin. In other
embodiments, the adhesive pad includes a foam layer, for example, a
layer wherein the foam is disposed between the adhesive pad's side
edges and acts as a shock absorber.
[0323] In some embodiments, the surface area of the adhesive pad is
greater than the surface area of the mounting unit's back surface.
Alternatively, the adhesive pad can be sized with substantially the
same surface area as the back surface of the base portion.
Preferably, the adhesive pad has a surface area on the side to be
mounted on the host's skin that is greater than about 1, 1.25, 1.5,
1.75, 2, 2.25, or 2.5, times the surface area of the back surface
of the mounting unit base. Such a greater surface area can increase
adhesion between the mounting unit and the host's skin, minimize
movement between the mounting unit and the host's skin, and/or
protect the wound exit-site (sensor insertion site) from
environmental and/or biological contamination. In some alternative
embodiments, however, the adhesive pad can be smaller in surface
area than the back surface assuming a sufficient adhesion can be
accomplished.
[0324] In some embodiments, the adhesive pad is substantially the
same shape as the back surface of the base, although other shapes
can also be advantageously employed, for example, butterfly-shaped,
round, square, or rectangular. The adhesive pad backing can be
designed for two-step release, for example, a primary release
wherein only a portion of the adhesive pad is initially exposed to
allow adjustable positioning of the device, and a secondary release
wherein the remaining adhesive pad is later exposed to firmly and
securely adhere the device to the host's skin once appropriately
positioned. The adhesive pad is preferably waterproof. Preferably,
a stretch-release adhesive pad is provided on the back surface of
the base portion to enable easy release from the host's skin at the
end of the useable life of the sensor.
[0325] In some circumstances, it has been found that a conventional
bond between the adhesive pad and the mounting unit can not be
sufficient, for example, due to humidity that can cause release of
the adhesive pad from the mounting unit. Accordingly, in some
embodiments, the adhesive pad can be bonded using a bonding agent
activated by or accelerated by an ultraviolet, acoustic, radio
frequency, or humidity cure. In some embodiments, a eutectic bond
of first and second composite materials can form a strong adhesion.
In some embodiments, the surface of the mounting unit can be
pretreated utilizing ozone, plasma, chemicals, or the like, in
order to enhance the bondability of the surface.
[0326] A bioactive agent is preferably applied locally at the
insertion site prior to or during sensor insertion. Suitable
bioactive agents include those which are known to discourage or
prevent bacterial growth and infection, for example,
anti-inflammatory agents, antimicrobials, antibiotics, or the like.
It is believed that the diffusion or presence of a bioactive agent
can aid in prevention or elimination of bacteria adjacent to the
exit-site. Additionally or alternatively, the bioactive agent can
be integral with or coated on the adhesive pad, or no bioactive
agent at all is employed.
[0327] In some embodiments, an applicator is provided for inserting
the sensor through the host's skin at the appropriate insertion
angle with the aid of a needle, and for subsequent removal of the
needle using a continuous push-pull action. Preferably, the
applicator comprises an applicator body that guides the applicator
and includes an applicator body base configured to mate with the
mounting unit during insertion of the sensor into the host. The
mate between the applicator body base and the mounting unit can use
any known mating configuration, for example, a snap-fit, a
press-fit, an interference-fit, or the like, to discourage
separation during use. One or more release latches enable release
of the applicator body base, for example, when the applicator body
base is snap fit into the mounting unit.
[0328] The sensor electronics includes hardware, firmware, and/or
software that enable measurement of levels of the analyte via the
sensor. For example, the sensor electronics can comprise a
potentiostat, a power source for providing power to the sensor,
other components useful for signal processing, and preferably an RF
module for transmitting data from the sensor electronics to a
receiver. Electronics can be affixed to a printed circuit board
(PCB), or the like, and can take a variety of forms. For example,
the electronics can take the form of an integrated circuit (IC),
such as an Application-Specific Integrated Circuit (ASIC), a
microcontroller, or a processor. Preferably, sensor electronics
comprise systems and methods for processing sensor analyte data.
Examples of systems and methods for processing sensor analyte data
are described in more detail below and in U.S. Publication No.
US-2005-0027463-A1.
[0329] In this embodiment, after insertion of the sensor using the
applicator, and subsequent release of the applicator from the
mounting unit, the sensor electronics are configured to releasably
mate with the mounting unit. In one embodiment, the electronics are
configured with programming, for example initialization,
calibration reset, failure testing, or the like, each time it is
initially inserted into the mounting unit and/or each time it
initially communicates with the sensor.
Sensor Electronics
[0330] The following description of electronics associated with the
sensor is applicable to a variety of continuous analyte sensors,
such as non-invasive, minimally invasive, and/or invasive (e.g.,
transcutaneous and wholly implantable) sensors. For example, the
sensor electronics and data processing as well as the receiver
electronics and data processing described below can be incorporated
into the wholly implantable glucose sensor disclosed in U.S.
Publication No. US-2005-0245799-A1 and U.S. Publication No.
US-2006-0015020-A1.
[0331] In one embodiment, a potentiostat, which is operably
connected to an electrode system (such as described above) provides
a voltage to the electrodes, which biases the sensor to enable
measurement of a current signal indicative of the analyte
concentration in the host (also referred to as the analog portion).
In some embodiments, the potentiostat includes a resistor that
translates the current into voltage. In some alternative
embodiments, a current to frequency converter is provided that is
configured to continuously integrate the measured current, for
example, using a charge counting device. An A/D converter digitizes
the analog signal into a digital signal, also referred to as
"counts" for processing. Accordingly, the resulting raw data stream
in counts, also referred to as raw sensor data, is directly related
to the current measured by the potentiostat.
[0332] A processor module includes the central control unit that
controls the processing of the sensor electronics. In some
embodiments, the processor module includes a microprocessor,
however a computer system other than a microprocessor can be used
to process data as described herein, for example an ASIC can be
used for some or all of the sensor's central processing. The
processor typically 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 is
described in U.S. Publication No. US-2005-0043598-A1. The processor
additionally can be used for the system's cache memory, for example
for temporarily storing recent sensor data. In some embodiments,
the processor module comprises memory storage components such as
ROM, RAM, dynamic-RAM, static-RAM, non-static RAM, EEPROM,
rewritable ROMs, flash memory, or the like.
[0333] In some embodiments, the processor module comprises a
digital filter, for example, an IIR or FIR filter, configured to
smooth the raw data stream from the A/D converter. Generally,
digital filters are programmed to filter data sampled at a
predetermined time interval (also referred to as a sample rate.) In
some embodiments, wherein the potentiostat is configured to measure
the analyte at discrete time intervals, these time intervals
determine the sample rate of the digital filter. In some
alternative embodiments, wherein the potentiostat is configured to
continuously measure the analyte, for example, using a
current-to-frequency converter as described above, the processor
module can be programmed to request a digital value from the A/D
converter at a predetermined time interval, also referred to as the
acquisition time. In these alternative embodiments, the values
obtained by the processor are advantageously averaged over the
acquisition time due the continuity of the current measurement.
Accordingly, the acquisition time determines the sample rate of the
digital filter. In preferred embodiments, the processor module is
configured with a programmable acquisition time, namely, the
predetermined time interval for requesting the digital value from
the A/D converter is programmable by a user within the digital
circuitry of the processor module. An acquisition time of from
about 2 seconds to about 512 seconds is preferred; however any
acquisition time can be programmed into the processor module. A
programmable acquisition time is advantageous in optimizing noise
filtration, time lag, and processing/battery power.
[0334] Preferably, the processor module is configured to build the
data packet for transmission to an outside source, for example, an
RF transmission to a receiver as described in more detail below.
Generally, the data packet comprises a plurality of bits that can
include a sensor ID code, raw data, filtered data, and/or error
detection or correction. The processor module can be configured to
transmit any combination of raw and/or filtered data.
[0335] In some embodiments, the processor module further comprises
a transmitter portion that determines the transmission interval of
the sensor data to a receiver, or the like. In some embodiments,
the transmitter portion, which determines the interval of
transmission, is configured to be programmable. In one such
embodiment, a coefficient can be chosen (e.g., a number of from
about 1 to about 100, or more), wherein the coefficient is
multiplied by the acquisition time (or sampling rate), such as
described above, to define the transmission interval of the data
packet. Thus, in some embodiments, the transmission interval is
programmable between about 2 seconds and about 850 minutes, more
preferably between about 30 second and 5 minutes; however, any
transmission interval can be programmable or programmed into the
processor module. However, a variety of alternative systems and
methods for providing a programmable transmission interval can also
be employed. By providing a programmable transmission interval,
data transmission can be customized to meet a variety of design
criteria (e.g., reduced battery consumption, timeliness of
reporting sensor values, etc.)
[0336] Conventional glucose sensors measure current in the nanoAmp
range. In contrast to conventional glucose sensors, the preferred
embodiments are configured to measure the current flow in the
picoAmp range, and in some embodiments, femtoAmps. Namely, for
every unit (mg/dL) of glucose measured, at least one picoAmp of
current is measured. Preferably, the analog portion of the A/D
converter is configured to continuously measure the current flowing
at the working electrode and to convert the current measurement to
digital values representative of the current. In one embodiment,
the current flow is measured by a charge counting device (e.g., a
capacitor). Thus, a signal is provided, whereby a high sensitivity
maximizes the signal received by a minimal amount of measured
hydrogen peroxide (e.g., minimal glucose requirements without
sacrificing accuracy even in low glucose ranges), reducing the
sensitivity to oxygen limitations in vivo (e.g., in
oxygen-dependent glucose sensors).
[0337] A battery is operably connected to the sensor electronics
and provides the power for the sensor. 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, and/or a plurality of batteries can be used to power
the system. The sensor can be transcutaneously powered via an
inductive coupling, for example. In some embodiments, a quartz
crystal is operably connected to the processor and maintains system
time for the computer system as a whole, for example for the
programmable acquisition time within the processor module.
[0338] Optional temperature probe can be provided, wherein the
temperature probe is located on the electronics assembly or the
glucose sensor itself. The temperature probe can be used to measure
ambient temperature in the vicinity of the glucose sensor. This
temperature measurement can be used to add temperature compensation
to the calculated glucose value.
[0339] An RF module is operably connected to the processor and
transmits the sensor data from the sensor to a receiver within a
wireless transmission via antenna. In some embodiments, a second
quartz crystal provides the time base for the RF carrier frequency
used for data transmissions from the RF transceiver. In some
alternative embodiments, however, other mechanisms, such as
optical, infrared radiation (IR), ultrasonic, or the like, can be
used to transmit and/or receive data.
[0340] In 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 from about 3 to about 24 months, or more) with maximum RF
transmittance from the in vivo environment to the ex vivo
environment for wholly implantable sensors (for example, a distance
of from about one to ten meters or more). Preferably, a high
frequency carrier signal of from about 402 MHz to about 433 MHz is
employed in order to maintain lower power requirements.
Additionally, in wholly implantable devices, 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, the preferred glucose sensor can sustain sensor
function for 3 months, 6 months, 12 months, or 24 months or
more.
[0341] In some embodiments, output signal (from the sensor
electronics) is sent to a receiver (e.g., a computer or other
communication station). The 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 a doctor, for example. In
some embodiments, the raw data stream can be continuously or
periodically algorithmically smoothed or otherwise modified to
diminish outlying points that do not accurately represent the
analyte concentration, for example due to signal noise or other
signal artifacts, such as described in U.S. Pat. No. 6,931,327.
[0342] When a sensor is first implanted into host tissue, the
sensor and receiver are initialized. This can be referred to as
start-up mode, and involves optionally resetting the sensor data
and calibrating the sensor 32. In selected embodiments, mating the
electronics unit 16 to the mounting unit triggers a start-up mode.
In other embodiments, the start-up mode is triggered by the
receiver.
Receiver
[0343] In some embodiments, the sensor electronics are wirelessly
connected to a receiver via one- or two-way RF transmissions or the
like. However, a wired connection is also contemplated. The
receiver provides much of the processing and display of the sensor
data, and can be selectively worn and/or removed at the host's
convenience. Thus, the sensor system can be discreetly worn, and
the receiver, which provides much of the processing and display of
the sensor data, can be selectively worn and/or removed at the
host's convenience. Particularly, the receiver includes programming
for retrospectively and/or prospectively initiating a calibration,
converting sensor data, updating the calibration, evaluating
received reference and sensor data, and evaluating the calibration
for the analyte sensor, such as described in more detail with
reference to co-pending U.S. Publication No.
US-2005-0027463-A1.
[0344] FIG. 4C is a side schematic view of a wholly implantable
analyte sensor 58 in one embodiment. The sensor includes a sensor
body 60 suitable for subcutaneous implantation and includes a small
structured sensor 34 as defined herein. Published U.S. Patent
Application No. 2004/0199059 to Brauker et al. describe systems and
methods suitable for the sensor body 60, and is incorporated herein
by reference in its entirety. In the preferred embodiments, a
biointerface membrane 68 is formed onto the sensing mechanism 34 as
described in more detail elsewhere herein. The sensor body 60
includes sensor electronics and preferably communicates with a
receiver as described in more detail, above.
[0345] FIG. 4D is a side schematic view of a wholly implantable
analyte sensor 62 in an alternative embodiment. The sensor 62
includes a sensor body 60 and a small structured sensor 34 as
defined herein. The sensor body 60 includes sensor electronics and
preferably communicates with a receiver as described in more
detail, above.
[0346] In preferred embodiments, a biointerface membrane 68 is
formed onto the sensing mechanism 34 as described in more detail
elsewhere herein. Preferably, a matrix or framework 64 surrounds
the sensing mechanism 34 for protecting the sensor from some
foreign body processes, for example, by causing tissue to compress
against or around the framework 64 rather than the sensing
mechanism 34.
[0347] In general, the optional protective framework 64 is formed
from a two-dimensional or three-dimensional flexible, semi-rigid,
or rigid matrix (e.g., mesh), and which includes spaces or pores
through which the analyte can pass. In some embodiments, the
framework is incorporated as a part of the biointerface membrane,
however a separate framework can be provided. While not wishing to
be bound by theory, it is believed that the framework 64 protects
the small structured sensing mechanism from mechanical forces
created in vivo.
[0348] FIG. 4E is a side schematic view of a wholly implantable
analyte sensor 66 in another alternative embodiment. The sensor 66
includes a sensor body 60 and a small structured sensor 34, as
defined herein, with a biointerface membrane 68 such as described
in more detail elsewhere herein. Preferably, a framework 64
protects the sensing mechanism 34 such as described in more detail
above. The sensor body 60 includes sensor electronics and
preferably communicates with a receiver as described in more
detail, above.
[0349] In certain embodiments, the sensing device, which is adapted
to be wholly implanted into the host, such as in the soft tissue
beneath the skin, is implanted subcutaneously, such as in the
abdomen of the host, for example. One skilled in the art
appreciates a variety of suitable implantation sites available due
to the sensor's small size. In some embodiments, the sensor
architecture is less than about 0.5 mm in at least one dimension,
for example a wire-based sensor with a diameter of less than about
0.5 mm. In another exemplary embodiment, for example, the sensor
can be 0.5 mm thick, 3 mm in length and 2 cm in width, such as
possibly a narrow substrate, needle, wire, rod, sheet or pocket. In
another exemplary embodiment, a plurality of about 1 mm wide wires
about 5 mm in length could be connected at their first ends,
producing a forked sensor structure. In still another embodiment, a
1 mm wide sensor could be coiled, to produce a planar, spiraled
sensor structure. Although a few examples are cited above, numerous
other useful embodiments are contemplated by the present invention,
as is appreciated by one skilled in the art.
[0350] Post implantation, a period of time is allowed for tissue
ingrowth within the biointerface. The length of time required for
tissue ingrowth varies from host to host, such as about a week to
about 3 weeks, although other time periods are also possible. Once
a mature bed of vascularized tissue has grown into the
biointerface, a signal can be detected from the sensor, as
described elsewhere herein and in U.S. Publication No.
2005-0245799-A1. Long-term sensors can remain implanted and produce
glucose signal information from months to years, as described in
the above-cited patent application.
[0351] In certain embodiments, the device is configured such that
the sensing unit is separated from the electronics unit by a tether
or cable, or a similar structure, similar to that illustrated in
FIG. 4B. One skilled in the art will recognize that a variety of
known and useful means can be used to tether the sensor to the
electronics. While not wishing to be bound by theory, it is
believed that the FBR to the electronics unit alone can be greater
than the FBR to the sensing unit alone, due to the electronics
unit's greater mass, for example. Accordingly, separation of the
sensing and electronics units effectively reduces the FBR to the
sensing unit and results in improved device function. As described
elsewhere herein, the architecture and/or composition of the
sensing unit (e.g., inclusion of a biointerface with certain
bioactive agents) can be implemented to further reduce the foreign
body response to the tethered sensing unit.
[0352] In another embodiment, an analyte sensor is designed with
separate electronics and sensing units, wherein the sensing unit is
inductively coupled to the electronics unit. In this embodiment,
the electronics unit provides power to the sensing unit and/or
enables communication of data therebetween. FIGS. 3F and 3G
illustrate exemplary systems that employ inductive coupling between
an electronics unit 52 and a sensing unit 58.
[0353] FIG. 4F is a side view of one embodiment of an implanted
sensor inductively coupled to an electronics unit within a
functionally useful distance on the host's skin. FIG. 4F
illustrates a sensing unit 58, including a sensing mechanism 34,
biointerface 68 and small electronics chip 216 implanted below the
host's skin 212, within the host's tissue 210. In this example, the
majority of the electronics associated with the sensor are housed
in an electronics unit 52 (also referred to as a mounting unit)
located within suitably close proximity on the host's skin. The
electronics unit 52 is inductively coupled to the small electronics
chip 216 on the sensing unit 58 and thereby transmits power to the
sensor and/or collects data, for example. The small electronics
chip 216 coupled to the sensing unit 58 provides the necessary
electronics to provide a bias potential to the sensor, measure the
signal output, and/or other necessary requirements to allow the
sensing mechanism 58 to function (e.g., chip 216 can include an
ASIC (application specific integrated circuit), antenna, and other
necessary components appreciated by one skilled in the art).
[0354] In yet another embodiment, the implanted sensor additionally
includes a capacitor to provide necessary power for device
function. A portable scanner (e.g., wand-like device) is used to
collect data stored on the circuit and/or to recharge the
device.
[0355] In general, inductive coupling, as described herein, enables
power to be transmitted to the sensor for continuous power,
recharging, and the like. Additionally, inductive coupling utilizes
appropriately spaced and oriented antennas (e.g., coils) on the
sensing unit and the electronics unit so as to efficiently
transmit/receive power (e.g., current) and/or data communication
therebetween. One or more coils in each of the sensing and
electronics unit can provide the necessary power induction and/or
data transmission.
[0356] In this embodiment, the sensing mechanism can be, for
example, a wire-based sensor as described in more detail with
reference to FIGS. 4A and 4B and as described in published U.S.
Patent Application US2006-0020187, or a planar substrate-based
sensor such as described in U.S. Pat. No. 6,175,752 to Say et al.
and U.S. Pat. No. 5,779,665 to Mastrototaro et al., all of which
are incorporated herein by reference in their entirety. The
biointerface 68 can be any suitable biointerface as described in
more detail elsewhere herein, for example, a layer of porous
biointerface membrane material, a mesh cage and the like. In one
exemplary embodiment, the biointerface 68 is a single- or
multi-layer sheet (e.g., pocket) of porous membrane material, such
as ePTFE, in which the sensing mechanism 34 is incorporated.
[0357] FIG. 4G is a side view of on embodiment of an implanted
sensor inductively coupled to an electronics unit implanted in the
host's tissue at a functionally useful distance. FIG. 4G
illustrates a sensor unit 58 and an electronics unit 52 similar to
that described with reference to FIG. 4F, above, however both are
implanted beneath the host's skin in a suitably close
proximity.
[0358] In general, it is believed that when the electronics unit
52, which carries the majority of the mass of the implantable
device, is separate from the sensing unit 58, a lesser foreign body
response will occur surrounding the sensing unit (e.g., as compared
to a device of greater mass, for example, a device including
certain electronics and/or power supply). Thus, the configuration
of the sensing unit, including a biointerface, can be optimized to
minimize and/or modify the host's tissue response, for example with
minimal mass as described in more detail elsewhere.
Biointerface
[0359] In some embodiments, the sensor includes a porous material
disposed over some portion thereof, which modifies the host's
tissue response to the sensor. In some embodiments, the porous
material surrounding the sensor advantageously enhances and extends
sensor performance and lifetime in the short-term by slowing or
reducing cellular migration to the sensor and associated
degradation that would otherwise be caused by cellular invasion if
the sensor were directly exposed to the in vivo environment.
Alternatively, the porous material can provide stabilization of the
sensor via tissue ingrowth into the porous material in the
long-term. Suitable porous materials include silicone,
polytetrafluoroethylene, expanded polytetrafluoroethylene,
polyethylene-co-tetrafluoroethylene, polyolefin, polyester,
polycarbonate, biostable polytetrafluoroethylene, homopolymers,
copolymers, terpolymers of polyurethanes, polypropylene (PP),
polyvinylchloride (PVC), polyvinylidene fluoride (PVDF), polyvinyl
alcohol (PVA), polybutylene terephthalate (PBT),
polymethylmethacrylate (PMMA), polyether ether ketone (PEEK),
polyamides, polyurethanes, cellulosic polymers, poly(ethylene
oxide), poly(propylene oxide) and copolymers and blends thereof,
polysulfones and block copolymers thereof including, for example,
di-block, tri-block, alternating, random and graft copolymers, as
well as metals, ceramics, cellulose, hydrogel polymers, poly
(2-hydroxyethyl methacrylate, pHEMA), hydroxyethyl methacrylate,
(HEMA), polyacrylonitrile-polyvinyl chloride (PAN-PVC), high
density polyethylene, acrylic copolymers, nylon, polyvinyl
difluoride, polyanhydrides, poly(l-lysine), poly (L-lactic acid),
hydroxyethylmethacrylate, hydroxyapeptite, alumina, zirconia,
carbon fiber, aluminum, calcium phosphate, titanium, titanium
alloy, nintinol, stainless steel, and CoCr alloy, or the like, such
as are described in U.S. Publication No. US-2005-0031689-A1 and
U.S. Publication No. US-2005-0112169-A1.
[0360] In some embodiments, the porous material surrounding the
sensor provides unique advantages in the short-term (e.g., one to
14 days) that can be used to enhance and extend sensor performance
and lifetime. However, such materials can also provide advantages
in the long-term too (e.g., greater than 14 days). Particularly,
the in vivo portion of the sensor (the portion of the sensor that
is implanted into the host's tissue) is encased (partially or
fully) in a porous material. The porous material can be wrapped
around the sensor (for example, by wrapping the porous material
around the sensor or by inserting the sensor into a section of
porous material sized to receive the sensor). Alternately, the
porous material can be deposited on the sensor (for example, by
electrospinning of a polymer directly thereon). In yet other
alternative embodiments, the sensor is inserted into a selected
section of porous biomaterial. Other methods for surrounding the in
vivo portion of the sensor with a porous material can also be used
as is appreciated by one skilled in the art.
[0361] The porous material surrounding the sensor advantageously
slows or reduces cellular migration to the sensor and associated
degradation that would otherwise be caused by cellular invasion if
the sensor were directly exposed to the in vivo environment.
Namely, the porous material provides a barrier that makes the
migration of cells towards the sensor more tortuous and therefore
slower (providing short-term advantages). It is believed that this
reduces or slows the sensitivity loss normally observed in a
short-term sensor over time.
[0362] In an embodiment wherein the porous material is a high
oxygen solubility material, such as porous silicone, the high
oxygen solubility porous material surrounds some of or the entire
in vivo portion of the sensor. In some embodiments, a lower ratio
of oxygen-to-glucose can be sufficient to provide excess oxygen by
using a high oxygen soluble domain (for example, a silicone- or
fluorocarbon-based material) to enhance the supply/transport of
oxygen to the enzyme membrane and/or electroactive surfaces. It is
believed that some signal noise normally seen by a conventional
sensor can be attributed to an oxygen deficit. Silicone has high
oxygen permeability, thus promoting oxygen transport to the enzyme
layer. By enhancing the oxygen supply through the use of a silicone
composition, for example, glucose concentration can be less of a
limiting factor. In other words, if more oxygen is supplied to the
enzyme and/or electroactive surfaces, then more glucose can also be
supplied to the enzyme without creating an oxygen rate-limiting
excess. While not being bound by any particular theory, it is
believed that silicone materials provide enhanced bio-stability
when compared to other polymeric materials such as
polyurethane.
[0363] In certain aspects, modifying a small structured sensor with
a biointerface structure, material, matrix, and/or membrane that
creates a space appropriate for filling with fluid in vivo can
enhance sensor performance. In some embodiments, the small
structured sensor includes a porous biointerface material, which
allows fluid from the surrounding tissues to form a fluid-filled
pocket around at least a portion of the sensor. It is believed that
the fluid-filled pocket provides a sufficient source of
analyte-containing fluid for accurate sensor measurement in the
short-term. Additionally or alternatively, inclusion of bioactive
agents can modify the host's tissue response, for example to reduce
or eliminate tissue ingrowth or other cellular responses into the
biointerface.
[0364] In some aspects, modifying a small structured sensor with a
structure, material, and/or membrane/matrix that allows tissue
ingrowth without barrier cell formation can enhance sensor
performance. For example, a vascularized bed of tissue for
long-term analyte sensor measurement. In some embodiments, a porous
biointerface membrane, including a plurality of interconnected
cavities and a solid portion, covering at least the sensing portion
of a small structured sensor allows vascularized tissue ingrowth
therein. Vascularized tissue ingrowth provides a sufficient source
of analyte-containing tissue in the long-term. Additionally or
alternatively, inclusion of bioactive agents can modify the host's
tissue response, for example to reduce or eliminate barrier cell
layer formation within the membrane.
[0365] When used herein, the terms "membrane" and "matrix" are
meant to be interchangeable. In these embodiments first domain is
provided that includes an architecture, including cavity size,
configuration, and/or overall thickness, that modifies the host's
tissue response, for example, by creating a fluid pocket,
encouraging vascularized tissue ingrowth, disrupting downward
tissue contracture, resisting fibrous tissue growth adjacent to the
device, and/or discouraging barrier cell formation. The
biointerface preferably covers at least the sensing mechanism of
the sensor and can be of any shape or size, including uniform,
asymmetrically, or axi-symmetrically covering or surrounding a
sensing mechanism or sensor.
[0366] A second domain is optionally provided that is impermeable
to cells and/or cell processes. A bioactive agent is optionally
provided that is incorporated into the at least one of the first
domain, the second domain, the sensing membrane, or other part of
the implantable device, wherein the bioactive agent is configured
to modify a host tissue response.
[0367] FIG. 5A is a cross-sectional schematic view of a
biointerface membrane 70 in vivo in one exemplary embodiment,
wherein the membrane comprises a first domain 72 and an optional
second domain 74. In the short-term, the architecture of the
biointerface membrane provides a space between the sensor and the
host's tissue that allows a fluid filled pocket to form for
transport of fluid therein. In the long-term, the architecture of
the membrane provides a robust, implantable membrane that
facilitates the transport of analytes through vascularized tissue
ingrowth without the formation of a barrier cell layer.
[0368] The first domain 72 comprises a solid portion 76 and a
plurality of interconnected three-dimensional cavities 78 formed
therein. In this embodiment, the cavities 78 have sufficient size
and structure to allow invasive cells, such as fibroblasts 75, a
fibrous matrix 77, and blood vessels 79 to enter into the apertures
80 that define the entryway into each cavity 78, and to pass
through the interconnected cavities toward the interface 82 between
the first and second domains. The cavities comprise an architecture
that encourages the ingrowth of vascular tissue in vivo, as
indicated by the blood vessels 79 formed throughout the cavities.
Because of the vascularization within the cavities, solutes 73 (for
example, oxygen, glucose and other analytes) pass through the first
domain with relative ease, and/or the diffusion distance (namely,
distance that the glucose diffuses) is reduced.
Architecture of the First Domain
[0369] In some embodiments, the first domain of the biointerface
membrane includes an architecture that supports tissue ingrowth,
disrupts contractile forces typically found in a foreign body
response, encourages vascularity within the membrane, and disrupts
the formation of a barrier cell layer. In some alternative
embodiments, the first domain of the biointerface membrane includes
an architecture that creates a fluid-filled space surrounding an
implanted device, which allows the passage of the analyte, but
protects sensitive portions of the device from substantial fibrous
tissue ingrowth and associated forces.
[0370] In general, the first domain, also referred to as the cell
disruptive domain, comprises an open-celled configuration
comprising interconnected cavities and solid portions. The
distribution of the solid portion and cavities of the first domain
preferably includes a substantially co-continuous solid domain and
includes more than one cavity in three dimensions substantially
throughout the entirety of the first domain. However, some
short-term embodiments cannot require co-continuity of the
cavities. Generally, cells can enter into the cavities; however,
they cannot travel through or wholly exist within the solid
portions. The cavities permit most substances to pass through,
including, for example, cells and molecules. One example of a
suitable material is expanded polytetrafluoraethylene (ePTFE).
[0371] Reference is now made to FIG. 5B, which is an illustration
of the membrane of FIG. 5A, showing contractile forces 81 caused by
the fibrous tissue in the long-term (e.g., after about 3 weeks),
for example, from the fibroblasts and fibrous matrix, of the FBR.
Specifically, the architecture of the first domain, including the
cavity interconnectivity and multiple-cavity depth, (namely, 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.
[0372] The architecture of the first domain of the biointerface
membrane, including the interconnected cavities and solid portion,
is 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 81 that contract around the solid portions 76 (for example,
throughout the interconnected cavities 78) away from the device.
That is, the architecture of the solid portions 76 and cavities 78
of the first domain cause contractile forces 81 to disperse away
from the interface between the first domain 72 and second domain
74. Without the organized contracture of fibrous tissue toward the
tissue-device interface 82 typically found in a FBC (FIG. 5B),
macrophages and foreign body giant cells do not form a substantial
monolayer of cohesive cells (namely, a barrier cell layer) and
therefore the transport of molecules across the second domain
and/or membrane is not blocked, as indicated by free transport of
analyte 73 through the first and second domains in FIGS. 5A and
5B.
[0373] Various methods are suitable for use in manufacturing the
first domain in order to create an architecture with preferred
dimensions and overall structure. The first domain can be
manufactured by forming particles, for example, sugar granules,
salt granules, and other natural or synthetic uniform or
non-uniform particles, in a mold, wherein the particles have shapes
and sizes substantially corresponding to the desired cavity
dimensions, such as described in more detail below. In some
methods, the particles are made to coalesce to provide the desired
interconnectivity between the cavities. The desired material for
the solid portion can be introduced into the mold using methods
common in the art of polymer processing, for example, injecting,
pressing, vacuuming, vapor depositing, pouring, and the like. After
the solid portion material is cured or solidified, the coalesced
particles are then dissolved, melted, etched, or otherwise removed,
leaving interconnecting cavities within the solid portion. In such
embodiments, sieving can be used to determine the dimensions of the
particles, which substantially correspond to the dimensions of
resulting cavities. In sieving, also referred to as screening, the
particles are added to the sieve and then shaken to produce overs
and unders. The overs are the particles that remain on the screen
and the unders are the particles that pass through the screen.
Other methods and apparatus known in the art are also suitable for
use in determining particle size, for example, air classifiers,
which apply opposing air flows and centrifugal forces to separate
particles having sizes down to 2 .mu.m, can be used to determine
particle size when particles are smaller than 100 .mu.m.
[0374] In one embodiment, the cavity size of the cavities 78 of the
first domain is substantially defined by the particle size(s) used
in creating the cavities. In some embodiments, the particles used
to form the cavities can be substantially spherical, thus the
dimensions below describe a diameter of the particle and/or a
diameter of the cavity. In some alternative embodiments, the
particles used to form the cavities can be non-spherical (for
example, rectangular, square, diamond, or other geometric or
non-geometric shapes), thus the dimensions below describe one
dimension (for example, shortest, average, or longest) of the
particle and/or cavity.
[0375] In some embodiments, a variety of different particle sizes
can be used in the manufacture of the first domain. In some
embodiments, the dimensions of the particles can be somewhat
smaller or larger than the dimensions of the resulting cavities,
due to dissolution or other precipitation that can occur during the
manufacturing process.
[0376] Although one method of manufacturing porous domains is
described above, a variety of methods known to one of ordinary
skill in the art can be employed to create the structures of
preferred embodiments, see section entitled, "Formation of the
Biointerface onto the Sensor," below. For example, molds can be
used in the place of the particles described above, such as coral,
self-assembly beads, etched or broken silicon pieces, glass frit
pieces, and the like. The dimensions of the mold can define the
cavity sizes, which can be determined by measuring the cavities of
a model final product, and/or by other measuring techniques known
in the art, for example, by a bubble point test. In U.S. Pat. No.
3,929,971, Roy discloses a method of making a synthetic membrane
having a porous microstructure by converting calcium carbonate
coral materials to hydroxyapatite while at the same time retaining
the unique microstructure of the coral material.
[0377] Other methods of forming a three-dimensional first domain
can be used, for example holographic lithography,
stereolithography, and the like, wherein cavity sizes are defined
and precisely formed by the lithographic or other such process to
form a lattice of unit cells, as described in U.S. Publication No.
US-2005-0251083-A1, and in U.S. Pat. No. 6,520,997, which discloses
a photolithographic process for creating a porous membrane.
[0378] The first domain 72 can be defined using alternative
methods. In an alternative preferred embodiment, fibrous non-woven
or woven materials, or other such materials, such as electrospun,
felted, velvet, scattered, or aggregate materials, are manufactured
by forming the solid portions without particularly defining the
cavities therebetween. Accordingly, in these alternative
embodiments, structural elements that provide the three-dimensional
conformation can include fibers, strands, globules, cones, and/or
rods of amorphous or uniform geometry. These elements are
hereinafter referred to as "strands." The solid portion of the
first domain can include a plurality of strands, which generally
define apertures formed by a frame of the interconnected strands.
The apertures of the material form a framework of interconnected
cavities. Formed in this manner, the first domain is defined by a
cavity size of about 0.6 to about 1 mm in at least one
dimension.
[0379] Referring to the dimensions and architecture of the first
domain 72, the porous biointerface membranes can be loosely
categorized into at least two groups: those having a
micro-architecture and those having a macro-architecture.
[0380] FIGS. 5A and 5B illustrate one preferred embodiment wherein
the biointerface membrane includes a macro-architecture as defined
herein. In general, the cavity size of a macro-architecture
provides a configuration and overall thickness that encourages
vascular tissue ingrowth and disrupts tissue contracture that is
believed to cause barrier cell formation in the long-term in vivo
(as indicated by the blood vessels 79 formed throughout the
cavities), while providing a long-term, robust structure. Referring
to the macro-architecture, a substantial number of the cavities 78,
defined using any of the methods described above, are greater than
or equal to about 20 .mu.m in one dimension. In some other
embodiments, a substantial number of the cavities are greater than
or equal to about 30, 40, 50, 60, 70, 80, 90, 100, 120, 180, 160,
180, 200, 280, 280, 320, 360, 400, 500, 600, 700 .mu.m, and
preferably less than about 1 mm in one dimension.
[0381] The biointerface membrane can also be formed with a
micro-architecture as defined herein. Generally, at least some of
the cavities of a micro-architecture have a sufficient size and
structure to allow inflammatory cells to partially or completely
enter into the cavities. However, in contrast to the
macro-architecture, the micro-architecture does not allow extensive
ingrowth of vascular and connective tissues within the cavities.
Therefore, in some embodiments, the micro-architecture of preferred
embodiments is defined by the actual size of the cavity, wherein
the cavities are formed from a mold, for example, such as described
in more detail above. However, in the context of the
micro-architecture it is preferable that the majority of the mold
dimensions, whether particles, beads, crystals, coral,
self-assembly beads, etched or broken silicon pieces, glass frit
pieces, or other mold elements that form cavities, are less than
about 20 .mu.m in at least one dimension.
[0382] In some alternative embodiments, wherein the biointerface
membrane is formed from a substantially fibrous material, the
micro-architecture is defined by a strand size of less than about 6
.mu.m in all but the longest dimension, and a sufficient number of
cavities are provided of a size and structure to allow inflammatory
cells, for example, macrophages, to completely enter through the
apertures that define the cavities, without extensive ingrowth of
vascular and connective tissues.
[0383] In certain embodiments, the micro-architecture is
characterized, or defined, by standard pore size tests, such as the
bubble point test. The micro-architecture is selected with a
nominal pore size of from about 0.6 .mu.m to about 20 .mu.m. In
some embodiments, the nominal pore size from about 1, 2, 3, 4, 5,
6, 7, 8, or 9 .mu.m to about 10, 11, 12, 13, 14, 15, 16, 17, 18, or
19 .mu.m. It has been found that a porous polymer membrane having
an average nominal pore size of about 0.6 to about 20 .mu.m
functions satisfactorily in creating a vascular bed within the
micro-architecture at the device-tissue interface. The term
"nominal pore size" in the context of the micro-architecture in
certain embodiments is derived from methods of analysis common to
membrane, such as the ability of the membrane to filter particles
of a particular size, or the resistance of the membrane to the flow
of fluids. Because of the amorphous, random, and irregular nature
of most of these commercially available membranes, the "nominal
pore size" designation cannot actually indicate the size or shape
of the apertures and cavities, which in reality have a high degree
of variability. Accordingly, as used herein with reference to the
micro-architecture, the term "nominal pore size" is a
manufacturer's convention used to identify a particular membrane of
a particular commercial source which has a certain bubble point; as
used herein, the term "pore" does not describe the size of the
cavities of the material in the preferred embodiments. The bubble
point measurement is described in Pharmaceutical Technology, May
1983, pp. 76 to 82.
[0384] The optimum dimensions, architecture (for example,
micro-architecture or macro-architecture), and overall structural
integrity of the membrane can be adjusted according to the
parameters of the device that it supports. For example, if the
membrane is employed with a glucose-measuring device, the
mechanical requirements of the membrane can be greater for devices
having greater overall weight and surface area when compared to
those that are relatively smaller.
[0385] In some embodiments, improved vascular tissue ingrowth in
the long-term is observed when the first domain has a thickness
that accommodates a depth of at least two cavities throughout a
substantial portion of the thickness. Improved vascularization
results at least in part from multi-layered interconnectivity of
the cavities, such as in the preferred embodiments, as compared to
a surface topography such as seen in the prior art, for example,
wherein the first domain has a depth of only one cavity throughout
a substantial portion thereof. The multi-layered interconnectivity
of the cavities enables vascularized tissue to grow into various
layers of cavities in a manner that provides mechanical anchoring
of the device with the surrounding tissue. Such anchoring resists
movement that can occur in vivo, which results in reduced sheer
stress and scar tissue formation. The optimum depth or number of
cavities can vary depending upon the parameters of the device that
it supports. For example, if the membrane is employed with a
glucose-measuring device, the anchoring that is required of the
membrane is greater for devices having greater overall weight and
surface area as compared to those that are relatively smaller.
[0386] The thickness of the first domain can be optimized for
decreased time-to-vascularize in vivo, that is, vascular tissue
ingrowth can occur somewhat faster with a membrane that has a thin
first domain as compared to a membrane that has a relatively
thicker first domain. Decreased time-to-vascularize results in
faster stabilization and functionality of the biointerface in vivo.
For example, in a subcutaneous implantable glucose device,
consistent and increasing functionality of the device is at least
in part a function of consistent and stable glucose transport
across the biointerface membrane, which is at least in part a
function of the vascularization thereof. Thus, quicker start-up
time and/or shortened time lag (as when, for example, the diffusion
path of the glucose through the membrane is reduced) can be
achieved by decreasing the thickness of the first domain.
[0387] The thickness of the first domain is typically from about 20
.mu.m to about 2000 .mu.m, preferably from about 50, 60, 70, 80,
90, or 100 .mu.m to about 800, 900, 1000, 1100, 1200, 1300, 1400,
1500, 1600, 1700, 1800, or 1900 .mu.m, and most preferably from
about 150, 200, 250, 300, 350, or 400 .mu.m to about 450, 500, 550,
600, 650, 700, or 750 .mu.m. However, in some alternative
embodiments a thinner or thicker cell disruptive domain (first
domain) can be desired.
[0388] The solid portion preferably includes one or more materials
such as silicone, polytetrafluoroethylene, expanded
polytetrafluoroethylene, polyethylene-co-tetrafluoroethylene,
polyolefin, polyester, polycarbonate, biostable
polytetrafluoroethylene, homopolymers, copolymers, terpolymers of
polyurethanes, polypropylene (PP), polyvinylchloride (PVC),
polyvinylidene fluoride (PVDF), polyvinyl alcohol (PVA),
polybutylene terephthalate (PBT), polymethylmethacrylate (PMMA),
polyether ether ketone (PEEK), polyamides, polyurethanes,
cellulosic polymers, polysulfones and block copolymers thereof
including, for example, di-block, tri-block, alternating, random
and graft copolymers. In some embodiments, the material selected
for the first domain is an elastomeric material, for example,
silicone, which is able to absorb stresses that can occur in vivo,
such that sheer and other environmental forces are significantly
minimized at the second domain. The solid portion can comprises a
silicone composition with a hydrophile such as Polyethylene Glycol
(PEG) covalently incorporated or grafted therein, such as described
in U.S. Publication No. US-2005-0090607-A1 or as disclosed in
copending U.S. patent application Ser. No. 11/404,417, filed Apr.
14, 2006 and entitled "SILICONE BASED MEMBRANES FOR USE IN
IMPLANTABLE GLUCOSE SENSORS."
[0389] One preferred material that can be used to form the solid
portion of the biointerface matrix is a material that allows the
passage of the analyte (e.g., glucose) there through. For example,
the biointerface matrix can be formed from a silicone
polymer/hydrophobic-hydrophilic polymer blend. In one embodiment,
The hydrophobic-hydrophilic polymer for use in the blend can be any
suitable hydrophobic-hydrophilic polymer, including but not limited
to components such as polyvinylpyrrolidone (PVP), polyhydroxyethyl
methacrylate, polyvinylalcohol, polyacrylic acid, polyethers such
as polyethylene glycol or polypropylene oxide, and copolymers
thereof, including, for example, di-block, tri-block, alternating,
random, comb, star, dendritic, and graft copolymers (block
copolymers are discussed in U.S. Pat. Nos. 4,803,243 and 4,686,044,
which are incorporated herein by reference). In one embodiment, the
hydrophobic-hydrophilic polymer is a copolymer of poly(ethylene
oxide) (PEO) and poly(propylene oxide) (PPO). Suitable such
polymers include, but are not limited to, PEO-PPO diblock
copolymers, PPO-PEO-PPO triblock copolymers, PEO-PPO-PEO triblock
copolymers, alternating block copolymers of PEO-PPO, random
copolymers of ethylene oxide and propylene oxide, and blends
thereof. In some embodiments, the copolymers can be optionally
substituted with hydroxy substituents. Commercially available
examples of PEO and PPO copolymers include the PLURONIC.RTM. brand
of polymers available from BASF.RTM.. In one embodiment,
PLURONIC.RTM. F-127 is used. Other PLURONIC.RTM. polymers include
PPO-PEO-PPO triblock copolymers (e.g., PLURONIC.RTM. R products).
Other suitable commercial polymers include, but are not limited to,
SYNPERONICS.RTM. products available from UNIQEMA.RTM..
[0390] The silicone polymer for use in the
silicone/hydrophobic-hydrophilic polymer blend can be any suitable
silicone polymer. In some embodiments, the silicone polymer is a
liquid silicone rubber that can be vulcanized using a metal- (e.g.,
platinum), peroxide-, heat-, ultraviolet-, or other
radiation-catalyzed process. In some embodiments, the silicone
polymer is a dimethyl- and methylhydrogen-siloxane copolymer. In
some embodiments, the copolymer has vinyl substituents. In some
embodiments, commercially available silicone polymers can be used.
For example, commercially available silicone polymer precursor
compositions can be used to prepare the blends, such as described
below. In one embodiment, MED-4840 available from NUSIL.RTM.
Technology LLC is used as a precursor to the silicone polymer used
in the blend. MED-4840 consists of a 2-part silicone elastomer
precursor including vinyl-functionalized dimethyl- and
methylhydrogen-siloxane copolymers, amorphous silica, a platinum
catalyst, a crosslinker, and an inhibitor. The two components can
be mixed together and heated to initiate vulcanization, thereby
forming an elastomeric solid material. Other suitable silicone
polymer precursor systems include, but are not limited to, MED-2174
peroxide-cured liquid silicone rubber available from NUSIL.RTM.
Technology LLC, SILASTIC.RTM. MDX4-4210 platinum-cured biomedical
grade elastomer available from DOW CORNING.RTM., and Implant Grade
Liquid Silicone Polymer (durometers 10-50) available from Applied
Silicone Corporation.
[0391] Silicone polymer/hydrophobic-hydrophilic polymer blends are
described in more detail in U.S. patent application Ser. No.
11/404,417, entitled "SILICONE BASED MEMBRANES FOR USE IN
IMPLANTABLE GLUCOSE SENSORS," filed on Apr. 14, 2006.
[0392] Additionally, elastomeric materials with a memory of the
original configuration can withstand greater stresses without
affecting the configuration, and thus the function, of the
device.
[0393] In some embodiments, the first domain can include a
macro-architecture and a micro-architecture located within at least
a portion of the macro-architecture, such as is described in U.S.
Publication No. US-2005-0251083-A1. For example, the
macro-architecture includes a porous structure with interconnected
cavities such as described with reference to the solid portion of
the first domain, wherein at least some portion of the cavities of
the first domain are filled with the micro-architecture that
includes a fibrous or other fine structured material that aids in
preventing formation of a barrier cell layer, for example in
pockets in the bottom of the cavities of the macro-architecture
adjacent to the implantable device.
[0394] In certain embodiments, other non-resorbable implant
materials can be used in forming the first domain, including but
not limited to, metals, ceramics, cellulose,
polyacrylonitrile-polyvinyl chloride (PAN-PVC), high density
polyethylene, acrylic copolymers, nylon, polyvinyl difluoride,
polyanhydrides, poly(l-lysine), hydroxyethylmethacrylate, alumina,
zirconia, carbon fiber, aluminum, titanium, titanium alloy,
nintinol, stainless steel, and CoCr alloy.
Architecture of the Second Domain
[0395] FIGS. 5A and 5B, illustrate the optional second domain of
the membrane. The second domain is impermeable to cells or cell
processes, and is composed of a biostable material. In one
exemplary embodiment, the second domain is comprised of
polyurethane and a hydrophilic polymer, such as is described in
U.S. Pat. No. 6,862,465 to Shults et al., which is incorporated
herein by reference in its entirety. Alternatively, the outermost
layer of the sensing membrane 32 can function as a cell impermeable
domain and therefore a second domain cannot be a discrete component
of the biointerface membrane.
[0396] In general, the materials preferred for the second domain
prevent or hinder cell entry or contact with device elements
underlying the membrane and prevent or hinder the adherence of
cells, thereby further discouraging formation of a barrier cell
layer. Additionally, because of the resistance of the materials to
barrier cell layer formation, membranes prepared therefrom are
robust long-term in vivo.
[0397] The thickness of the cell impermeable biomaterial of the
second domain (also referred to as a cell impermeable domain) is
typically about 1 .mu.m or more, preferably from about 1, 5, 10,
15, 20, 25, 30, 35, 40, 45, or 50, 55, 60, 65, 70, 75, 80, 85, 90,
95, 100, 110, 120, 130, 140, 150, 160, 170, 180, 190, or 200 .mu.m
to about 500, 600, 700, 800, 900, or 1000 .mu.m. In some
embodiments, thicker or thinner cell impermeable domains can be
desired. Alternatively, the function of the cell impermeable domain
is accomplished by the implantable device, or a portion of the
implantable device, which can or cannot include a distinct domain
or layer.
[0398] The characteristics of the cell impermeable membrane prevent
or hinder cells from entering the membrane, but permit or
facilitate transport of the analyte of interest or a substance
indicative of the concentration or presence of the analyte.
Additionally the second domain, similar to the first domain, is
preferably constructed of a biodurable material (for example, a
material durable for a period of several years in vivo) that is
impermeable to host cells, for example, macrophages, such as
described above.
[0399] In embodiments wherein the biointerface membrane is employed
in an implantable glucose-measuring device, the biointerface
membrane is permeable to oxygen and glucose or a substance
indicative of the concentration of glucose. In embodiments wherein
the membrane is employed in a drug delivery device or other device
for delivering a substance to the body, the cell impermeable
membrane is permeable to the drug or other substance dispensed from
the device. In embodiments wherein the membrane is employed for
cell transplantation, the membrane is semi-permeable, for example,
impermeable to immune cells and soluble factors responsible for
rejecting transplanted tissue, but permeable to the ingress of
glucose and oxygen for the purpose of sustaining the transplanted
tissue; additionally, the second domain is permeable to the egress
of the gene product of interest (for example, insulin).
[0400] The cell disruptive (first) domain and the cell impermeable
(second) domain can be secured to each other by any suitable method
as is known in the art. For example, the cell impermeable domain
can simply be layered or cast upon the porous cell disruptive
domain so as to form a mechanical attachment. Alternatively,
chemical and/or mechanical attachment methods can be suitable for
use. Chemical attachment methods can include adhesives, glues,
lamination, and/or wherein a thermal bond is formed through the
application of heat and pressure, and the like. Suitable adhesives
are those capable of forming a bond between the materials that make
up both the barrier cell disruptive domain and the cell impermeable
domain, and include liquid and/or film applied adhesives. An
appropriate material can be designed that can be used for preparing
both domains such that the composite is prepared in one step,
thereby forming a unitary structure. For example, when the cell
disruptive domain and the cell impermeable domain comprise
silicone, the materials can be designed so that they can be
covalently cured to one another. However in some embodiments
wherein the second domain comprises a part of the implantable
device, it can be attached to or simply lie adjacent to the first
domain.
[0401] In some embodiments wherein an adhesive is employed, the
adhesive can comprise a biocompatible material. However, in some
embodiments adhesives not generally considered to have a high
degree of biocompatibility can also be employed. Adhesives with
varying degrees of biocompatibility suitable for use include
acrylates, for example, cyanoacrylates, epoxies, methacrylates,
polyurethanes, and other polymers, resins, RTV silicone, and
crosslinking agents as are known in the art. In some embodiments, a
layer of woven or non-woven material (such as ePTFE) is cured to
the first domain after which the material is bonded to the second
domain, which allows a good adhesive interface between the first
and second domains using a biomaterial known to respond well at the
tissue-device interface, for example.
Bioactive Agents
[0402] In some alternative embodiment, the biointerface membranes
include a bioactive agent, which is incorporated into at least one
of the first and second domains 72, 74 of the biointerface
membrane, or which is incorporated into the device (e.g., sensing
membrane 32) and adapted to diffuse through the first and/or second
domains, in order to modify the tissue response of the host to the
membrane. The architectures of the first and second domains have
been shown to create a fluid pocket, support vascularized tissue
ingrowth, to interfere with and resist barrier cell layer
formation, and to facilitate the transport of analytes across the
membrane. However, the bioactive agent can further enhance
formation of a fluid pocket, alter or enhance vascularized tissue
ingrowth, resistance to barrier cell layer formation, and thereby
facilitate the passage of analytes 73 across the device-tissue
interface 82.
[0403] In embodiments wherein the biointerface includes a bioactive
agent, the bioactive agent is incorporated into at least one of the
first and second domains of the biointerface membrane, or into the
device and adapted to diffuse through the first and/or second
domains, in order to modify the tissue response of the host to the
membrane. In general, the architectures of the first and second
domains support vascularized tissue growth in or around the
biointerface membrane, interfere with and resist barrier cell layer
formation, and/or allow the transport of analytes across the
membrane. However, certain outside influences, for example, faulty
surgical techniques, acute or chronic movement of the implant, or
other surgery-, host-, and/or implantation site-related conditions,
can create acute and/or chronic inflammation at the implant site.
When this occurs, the biointerface membrane architecture alone
cannot be sufficient to overcome the acute and/or chronic
inflammation. Alternatively, the membrane architecture can benefit
from additional mechanisms that aid in reducing this acute and/or
chronic inflammation that can produce a barrier cell layer and/or a
fibrotic capsule surrounding the implant, resulting in compromised
solute transport through the membrane.
[0404] In general, the inflammatory response to biomaterial
implants can be divided into two phases. The first phase consists
of mobilization of mast cells and then infiltration of
predominantly polymorphonuclear (PMN) cells. This phase is termed
the acute inflammatory phase. Over the course of days to weeks,
chronic cell types that comprise the second phase of inflammation
replace the PMNs. Macrophage and lymphocyte cells predominate
during this phase. While not wishing to be bound by any particular
theory, it is believed that short-term stimulation of
vascularization, or short-term inhibition of scar formation or
barrier cell layer formation, provides protection from scar tissue
formation, thereby providing a stable platform for sustained
maintenance of the altered foreign body response, for example.
[0405] Accordingly, bioactive intervention can modify the foreign
body response in the early weeks of foreign body capsule formation
and alter the long-term behavior of the foreign body capsule.
Additionally, it is believed that in some circumstances the
biointerface membranes of the preferred embodiments can benefit
from bioactive intervention to overcome sensitivity of the membrane
to implant procedure, motion of the implant, or other factors,
which are known to otherwise cause inflammation, scar formation,
and hinder device function in vivo.
[0406] In general, bioactive agents that are believed to modify
tissue response include anti-inflammatory agents, anti-infective
agents, anesthetics, inflammatory agents, growth factors,
angiogenic (growth) factors, adjuvants, immunosuppressive agents,
antiplatelet agents, anticoagulants, ACE inhibitors, cytotoxic
agents, anti-barrier cell compounds, vascularization compounds,
anti-sense molecules, and the like. In some embodiments, preferred
bioactive agents include S1P (Sphingosine-1-phosphate),
Monobutyrin, Cyclosporin A, Anti-thrombospondin-2, Rapamycin (and
its derivatives), and Dexamethasone. However, other bioactive
agents, biological materials (for example, proteins), or even
non-bioactive substances can incorporated into the membranes of
preferred embodiments.
[0407] Bioactive agents suitable for use in the preferred
embodiments are loosely organized into two groups: anti-barrier
cell agents and vascularization agents. These designations reflect
functions that are believed to provide short-term solute transport
through the biointerface membrane, and additionally extend the life
of a healthy vascular bed and hence solute transport through the
biointerface membrane long-term in vivo. However, not all bioactive
agents can be clearly categorized into one or other of the above
groups; rather, bioactive agents generally comprise one or more
varying mechanisms for modifying tissue response and can be
generally categorized into one or both of the above-cited
categories.
Anti-Barrier Cell Agents
[0408] Generally, anti-barrier cell agents include compounds
exhibiting affects on macrophages and foreign body giant cells
(FBGCs). It is believed that anti-barrier cell agents prevent
closure of the barrier to solute transport presented by macrophages
and FBGCs at the device-tissue interface during FBC maturation.
[0409] Anti-barrier cell agents generally include mechanisms that
inhibit foreign body giant cells and/or occlusive cell layers. For
example, Super Oxide Dismutase (SOD) Mimetic, which utilizes a
manganese catalytic center within a porphyrin like molecule to
mimic native SOD and effectively remove superoxide for long
periods, thereby inhibiting FBGC formation at the surfaces of
biomaterials in vivo, is incorporated into a biointerface membrane
of a preferred embodiment.
[0410] Anti-barrier cell agents can include anti-inflammatory
and/or immunosuppressive mechanisms that affect early FBC
formation. Cyclosporine, which stimulates very high levels of
neovascularization around biomaterials, can be incorporated into a
biointerface membrane of a preferred embodiment (see U.S. Pat. No.
5,569,462 to Martinson et al.). Alternatively, Dexamethasone, which
abates the intensity of the FBC response at the tissue-device
interface, can be incorporated into a biointerface membrane of a
preferred embodiment. Alternatively, Rapamycin, which is a potent
specific inhibitor of some macrophage inflammatory functions, can
be incorporated into a biointerface membrane of a preferred
embodiment.
[0411] Other suitable medicaments, pharmaceutical compositions,
therapeutic agents, or other desirable substances can be
incorporated into the membranes of preferred embodiments,
including, but not limited to, anti-inflammatory agents,
anti-infective agents, necrosing agents, and anesthetics.
[0412] Generally, anti-inflammatory agents reduce acute and/or
chronic inflammation adjacent to the implant, in order to decrease
the formation of a FBC capsule to reduce or prevent barrier cell
layer formation. Suitable anti-inflammatory agents include but are
not limited to, for example, nonsteroidal anti-inflammatory drugs
(NSAIDs) such as acetometaphen, aminosalicylic acid, aspirin,
celecoxib, choline magnesium trisalicylate, diclofenac potassium,
diclofenac sodium, diflunisal, etodolac, fenoprofen, flurbiprofen,
ibuprofen, indomethacin, interleukin (IL)-10, IL-6 mutein,
anti-IL-6 iNOS inhibitors (for example, L-NAME or L-NMDA),
Interferon, ketoprofen, ketorolac, leflunomide, melenamic acid,
mycophenolic acid, mizoribine, nabumetone, naproxen, naproxen
sodium, oxaprozin, piroxicam, rofecoxib, salsalate, sulindac, and
tolmetin; and corticosteroids such as cortisone, hydrocortisone,
methylprednisolone, prednisone, prednisolone, betamethesone,
beclomethasone dipropionate, budesonide, dexamethasone sodium
phosphate, flunisolide, fluticasone propionate, paclitaxel,
tacrolimus, tranilast, triamcinolone acetonide, betamethasone,
fluocinolone, fluocinonide, betamethasone dipropionate,
betamethasone valerate, desonide, desoximetasone, fluocinolone,
triamcinolone, triamcinolone acetonide, clobetasol propionate, and
dexamethasone.
[0413] Generally, immunosuppressive and/or immunomodulatory agents
interfere directly with several key mechanisms necessary for
involvement of different cellular elements in the inflammatory
response. Suitable immunosuppressive and/or immunomodulatory agents
include anti-proliferative, cell-cycle inhibitors, (for example,
paclitaxol (e.g., Sirolimus), cytochalasin D, infiximab), taxol,
actinomycin, mitomycin, thospromote VEGF, estradiols, NO donors,
leptin, QP-2, tacrolimus, tranilast, actinomycin, everolimus,
methothrexate, mycophenolic acid, angiopeptin, vincristing,
mitomycine, statins, C MYC antisense, sirolimus (and analogs),
RestenASE, 2-chloro-deoxyadenosine, PCNA Ribozyme, batimstat,
prolyl hydroxylase inhibitors, PPAR.gamma. ligands (for example
troglitazone, rosiglitazone, pioglitazone), halofuginone,
C-proteinase inhibitors, probucol, BCP671, EPC antibodies,
catchins, glycating agents, endothelin inhibitors (for example,
Ambrisentan, Tesosentan, Bosentan), Statins (for example,
Cerivasttin), E. coli heat-labile enterotoxin, and advanced
coatings.
[0414] Generally, anti-infective agents are substances capable of
acting against infection by inhibiting the spread of an infectious
agent or by killing the infectious agent outright, which can serve
to reduce immuno-response without inflammatory response at the
implant site. Anti-infective agents include, but are not limited
to, anthelmintics (mebendazole), antibiotics including
aminoclycosides (gentamicin, neomycin, tobramycin), antifungal
antibiotics (amphotericin b, fluconazole, griseofulvin,
itraconazole, ketoconazole, nystatin, micatin, tolnaftate),
cephalosporins (cefaclor, cefazolin, cefotaxime, ceftazidime,
ceftriaxone, cefuroxime, cephalexin), beta-lactam antibiotics
(cefotetan, meropenem), chloramphenicol, macrolides (azithromycin,
clarithromycin, erythromycin), penicillins (penicillin G sodium
salt, amoxicillin, ampicillin, dicloxacillin, nafcillin,
piperacillin, ticarcillin), tetracyclines (doxycycline,
minocycline, tetracycline), bacitracin; clindamycin; colistimethate
sodium; polymyxin b sulfate; vancomycin; antivirals including
acyclovir, amantadine, didanosine, efavirenz, foscarnet,
ganciclovir, indinavir, lamivudine, nelfinavir, ritonavir,
saquinavir, silver, stavudine, valacyclovir, valganciclovir,
zidovudine; quinolones (ciprofloxacin, levofloxacin); sulfonamides
(sulfadiazine, sulfisoxazole); sulfones (dapsone); furazolidone;
metronidazole; pentamidine; sulfanilamidum crystallinum;
gatifloxacin; and sulfamethoxazole/trimethoprim.
[0415] Generally, necrosing agents are any drugs that cause tissue
necrosis or cell death. Necrosing agents include cisplatin, BCNU,
taxol or taxol derivatives, and the like.
Vascularization Agents
[0416] Generally, vascularization agents include substances with
direct or indirect angiogenic properties. In some cases,
vascularization agents can additionally affect formation of barrier
cells in vivo. By indirect angiogenesis, it is meant that the
angiogenesis can be mediated through inflammatory or immune
stimulatory pathways. It is not fully known how agents that induce
local vascularization indirectly inhibit barrier-cell formation,
however it is believed that some barrier-cell effects can result
indirectly from the effects of vascularization agents.
[0417] Vascularization agents include mechanisms that promote
neovascularization around the membrane and/or minimize periods of
ischemia by increasing vascularization close to the tissue-device
interface. Sphingosine-1-Phosphate (S1P), which is a phospholipid
possessing potent angiogenic activity, is incorporated into a
biointerface membrane of a preferred embodiment. Monobutyrin, which
is a potent vasodilator and angiogenic lipid product of adipocytes,
is incorporated into a biointerface membrane of a preferred
embodiment. In another embodiment, an anti-sense molecule (for
example, thrombospondin-2 anti-sense), which increases
vascularization, is incorporated into a biointerface membrane.
[0418] Vascularization agents can include mechanisms that promote
inflammation, which is believed to cause accelerated
neovascularization in vivo. In one embodiment, a xenogenic carrier,
for example, bovine collagen, which by its foreign nature invokes
an immune response, stimulates neovascularization, and is
incorporated into a biointerface membrane of the preferred
embodiments. In another embodiment, Lipopolysaccharide, which is a
potent immunostimulant, is incorporated into a biointerface
membrane. In another embodiment, a protein, for example, a bone
morphogenetic protein (BMP), which is known to modulate bone
healing in tissue, is incorporated into a biointerface membrane of
a preferred embodiment.
[0419] Generally, angiogenic agents are substances capable of
stimulating neovascularization, which can accelerate and sustain
the development of a vascularized tissue bed at the tissue-device
interface. Angiogenic agents include, but are not limited to,
copper ions, iron ions, tridodecylmethylammonium chloride, Basic
Fibroblast Growth Factor (bFGF), (also known as Heparin Binding
Growth Factor-II and Fibroblast Growth Factor II), Acidic
Fibroblast Growth Factor (aFGF), (also known as Heparin Binding
Growth Factor-I and Fibroblast Growth Factor-I), Vascular
Endothelial Growth Factor (VEGF), Platelet Derived Endothelial Cell
Growth Factor BB (PDEGF-BB), Angiopoietin-1, Transforming Growth
Factor Beta (TGF-Beta), Transforming Growth Factor Alpha
(TGF-Alpha), Hepatocyte Growth Factor, Tumor Necrosis Factor-Alpha
(TNF-Alpha), Placental Growth Factor (PLGF), Angiogenin,
Interleukin-8 (IL-8), Hypoxia Inducible Factor-I (HIF-1),
Angiotensin-Converting Enzyme (ACE) Inhibitor Quinaprilat,
Angiotropin, Thrombospondin, Peptide KGHK, Low Oxygen Tension,
Lactic Acid, Insulin, Leptin, Copper Sulphate, Estradiol,
prostaglandins, cox inhibitors, endothelial cell binding agents
(for example, decorin or vimentin), glenipin, hydrogen peroxide,
nicotine, and Growth Hormone.
[0420] Generally, pro-inflammatory agents are substances capable of
stimulating an immune response in host tissue, which can accelerate
or sustain formation of a mature vascularized tissue bed. For
example, pro-inflammatory agents are generally irritants or other
substances that induce chronic inflammation and chronic granular
response at the implantation-site. While not wishing to be bound by
theory, it is believed that formation of high tissue granulation
induces blood vessels, which supply an adequate, or rich supply of
analytes to the device-tissue interface. Pro-inflammatory agents
include, but are not limited to, xenogenic carriers,
Lipopolysaccharides, S. aureus peptidoglycan, and proteins.
[0421] Other substances that can be incorporated into membranes of
preferred embodiments include various pharmacological agents,
excipients, and other substances well known in the art of
pharmaceutical formulations.
[0422] U.S. Publication No. US-2005-0031689-A1 discloses a variety
of systems and methods by which the bioactive agent can be
incorporated into the biointerface membranes and/or implantable
device. Although the bioactive agent is preferably incorporated
into the biointerface membrane and/or implantable device, in some
embodiments the bioactive agent can be administered concurrently
with, prior to, or after implantation of the device systemically,
for example, by oral administration, or locally, for example, by
subcutaneous injection near the implantation site. A combination of
bioactive agent incorporated in the biointerface membrane and
bioactive agent administration locally and/or systemically can be
preferred in certain embodiments.
[0423] Generally, numerous variables can affect the
pharmacokinetics of bioactive agent release. The bioactive agents
of the preferred embodiments can be optimized for short- and/or
long-term release. In some embodiments, the bioactive agents of the
preferred embodiments are designed to aid or overcome factors
associated with short-term effects (for example, acute
inflammation) of the foreign body response, which can begin as
early as the time of implantation and extend up to about one month
after implantation. In some embodiments, the bioactive agents of
the preferred embodiments are designed to aid or overcome factors
associated with long-term effects, for example, chronic
inflammation, barrier cell layer formation, or build-up of fibrotic
tissue of the foreign body response, which can begin as early as
about one week after implantation and extend for the life of the
implant, for example, months to years. In some embodiments, the
bioactive agents of the preferred embodiments combine short- and
long-term release to exploit the benefits of both. U.S. Publication
No. US-2005-0031689-A1 discloses a variety of systems and methods
for release of the bioactive agents.
[0424] The amount of loading of the bioactive agent into the
biointerface membrane can depend upon several factors. For example,
the bioactive agent dosage and duration can vary with the intended
use of the biointerface membrane, for example, cell
transplantation, analyte measuring-device, and the like;
differences among hosts in the effective dose of bioactive agent;
location and methods of loading the bioactive agent; and release
rates associated with bioactive agents and optionally their carrier
matrix. Therefore, one skilled in the art will appreciate the
variability in the levels of loading the bioactive agent, for the
reasons described above. U.S. Publication No. US-2005-0031689-A1 to
Shults et al. discloses a variety of systems and methods for
loading of the bioactive agents.
Biointerface Membrane Formation onto the Sensor
[0425] Due to the small dimension(s) of the sensor (sensing
mechanism) of the preferred embodiments, some conventional methods
of porous membrane formation and/or porous membrane adhesion are
inappropriate for the formation of the biointerface membrane onto
the sensor as described herein. Accordingly, the following
embodiments exemplify systems and methods for forming and/or
adhering a biointerface membrane onto a small structured sensor as
defined herein. For example, the biointerface membrane of the
preferred embodiments can be formed onto the sensor using
techniques such as electrospinning, molding, weaving,
direct-writing, lyophilizing, wrapping, and the like.
[0426] Although FIGS. 6 to 10 describe systems and methods for the
formation of porous biointerface membranes, including
interconnected cavities and solid portion(s). In some embodiments,
a cell impermeable (second domain) can additionally be formed using
known thin film techniques, such as dip coating, spray coating,
spin coating, tampo printing, and the like, prior to formation of
the interconnected cavities and solid portion(s). Alternatively,
the porous biointerface membrane (e.g., first domain) can be formed
directly onto the sensing membrane.
[0427] FIG. 6 is a flow chart that illustrates the process 150 of
forming a biointerface-coated small structured sensor in one
embodiment. In this embodiment, the biointerface membrane includes
woven or non-woven fibers formed directly onto the sensor.
Generally, fibers can be deposited onto the sensor using methods
suitable for formation of woven- or non-woven fibrous materials. In
some embodiments, the biointerface membrane is electrospun directly
onto the sensor; electrospinning advantageously allows the
biointerface membranes to be made with small consistent fiber
diameters that are fused at the nodes and are without
aggregation.
[0428] In some embodiments, the biointerface membrane is directly
written onto the sensor; direct writing can advantageously allow
uniform deposition of stored patterns (e.g., in a computer system)
for providing consistent and reproducible architectures. In these
embodiments, a curing step is included either during or after the
writing step to solidify the material being written (e.g., heat, UV
curing, radiation, etc.). Direct writing is described in more
detail, below.
[0429] At block 152, one or more dispensers dispense a polymeric
material used to form the fibers. A variety of polymeric materials
are contemplated for use with the preferred embodiments, including
one or more of silicone, polytetrafluoroethylene, expanded
polytetrafluoroethylene, polyethylene-co-tetrafluoroethylene,
polyolefin, polyester, polycarbonate, biostable
polytetrafluoroethylene, homopolymers, copolymers, terpolymers of
polyurethanes, polypropylene (PP), polyvinylchloride (PVC),
polyvinylidene fluoride (PVDF), polyvinyl alcohol (PVA),
polybutylene terephthalate (PBT), polymethylmethacrylate (PMMA),
polyether ether ketone (PEEK), polyamides, polyurethanes,
cellulosic polymers, poly(ethylene oxide), poly(propylene oxide)
and copolymers and blends thereof, polysulfones and block
copolymers thereof including, for example, di-block, tri-block,
alternating, random and graft copolymers.
[0430] The coating process can be performed in a vacuum or in a
gaseous medium, which environment can affect the architecture of
the biointerface membrane as is appreciated by one skilled in the
art.
[0431] In embodiments wherein the biointerface is electrospun onto
the sensor, the dispenser dispenses a charged liquefied polymer
within an electric field, to thereby form a jet of polymer fibers,
for example, such as described in PCT International Publication No.
WO 2005/032400, which is incorporated herein by reference in its
entirety.
[0432] In embodiments wherein the biointerface is directly-written
onto the sensor, a dispenser dispenses a polymer solution using a
nozzle with a valve, or the like, for example as described in U.S.
Publication No. US-2004-0253365-A1. In general, a variety of
nozzles and/or dispensers can be used to dispense a polymeric
material to form the woven or non-woven fibers of the biointerface
membrane.
[0433] In general, a direct-write patterning system is suitable for
either fine-pattern micro dispensing and/or fine-focused laser-beam
writing over flat or conformal surfaces to create exact replicas of
a preferred biointerface structure. In certain embodiments, the
biointerface materials described herein can be deposited using
these integrated tool technologies for the direct-write deposition
and laser micromachining of a wide variety of biointerface
architectures described herein. Additionally, the direct-write
patterning system can provide the capability for concurrent
detection and imaging methods during additive and subtractive
processes.
[0434] In another aspect, alternative embodiments of the
direct-writing deposition technique utilize a tool in which
constituent materials can be dispensed through multiple, discrete
dispensing heads. In yet another alternative embodiment, the
biointerface structure is directly written onto a removable
substrate, after which the substrate is removed and the
biointerface applied to the sensor (e.g., wrapped around the sensor
or the sensor is inserted into the biointerface).
[0435] At block 154, the dispenser(s) is moved relative to the
sensor and/or the sensor is moved relative to the dispenser(s) so
as to coat the sensor with the fibers. In embodiments wherein the
biointerface membrane is electrospun onto the sensor, the
dispenser(s) can change the direction and/or magnitude of the
electric field during motion in order to effect the orientation of
the polymer fibers on the sensor. Additionally, the path of the
dispenser is preferably selected so as to coat the portions of or
the entire object. In one exemplary embodiment, wherein it is
desirable for the biointerface membrane to substantially
circumscribe the sensor (e.g., a substantially cylindrical shape),
the dispenser can be moved along a helix path, a circular path, a
zigzag path, or the like. Additionally, the dispenser can move
rotationally and/or translationally relative to the sensor. The
number of sweeps is preferably selected according to the desired
architecture of the biointerface membrane. Additionally, the
density of the fibers and/or the type of liquefied polymer can be
changed from one sweep to the other to thereby control the
architecture of the membrane.
[0436] In embodiments wherein the biointerface membrane is directly
written onto the sensor, the dispenser is programmed to write a
pattern that creates the desired membrane architecture, including
the interconnected cavities and solid portion(s). Namely, the
dispenser is programmed to move in the x, y, and optionally z
direction in order to create the desired membrane architecture.
See, for example, U.S. Publication No. US-2004-0253365-A1 cited
above.
[0437] Although the preferred embodiments described moving the
dispenser(s) relative to the sensor, alternatively, the dispenser
can remain stationary and the sensor moved, as is appreciated by
one skilled in the art.
[0438] In some embodiments, the sensor is moved in a rotational or
translational motion, which can be performed in combination with,
or instead of, movement of the dispenser. In this step, the sensor
is moved so as to ensure coating throughout the entirety of the
biointerface region (or a portion thereof). In one exemplary
embodiment, wherein a substantially circumscribing biointerface
membrane is desired (e.g., for a substantially cylindrically shaped
sensing sensor) such as illustrated in FIG. 3A, the sensor can be
rotated so to aid in coating the entire circumference of the
sensor. In another exemplary embodiment, wherein a substantially
planar biointerface membrane is desired (e.g., for a substantially
planar sensor), the sensor can be translated so as to aid in
coating the desired planar surface area.
[0439] FIG. 7 is a flow chart that illustrates the process 160 of
forming a biointerface-coated sensor in an alternative embodiment.
In this embodiment, the interconnected cavities and solid
portion(s) of the biointerface membrane are amorphous in
configuration, such as illustrated in FIGS. 5A and 5B, for
example.
[0440] At block 162, a selectively removable porogen (e.g., porous
mold) is formed by spraying, coating, rolling, or otherwise forming
selectively removable particles, for example, sugar crystals, onto
the surface of the sensor. Additional examples of materials
suitable as selectively removable mold material include
thermoplastic polymers such as waxes, paraffin, polyethylene,
nylon, polycarbonate, or polystyrene in naturally available
particles or processed into specific sizes, shapes, molded forms,
spheres or fibers, salt or other particles which cannot be made to
inherently stick together coated with sugar, and certain drug
crystals such as gentamycin, tetracycline, or cephalosporins. In
general, any dissolvable, burnable, meltable, or otherwise
removable particle, which can be made to stick together, could be
used. Preferably, the particles have shapes and sizes substantially
corresponding to the desired cavity dimensions, such as described
in more detail above. In some embodiments, the particles are made
to adhere to the sensor by environmental conditions, for example,
humidity can be used to cause sugar to adhere to the sensor.
[0441] In some embodiments, the particles are made to coalesce to
provide the desired interconnectivity between the cavities. In an
exemplary porous silicone embodiment, sugar crystals are exposed to
a humid environment sufficient to cause coalescence of the sugar
crystals. In some alternative embodiments, other molds can be used
in the place of the particles described above, for example, coral,
self-assembly beads, etched and broken silicon pieces, glass frit
pieces, and the like, as shown in FIG. 11B.
[0442] At block 164, a material (e.g., a moldable or conformable
material) is filled or coated into the interconnected cavities of
the mold using methods common in the art of polymer processing, for
example, injecting, pressing, vacuuming, vapor depositing,
extruding, pouring, and the like. Examples of materials suitable
for the resulting porous device include polymers, metals, metal
alloys, ceramics, biological derivatives, and combinations thereof,
in solid or fiber form. In an exemplary porous silicone embodiment,
silicone is pressed into the interconnected cavities of the
mold.
[0443] At block 166, the material is substantially cured or
solidified to form the solid portion(s) of the biointerface
membrane. Solidification of the material can be accelerated by
supplying dry air (which can be heated) to the material, for
example. Additionally, freezing, freeze drying or vacuum
desiccation, with or without added heat, can also be utilized to
cause the material to solidify. In some circumstances, a skin or
any excess material can be removed (e.g., shaved, etched, or the
like) after curing. In the exemplary porous silicone embodiment, an
outer skin of silicone is removed to expose the interconnected
cavities at an outer surface.
[0444] At block 168, the selectively removable porogen (e.g.,
porous mold) is dissolved, melted, etched, or otherwise removed,
leaving interconnecting cavities within the solid portion (FIG.
11A). Preferably, the selectively removable porogen is readily
removable without significantly altering the final product (or
product material). This removal can be by dissolution by some
solvent that does not significantly dissolve the final product
material. Alternatively, the mold material can be melted (or
burned) out of the final product material if the melting point (or
burning point) of the mold material is below that of the final
product material. In the exemplary porous silicone embodiment,
water is used to dissolve the sugar crystals.
[0445] FIG. 8 is a flow chart that illustrates the process 170 of
forming a biointerface-coated small structured sensor in another
alternative embodiment. In this embodiment, the interconnected
cavities and solid portion(s) of the biointerface membrane are
amorphous in configuration, such as illustrated in FIGS. 4A and 4B,
for example, and the solid portion is molded around the sensor.
[0446] At block 172, a selectively removable porogen is formed by
filling a shaped cavity with selectively removable particles, for
example, sugar crystals, wherein the sensor is located within the
shaped cavity, and wherein the selectively removable particles
substantially surround the sensor. Additional examples of materials
suitable as selectively removable mold material are described with
reference to block 162, above. In some embodiments, the shaped
cavity mold is formed from a selectively removable material (e.g.,
sacrificial cavity mold) similar the selectively removable
particles described above. One such example includes a tube formed
from a dissolvable polymer. Alternatively, the shaped cavity can be
a non-selectively removable material, and instead, a sacrificial
layer of selectively removable material is formed directly onto the
cavity walls, enabling the removal of the biointerface membrane
after dissolution of the sacrificial layer.
[0447] Preferably the shape of the cavity mold substantially
corresponds to the desired final shape of the biointerface
membrane. In one exemplary embodiment, the cavity mold is
substantially cylindrical, for example using a syringe or cannula
as the cavity mold.
[0448] In some embodiments, the particles are made to coalesce to
provide the desired interconnectivity between the cavities. In an
exemplary porous silicone embodiment, sugar crystals are exposed to
humidity or spray of water sufficient to cause coalescence of the
sugar crystals. In some alternative embodiments, other molds can be
used in the place of the particles described above, for example,
coral, self-assembly beads, etched and broken silicon pieces, glass
frit pieces, and the like.
[0449] At block 174, a material (e.g., a moldable or conformable
material) is filled into the interconnected cavities of the mold
using methods common in the art of polymer processing, for example,
injecting, pressing, vacuuming, vapor depositing, pouring, and the
like. Examples of materials suitable for the resulting porous
device are described in more detail with reference to block 164,
above. In an exemplary porous silicone embodiment, silicone is
pressed into the interconnected cavities of the mold.
[0450] At block 176, the material is substantially cured or
solidified to form the solid portion(s) of the biointerface
membrane. Solidification of the material can be accelerated as
described in more detail with reference to block 166, above.
[0451] At block 178, the selectively removable porogen is
dissolved, melted, etched, or otherwise removed, leaving
interconnecting cavities within the solid portion surrounding the
sensor. In some embodiments, wherein a sacrificial layer is formed
as described above, the sacrificial layer can be removed before,
during, or after the removal of the selectively removable porogen.
In some embodiments, the final product is removed from the cavity
mold before, during, or after the removal of the selectively
removable porogen.
[0452] Preferably, the selectively removable porogen is readily
removable without significantly altering the final product (or
product material). This removal can be by dissolution by some
solvent that does not significantly dissolve the final product
material. Alternatively, the mold material can be melted (or
burned) out of the final product material if the melting point (or
burning point) of the mold material is below that of the final
product material. In one exemplary embodiment, a sacrificial tube
forms the mold cavity; wherein the sacrificial tube is removed
prior to, during, or after dissolution of the selectively removable
porogen. One skilled in the art can appreciate a variety of
modifications or combinations of the above described removal step
without departing from the spirit of the invention.
[0453] FIG. 9 is a flow chart that illustrates the process 180 of
forming a biointerface-wrapped sensor in one embodiment. In this
embodiment, the interconnected cavities and solid portion(s) of the
biointerface membrane can be fibrous or amorphous in configuration.
In fact, substantially any biointerface membrane with an
architecture as described in more detail above, which is formed in
substantially any manner, can be used with this embodiment.
[0454] At block 182, a sensor is manufactured and provided, wherein
the sensor is formed with a small structure as defined herein.
[0455] At block 184, a biointerface membrane with an architecture
as described herein is manufactured in substantially any desired
manner, wherein the biointerface membrane is formed substantially
as a sheet or tube of membrane. Biointerface membranes suitable for
wrapping around the sensor and providing the desired host interface
are described in more detail above (see section entitled,
"Architecture of the First Domain.")
[0456] At block 186, the biointerface membrane is wrapped around
the sensor manually, or using an automated device, as can be
appreciated by one skilled in the art. Namely, the biointerface
membrane is wrapped such that it substantially surrounds the
sensor, or the sensing mechanism of the sensor (e.g., the
electroactive surfaces or sensing membrane). The number of wraps
can be from less than 1 to about 100, preferably 1, 11/2, 2, 21/2,
3, 31/2, 4, 5, 6, 7, 8, 9, 10, or more. The number of wraps depends
on the architecture of the sheet of biointerface membrane, and the
desired architecture of the biointerface surrounding the
sensor.
[0457] In some embodiments, the circumference (or a portion thereof
(e.g., an edge)) of the biointerface membrane with an architecture
as described herein can be adhered or otherwise attached or sealed
to form a substantially consistent outer surface (of the
biointerface membrane). In an aspect of this embodiment, the
biointerface membrane is wrapped around the sensor one time,
wherein the "wrap" includes a tubular biointerface membrane
configured to slide over the sensor (or sensing mechanism), for
example, be stretching the tubular biointerface membrane and
inserting the sensor therein.
[0458] FIG. 10 is a flow chart that illustrates the process 190 of
forming a sensing biointerface in one embodiment. In this
embodiment, the sensor is inserted into the biointerface membrane
so that it is encompassed therein.
[0459] At block 192, a biointerface membrane is manufactured in
substantially any desired manner. Biointerface membranes suitable
for the sensing biointerface are described in more detail above
(see for example, section entitled, "Architecture of the First
Domain"). In some embodiments, the biointerface membrane is molded
into the desired final shape to surround the sensor and implant
into a host. Alternatively, the biointerface membrane can be
provided as a sheet of bulk material.
[0460] At block 194, a particularly shaped or sized biointerface
membrane can be (optionally) cut. Namely, in embodiments wherein
the biointerface membrane is provided in bulk, e.g., as a sheet of
material, the desire shape or size can be cut there from. In these
embodiments, bulk biointerface membrane sheet is preferably of the
appropriate thickness for the desired final product. In one
exemplary embodiment, the biointerface membrane (bulk sheet) is
compressed, for example between two substantially rigid structures,
and the final size/shape biointerface membrane cut there from,
after which the biointerface membrane is released. While not
wishing to be bound by theory, it is believed that by compressing
the biointerface membrane during cutting, a more precise shape can
be achieved. Biointerface membranes can have sufficient elasticity,
such that the thickness is returned after release from compression,
as is appreciated by one skilled in the art.
[0461] At block 196, a sensor is inserted into the biointerface
membrane. Preferably, the sensor is inserted into the membrane such
that the sensing mechanism contacts at least one or more of the
interconnected cavities so that the host analyte can be measured.
Alternatively, the biointerface can be formed from a material that
allows the flux of the analyte there through. In some embodiments,
the sensor is inserted with the aid of a needle. Alternatively, the
sensor is formed with appropriate sharpness and rigidity to enable
insertion through the biointerface membrane.
[0462] In some embodiments, an anchoring mechanism, such as a barb,
is provided on the sensor, in order to anchor the sensor within the
biointerface membrane (and/or host tissue). A variety of additional
or alternative aspects can be provided to implement the
biointerface membrane surrounded sensors of the preferred
embodiments.
[0463] A porous membrane material applied to the sensor can act as
a spacer between the sensor and the surrounding tissue at the site
of sensor insertion, in either the short-term or long-term sensors.
For example, a spacer from 60-300 microns thick can be created of
porous silicone having pore sizes of 0.6 microns and greater (e.g.,
up to about 1,000 microns or more). When inserted into the tissues,
the adipose cells come to rest against the outermost aspects of the
porous membrane, rather than against the surface of the sensor
(FIG. 2C), allowing open space for transport of water-soluble
molecules such as oxygen and glucose.
[0464] Porous membrane material can be manufactured and applied to
a sensor using any advantageous method known to one skilled in the
art. As discussed elsewhere, porous membranes can be manufactured
from a variety of useful materials known in the art, depending upon
the desired membrane parameters.
[0465] FIG. 11A is a scanning electron micrograph showing a
cross-section of an exemplary porous silicone tube that does not
contain a sensor. Note the open porous structure of cavities and
channels within the solidified silicone. Porous silicone can be
manufactured and applied to the sensor by a variety of means. The
material in FIG. 11A, for example, was formed by sieving sugar to
give crystals having a size and shape approximate to that of the
desired pore size. The sugar was humidified and then compressed
into a mold. The mold was then baked, to harden the sugar within
the mold. Silicone was forced into the mold and then cured. After
the silicone was cured, the mold was removed and the sugar
dissolved away. A sensor could subsequently be inserted into the
porous silicone tube.
[0466] FIG. 11B is a scanning electron micrograph of sugar molded
onto a sensor. In this example, a sugar mold was formed directly on
the sensor. Note the clumps of sugar crystals attached to the
surface of the sensor. In this example, the sensor was placed into
the mold, which was then filled with humidified sugar crystals. The
mold containing the sensor and sugar was baked to solidify the
sugar on the surface of the sensor. The sensor, with sugar crystals
attached, was removed from the mold, in order to prepare the
electron micrograph. In some embodiments, the sensor can be rolled
in the humidified sugar, to attach a layer of sugar to the sensor
surface, and then baked to solidify the sugar. In some embodiments,
the sugarcoated sensor can be rolled in humidified sugar additional
times to form a thicker sugar mold (i.e., 2 or more layers) around
the sensor. In some embodiments, silicone is pumped or injected
into the solidified sugar and cured. After curing, the sugar is
removed, such as by washing, to give a porous silicone covered
sensor.
[0467] In an alternative embodiment, porous silicone is pre-formed
as a sheet or plug and then applied to the sensor. For example, a
sugar mold lacking a sensor therein is formed using the usual
means. As previously described, silicone is injected into the mold
and then cured. After the mold material is removed from the cured
silicone, the sensor is inserted into the plug, thereby creating a
sensor having a porous silicone biointerface membrane.
[0468] Alternatively, a thin sheet of porous silicone is
manufactured and then wrapped around the sensor. For example, a
thin porous silicone sheet is manufactured by pressing a thin layer
of sieved, humidified sugar into a Petri dish. The sugar is baked.
Silicone is applied to the sugar mold by injection, pressing, or
the like, and then cured. The sugar is removed from the porous
silicone sheet, such as by washing. The manufactured porous
silicone is then wrapped around the sensor to form a biointerface
membrane of a desired thickness.
[0469] In still another embodiment, other materials can be used to
manufacture the biointerface membrane. For example, the sensor can
be wrapped in a layer of ePTFE having a pore size of about 0.6
microns and above, to create a layer about 12-100 microns thick.
See U.S. Pat. No. 6,862,465. In yet another embodiment, the spacer
can be either a smooth or porous hydrogel.
Methods of Use
[0470] One aspect of the present invention contemplates new methods
of use to reduce noise. In one embodiment, noise is reduced by
first providing a device of the present invention, such as an
implantable analyte sensor, preferably a glucose sensor. The sensor
is pre-inserted through the host's skin and into the host. The term
"pre-insertion" or "pre-inserted" as used herein is a broad term
and is used in its ordinary sense, including, without limitation,
to refer to inserting a sensor a period of time (e.g., a "waiting
period") before it is to be used, such as about 1-24 hours or
longer, e.g., without operatively connecting the sensor to the
electronics. The period of time is associated with an amount of
time necessary for wound healing to occur. For example, the wound
healing process progresses for the first few hours or days.
Preferably, interferents that build up around the sensor will be
diluted or removed by bulk fluid flow and/or an increase in the
fluid bulk around at least a portion of the sensor.
[0471] In one embodiment of the present method, the host waits
about 1, 2, 3, 4, 5, 6, 7, 8, 9, 10, 11 or 12 hours or longer,
before operatively connecting the electronics to the sensor. In
another embodiment, the host can wait about 24, 36 or 48 hours or
longer, before connecting the electronics. In yet another
embodiment, the host can wait about 1-2, 2-4, 4-6, 6-8, 8-10,
10-12, 12-15, 12-24, 12-36, 12-48, 24-36 or 36-48 hours or longer,
before operatively connecting the electronics.
[0472] In some embodiments, a cap is provided to cover the
electrical components (e.g., contacts), until the electronics are
coupled to the sensor. The cap can be manufactured of any
convenient material, such as plastic, tape, foil, glass or a
combination thereof. The cap can attach to the sensor using any
convenient method known to those of skill in the art. For example,
the cap can attach with a snap fit, adhesive, pins, or the like.
After the waiting period has been completed, the host can remove
the cap and operably attach the electronics to the sensor.
[0473] After the electronics have been operatively connected to the
sensor (e.g., the sensor electronics are connected to the sensor),
a signal from the sensor is detected, as described in detail above.
The sensor will be used for a prescribed period of time, after
operably connecting the electronics to the sensor (e.g., in
addition to pre-insertion or a waiting period). For example, a
3-day sensor will be used for 3-days and then removed (after three
days of data collection). In another example, a 7-day sensor will
be removed after seven days of data collection. In the case of
sensors configured for shorter or longer periods of use, the sensor
will be removed after that period of time. In additional
embodiments, noise can be reduced by sensor pre-insertion and/or
overlapping sensor insertion as described in co-pending U.S. patent
application Ser. No. 11/373,628, filed Mar. 9, 2006 and entitled
"SYSTEM AND METHODS FOR PROCESSING ANALYTE SENSOR DATA FOR SENSOR
CALIBRATION."
[0474] In another embodiment, of the present invention, a second
sensor can be pre-inserted into the host before removal of the
first sensor. Preferably, the amount of time the second sensor is
pre-inserted, before the first sensor, corresponds to the waiting
period required for wound healing, as described above. For example,
if the waiting period is 6-hours, the host would pre-insert the
second sensor at least about 6 hours before he removed the first
sensor. In another example, if the waiting period is 24 hours, he
would pre-insert the second sensor on the second-to-the-last day
(e.g., about 24 hours before removal of the first sensor).
[0475] Pre-inserting a sensor, waiting a period of time associated
with wound healing and then operatively connecting the electronics,
allows time for a fluid pocket to form and/or wound healing to
progress, and thereby avoids presenting data to a user during the
early time after sensor insertion when early, sedentary noise is
most likely to occur, while maintaining the full period of sensor
utility (e.g., number of days the sensor is to be used, such as but
not limited to 1, 2, 3, 4, 5, 6, 7, 8, 9, 10, 11, 12, 13, 14, or 15
days). By pre-inserting a sensor, in a series of sensor used by a
host (e.g., a host changes to a new 5-day sensor after every 5 days
of use), the host can have daily use of a sensor without having
days of disuse due to a waiting period such as can be necessitated
by procedures such as calibration, break-in or wound healing.
Advantageously, the host is provided with an extended period of
continuous use (instead of the intermittent periods of use required
by some implantable analyte sensors, such as implantable glucose
sensors) and is provided with substantially increased or improved
information/data on his analyte levels (e.g., glucose
concentration) so that he can make more informed treatment
decisions. Accordingly, due to more informed treatment decisions,
the host can benefit from improved disease management, with
improved health and quality of life.
[0476] Methods and devices that are suitable for use in conjunction
with aspects of the preferred embodiments are disclosed in U.S.
Pat. No. 4,994,167; U.S. Pat. No. 4,757,022; U.S. Pat. No.
6,001,067; U.S. Pat. No. 6,741,877; U.S. Pat. No. 6,702,857; U.S.
Pat. No. 6,558,321; U.S. Pat. No. 6,931,327; and U.S. Pat. No.
6,862,465.
[0477] Methods and devices that are suitable for use in conjunction
with aspects of the preferred embodiments are disclosed in U.S.
Publication No. US-2005-0176136-A1; U.S. Publication No.
US-2005-0251083-A1; U.S. Publication No. US-2005-0143635-A1; U.S.
Publication No. US-2005-0181012-A1; U.S. Publication No.
US-2005-0177036-A1; U.S. Publication No. US-2005-0124873-A1; U.S.
Publication No. US-2005-0051440-A1; U.S. Publication No.
US-2005-0115832-A1; U.S. Publication No. US-2005-0245799-A1; U.S.
Publication No. US-2005-0245795-A1; U.S. Publication No.
US-2005-0242479-A1; U.S. Publication No. US-2005-0182451-A1; U.S.
Publication No. US-2005-0056552-A1; U.S. Publication No.
US-2005-0192557-A1; U.S. Publication No. US-2005-0154271-A1; U.S.
Publication No. US-2004-0199059-A1; U.S. Publication No.
US-2005-0054909-A1; U.S. Publication No. US-2005-0112169-A1; U.S.
Publication No. US-2005-0051427-A1; U.S. Publication No.
US-2003-0032874-A1; U.S. Publication No. US-2005-0103625-A1; U.S.
Publication No. US-2005-0203360-A1; U.S. Publication No.
US-2005-0090607-A1; U.S. Publication No. US-2005-0187720-A1; U.S.
Publication No. US-2005-0161346-A1; U.S. Publication No.
US-2006-0015020-A1; U.S. Publication No. US-2005-0043598-A1; U.S.
Publication No. US-2003-0217966-A1; U.S. Publication No.
US-2005-0033132-A1; U.S. Publication No. US-2005-0031689-A1; U.S.
Publication No. US-2004-0045879-A1; U.S. Publication No.
US-2004-0186362-A1; U.S. Publication No. US-2005-0027463-A1; U.S.
Publication No. US-2005-0027181-A1; U.S. Publication No.
US-2005-0027180-A1; U.S. Publication No. US-2006-0020187-A1; U.S.
Publication No. US-2006-0036142-A1; U.S. Publication No.
US-2006-0020192-A1; U.S. Publication No. US-2006-0036143-A1; U.S.
Publication No. US-2006-0036140-A1; U.S. Publication No.
US-2006-0019327-A1; U.S. Publication No. US-2006-0020186-A1; U.S.
Publication No. US-2006-0020189-A1; U.S. Publication No.
US-2006-0036139-A1; U.S. Publication No. US-2006-0020191-A1; U.S.
Publication No. US-2006-0020188-A1; U.S. Publication No.
US-2006-0036141-A1; U.S. Publication No. US-2006-0020190-A1; U.S.
Publication No. US-2006-0036145-A1; U.S. Publication No.
US-2006-0036144-A1; U.S. Publication No. US-2006-0016700-A1; U.S.
Publication No. US-2006-0142651-A1; U.S. Publication No.
US-2006-0086624-A1; U.S. Publication No. US-2006-0068208-A1; U.S.
Publication No. US-2006-0040402-A1; U.S. Publication No.
US-2006-0036142-A1; U.S. Publication No. US-2006-0036141-A1; U.S.
Publication No. US-2006-0036143-A1; U.S. Publication No.
US-2006-0036140-A1; U.S. Publication No. US-2006-0036139-A1; U.S.
Publication No. US-2006-0142651-A1; U.S. Publication No.
US-2006-0036145-A1; and U.S. Publication No.
US-2006-0036144-A1.
[0478] Methods and devices that are suitable for use in conjunction
with aspects of the preferred embodiments are disclosed in 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. 11/335,879 filed Jan. 18, 2006 and entitled
"CELLULOSIC-BASED INTERFERENCE DOMAIN FOR AN ANALYTE SENSOR"; U.S.
application Ser. No. 11/334,876 filed Jan. 18, 2006 and entitled
"TRANSCUTANEOUS ANALYTE SENSOR"; U.S. application Ser. No.
11/333,837 filed Jan. 17, 2006 and entitled "LOW OXYGEN IN VIVO
ANALYTE SENSOR".
[0479] All references cited herein, including but not limited to
published and unpublished applications, patents, and literature
references, are incorporated herein by reference in their entirety
and are hereby made a part of this specification. To the extent
publications and patents or patent applications incorporated by
reference contradict the disclosure contained in the specification,
the specification is intended to supersede and/or take precedence
over any such contradictory material.
[0480] The term "comprising" as used herein is synonymous with
"including," "containing," or "characterized by," and is inclusive
or open-ended and does not exclude additional, unrecited elements
or method steps.
[0481] All numbers expressing quantities of ingredients, reaction
conditions, and so forth used in the specification are to be
understood as being modified in all instances by the term "about."
Accordingly, unless indicated to the contrary, the numerical
parameters set forth herein are approximations that may vary
depending upon the desired properties sought to be obtained. At the
very least, and not as an attempt to limit the application of the
doctrine of equivalents to the scope of any claims in any
application claiming priority to the present application, each
numerical parameter should be construed in light of the number of
significant digits and ordinary rounding approaches.
[0482] 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.
* * * * *