Note: Descriptions are shown in the official language in which they were submitted.
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Implantable Analyte Sensor
BACKGROUND
The present invention relates to implantable analyte sensors.
Several implantable glucose sensors have been developed. Examples include
those described in
U.S. Patent numbers 5,387,327; 5,411,647; and 5,476,776; as well as those
described in PCT
International Publication numbers WO 91/15993; WO 94/20602; WO 96/06947; and
WO
97/19344. The implantable glucose sensors usually include a polymer substrate,
with metal
electrodes printed on the surface of the substrate. A biocompatible membrane
covers the elec-
trodes, allowing glucose to reach the electrodes, while excluding other
molecules, such as pro-
teins. Electrochemistry, often with the aid of enzymes at the electrodes, is
used to determine
the quantity of glucose present. The glucose sensor is implanted into a
patient, and the elec-
trodes may be attached via wires that pass out of the patient's body to
external circuitry that
controls the electrodes, measures and reports the glucose concentration.
Alternatively, all or
part of this external circuitry may be miniaturized and included in the
implantable glucose
sensor. A transmitter, such as that described in WO 97/19344, may even be
included in the
implantable glucose sensor, completely eliminating the need for leads that
pass out of the
patient.
A problem associated with an amperometric glucose sensor is unstable signals.
This may result
from degradation of the enzyme from interaction with protein, leakage of the
enzyme, and/or
fouling of the electrode. The usual way to overcome this is to use the above
described biocom-
patible membrane, or a coating. However, several problems are also associated
with these
membranes. For example, Nafion-based biosensor membranes exhibit cracking,
flaking, pro-
tein adhesion, and calcium deposits. Mineralization of polymer-based membranes
occurs in
the biological environment, resulting in cracking and changes in permeability.
The tortuous
porosity associated with polymer membranes has also been shown to be important
in mem-
brane stability and mineralization in vivo. Biological components, which enter
pores or voids
in the material, cause metabolic shadows, which are loci for ion and calcium
accumulation.
This situation, coupled with the fact that mineral deposits have
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been known to propagate surface fractures in polymeric membranes, presents a
potentially
serious problem for implantable glucose sensors.
In polymer membranes the pore size distribution usually follows some kind of
probability
distribution (e.g. gaussian), which leaves a finite probability for large
proteins to eventually
transfer through the membrane. Drift may be caused by this leakage or
inadequate diffusion
properties, and events at the body-sensor interface such as biofouling and
protein adsorption,
encapsulation with fibrotic tissue, and degradation of the device material
over time.
Currently, membranes with nominal pore sizes as small as 20 nm are available.
Even so, the
filtration at these dimensions is far from absolute. The most common filters
are polymeric
membranes formed from a solvent-casting process, which result in a pore size
distribution
with variations as large as 30%. The use of ion-track etching to form
membranes (e.g.
MILLPORE ISOPORE) produces a much tighter pore size distribution (~10%).
However,
these membranes have low porosities (<109 pores/cm2), limited pore sizes, and
the pores are
randomly distributed across the surface. Porous alumina (e.g. WHATMAN) has
also been
used to achieve uniform pores. Although the aluminas typically have higher
pore densities
(,101°/cm2), only certain pore sizes (typically greater than 20
nanometers) can be achieved
and the pore configurations and arrangements are difficult to control.
BRIEF SUMMARY
In one aspect, the present invention is an implantable analyte sensor,
comprising a substrate,
electrodes on the substrate, and a membrane on the electrodes. The membrane
comprises
elemental silicon.
In another aspect, the present invention relates to a method of making an
implantable analyte
sensor, comprising covering electrodes with a membrane. The electrodes are on
a substrate,
and the membrane comprises elemental silicon.
The invention also relates to implantable analyte sensors that exhibit a
signal drift of less than
20% per day in vivo.
BRIEF DESCRIPTION OF THE DRAWINGS
The following drawings form part of the present specification and are included
to further
demonstrate certain aspects of the present invention. The invention may be
better understood
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3
by reference to one or more of these drawings in combination with the detailed
description of
specific embodiments presented herein:
Figured-9 illustrate the process of making a membrane for use in an embodiment
of the pre-
sent invention;
Figure 10 shows a cut-away view of an implantable analyte sensor;
Figure 11 shows an exploded view of an implantable analyte sensor; and
Figure 12 shows a cut-away view of an implatable analyte sensor.
DETAILED DESCRIPTION
Figure 10 shows a cut awav view of an embodiment of the present invention. In
the figure, an
implantable analyte sensor 2 includes a substrate 6 on which are electrodes 8
and 8. The elec-
trodes are covered with a membrane 4. Leads 12 and 12 allow for electrically
connecting the
implantable analyte sensor to external circuitry (not shown). The implantable
analyte sensor
also includes an external coating 16 and an internal coating 14.
Figure 11 shows an exploded view of an embodiment of the present invention.
The internal
and external coatings are not included in the figure for clarity. As shown in
the figure, the
implantable analyte sensor 2 includes the electrodes 8 and 8 on the substrate
6 surface, which
are electrically connected with microelectronic circuitry 10. The
microelectronic circuitry is
electrically connected to leads 12 and 12, which allow for electrically
connecting the
implantable analyze sensor to external circuitry (not shown). The electrodes
are covered with
the membrane 4.
Figure 12 shows a cut away view of an embodiment of the present invention
similar to that
shown in Figure 10, except for the presence of a third electrode 8 and a third
lead 12. Although
so illustrated, the number of electrodes may be different from the number of
leads.
The membrane is composed of a hard material that has been micromachined.
Preferably, the
membrane comprises elemental silicon, but other hard, biocompatable materials
that can be
micromachined are possible, such as metals (for example titanium), ceramics
(for example,
silica or silicon nitride), and polymers (such as polytetrafluoroethylene,
polymethyl-
methacrylate, polystyrenes and silicones). Micromachining is a process that
includes photo-
lithography, such as that used in the semiconductor industry, to remove
material from, or add
material too, a substrate. These techniques are well known, and are described,
in Encyclopedia
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of Chemical Technology, Kirk-Othmer, Volume 14, pp. 677-709 ( 1995);
Semiconductor
Device Fundamentals, Robert F. Pierret, Addison-Wesley, 1996; and Microchip
Fabrication
3rd. edition, Peter Van Zant, McGraw-Hill, 1997. A detailed fabrication method
for a mem-
brane comprising elemental silicon is described in the dissertation of Derek
James Hansford,
submitted in partial satisfaction of the requirements for the degree of Doctor
of Philosophy in
Engineering-Materials Science and Mineral Engineering in the Graduate Division
of the
University of California, Berkeley, submitted in the spring of 1999.
A special property of the membrane is a defined pore size, which has a small
size distribution
compared to the size distribution of standard membranes. Due to tight
tolerances in the
manufacturing process, the pore size can be controlled at precise diameters,
for example 1 to
50 nm, or 5 to 20 nm, or even 5 to 15 nm (such as 12 nm, 18 nm or even 25 nm),
with a varia-
tion of +/- 0.01-20%, +/-0.1-10% or even +/-1-5%. Therefore molecules above
this size can be
excluded with high certainty, since the size distribution has the shape of a
top hat, rather than
a bell curve, and hence pore sizes above, for example 12 nm, 18 nm, 25 nm or
50 nm are not
present. These membranes can exclude interfering molecules, such as proteins,
which could
otherwise cause major drift problems of the sensor, when the sensor is
implanted in vivo.
Signal drift is a change in the magnitude of the signal from a sensor which is
unrelated to
changes in analyte concentration. The amount of signal drift is based on the
magnitude of the
signal prior to the drifting. f referably, the implantable analyte sensors of
the present invention
exhibit a signal drift of less than 20~% per day in vivo, more preferably less
than 10% per day in
vivo, most preferably less than S~?4~ per day in vivo.
Membranes for use in the present invention may be characterized by a glucose
diffusion test
and an albumin diffusion test. These tests are described below. Preferably,
the membrane has a
glucose diffusion test result of at least 1 mg/dl in 330 min., more preferably
at least 10 mg/dl in
330 min., even more preferably at least 30 mg/dl in 330 min., and most
preferably at least 60
mg/dl in 330 min. Preferably, the membrane has an albumin diffusion test
result of at most
0.1 g/dl in 420 min., more preferably at most 0.05 g/dl in 420 min., even more
preferably at
most 0.01 g/dl in 420 min., and most preferably at most 0.001 g/dl in 420 min.
The manufacturing process of the membranes may allow a simple and economical
production
of small, implantable analvte sensors. For example, the membranes can be first
manufactured,
and then on a substrate, the electrodes for the sensor and the electrical
connectors can be
formed. Preferably, the substrate is silicon, but other materials are
possible, such as ceramics,
or polymers. If desired, electronic components, for example, amplifiers,
filters, transmitters
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and/or signal preconditioning components, can easily be incorporated in this
layer. In par-
ticular, if the substrate comprises elemental silicon, well known integrated
circuit technology
may be used to place all the circuitry in miniaturized form on a single chip.
There are two possible approaches to attach the substrate and the membrane,
when a reagent
5 is included in the sensor:
The substrate and the membrane are thermally bonded before the reagent is
deposited on
the electrodes. In this case, an opening, preferably in the membrane is
provided (since this
may be manufactured with a micromachining process, an opening is easily
generated
during one of the processing steps). In the case where multiple membranes are
formed as
a single piece, and or multiple substrates are formed as a single piece, after
thermal
bonding, a further etching step may be used to separate the individual
membrane/
substrate units. The reagent is deposited through the individual openings and
the open-
ings are sealed using, for example a polymer sealant. The individual sensors
are then sepa-
rated, incorporated into a flexible, inner coating, for example silicone
rubber, and indi-
vidually coated with an outer coating, such as a biocompatible layer.
2. The reagent is deposited on the electrodes before the membrane and
substrate are
attached. In this case, thermal bonding is not possible, since the enzyme in
the reagent
would be destroyed. The individual membranes and substrates are first
separated and the
individual sensors are assembled by bonding one membrane with one substrate
using a
suitable bonding agent, for example, cyanoacrylate. As a final step, the
individual sensors
are incorporated into a flexible, inner coating, for example silicone rubber,
and indi-
vidually coated with an outer coating, such as a biocompatible layer. The
sensor can be
inserted into the skin using a needle applicator. The control unit typically
remains outside
the body and can be connected to the sensor element through electrical wires
(leads).
The electrodes are formed on the surface of the substrate. They may be formed
by well known
semiconductor processing techniques, from conductive materials, such as pure
metals or
alloys, or other materials which are metallic conductors. Examples include
aluminum, carbon
(such as graphite), cobalt, copper, gallium, gold, indium, iridium, iron,
lead, magnesium,
mercury (as an amalgam), nickel, niobium, osmium, palladium, platinum,
rhenium, rhodium,
selenium, silicon (such as highly doped polycrystalline silicon), silver,
tantalum, tin, titanium,
tungsten, uranium, vanadium, zinc, zirconium, mixtures thereof, and alloys or
metallic com-
pounds of these elements. Preferably, the electrodes include gold, platinum,
palladium,
iridium, or alloys of these metals, since such noble metals and their alloys
are unreactive in
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biological systems. The electrodes may be any thickness, but preferably are 10
nm to 1 mm,
more preferably, 20 nm to 100 pm, or even 25 nm to 1 pm.
At least two electrodes must be present. The number of electrodes may be 2-
1000, or 3-200, or
even 3-99. Individual electrode sets (2 or 3 electrodes) may be separated into
individual
chambers, each covered with the membrane. Furthermore, individual electrode
sets (2 or 3
electrodes) may each have a different reagent, allowing for an implantable
analvte sensor that
can measure at least two, such as 3-100, or 4-20, different analytes.
The remaining individual part of the implantable analyte sensors are well
known to those of
ordinary skill in the art, and are described, for example, in U.S. Patent
numbers 5,387,327;
5,411,647; and 5,476,776; as well as in PCT International Publication numbers
WO 91/15993;
WO 94/20602; WO 96/06947; and WO 97/ 19344.
Although illustrated with both leads and microelectronic circuitry, these
components are
optional. The microelectronic circuitry may include some or all of the
electrical components
normally external to the implantable analyte sensor, such as a microprocessor,
an amplifier, or
a power supply. If the microelectronic circuitry also includes a transmitter,
or another device
for sending information wirelessly, such as a laser which emits light through
the skin, then
there is no need to include the leads. Alternatively, the microelectronic
circuitry may not be
present, in which case the lead will directly electrically connect the
electrodes with external
electrical components.
Optionally, one or more internal coatings may be present. The internal coating
may function
to regulate diffusion. Examples of internal coatings include cellulose
acetate, polyurethanes,
polyallylamines (PAL), polyaziridine (PAZ), and silicon-containing polymers.
Some specific
examples are described in PCT Publications WO 98/17995, WO 98/13685 and WO
96/06947,
and in U.S. Patent Nos. 4,650,547 and 5,165,407.
Optionally, one or more external coatings may be present. The implantable
analvte sensors of
the present invention are intended to be used in vivo, preferably
subcutaneously in mammals,
such as humans, dogs or mice. The external coatings function to improve the
biocompatibility
of the implantable analyte sensor. Examples of external coatings include
nafion, polyure-
thanes, polytetrafiuoroethylenes (PTFE), poly (ethylene oxide) (PEO), and 2-
methacryloy-
loxyethyl phosphorylcholine-co-n-butyl methacrylate (MPC) membranes. Some
specific
examples are described in PCT Publication WO 96/06947, and in "Medical
Progress through
Technology", Nishida et al. 21: 91-103 (1995).
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The electrodes may be coated with a reagent. The reagent is optional, and may
be used to
provide electrochemical probes for specific analytes. The reagent may be as
simple as a single
enzyme, such as glucose oxidase or glucose hydrogenase for the detection of
glucose. The
enzyme may be immobilized or "wired" as described in PCT Publication WO
96/06947. The
reagents may optionally also include a mediator, to enhance sensitivity of the
sensor. The
starting reagents are the reactants or components of the reagent, and are
often compounded
together in liquid form before application to the electrodes. The liquid may
then evaporate,
leaving the reagent in solid form. The choice of specific reagent depends on
the specific analyte
or analvtes to be measured, and are well known to those of ordinary skill in
the art. For exam-
ple, a reagent for measurement of glucose can contain 62.2 mg polyethylene
oxide (mean
molecular weight of 100-900 kilodaltons)> 3.3 mg NATROSOL 250 M, 41.5 mg
AVICEL RC-
591 F, 89.4 mg monobasic potassium phosphate, 157.9 mg dibasic potassium
phosphate, 437.3
mg potassium ferricyanide, 46.0 mg sodium succinate, 148.0 mg trehalose, 2.6
mg TRITON X-
100 surfactant, and 2,000 to 9,000 units of enzyme activity per gram of
reagent. The enzyme is
prepared as an enzyme solution from 12.5 mg coenzyme PQQ and 1.21 million
units of the
apoenzyme of quinoprotein glucose dehydrogenase, forming a solution of
quinoprotein glu-
cose dehydrogenase. This reagent is described in WO 99/30152, pages 7-10,
hereby incor-
porated by referece.
Other non-limiting examples of enzymes and optional mediators that may be used
in meas-
wring particular analvtes in the present invention are listed below in Table
1.
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T ABLE 1
Analyte Enzymes Mediator Additional Mediator
(Oxidized
Form)
Glucose Glucose Dehydroge-Ferricyanide
nase and Diaphorase
Glucose Glucose-Dehydroge-Ferricyanide
nase
Cholesterol (Quinoprotein) Ferricyanide 2,6-Dimethyl-1,4-
Cholesterol Esterase Benzoquinone
and Cholesterol 2,5-Dichloro-1,4-Benzo-
Oxidase quinone or Phenazine
Ethosulfate
HDL CholesterolCholesterol EsteraseFerricyanide 2,6-Dimethyl-1,4-Benzo-
and Cholesterol quinone
Oxidase
2,5-Dichloro-1,4-Benzo-
quinone or Phenazine
Ethosulfate
TriglyceridesLipoprotein Lipase,Ferricyanide Phenazine Methosulfate
or
Glycerol Kinase, Phenazine
and
Glycerol-3-PhosphateEthosulfate
Oxidase
Lactate Lactate Oxidase Ferricyanide 2,6-Dichloro-1,4-Benzo-
quinone
Lactate Lactate Dehydroge-Ferricyanide
nase and DiaphorasePhenazine
Ethosulfate,
or
Phenazine
Methosulfate
Lactate Diaphorase Ferricyanide Phenazine Ethosulfate,
or
Dehydrogenase Phenazine Methosulfate
Pyruvate Pyruvate Oxidase Ferricyanide
Alcohol AlcoholOxidase Phenylenediamine
Bilirubin Bilirubin Oxidase1-Methoxy-
Phenazine
Methosulfate
Uric Acid Uricase Ferricyanide
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In some of the examples shown in Table 1, at least one additional enzyme is
used as a reaction
catalyst. Also, some of the examples shown in Table 1 may utilize an
additional mediator,
which facilitates electron transfer to the oxidized form of the mediator. The
additional
mediator may be provided to the reagent in lesser amount than the oxidized
form of the
mediator. While the above assays are described, it is appreciated that a
variety of electro-
chemical assays may be conducted in accordance with this disclosure.
Formation of membrane
The following describes how to make a membrane for use in the present
invention, based on
the description from the dissertation of Derek James Hansford, submitted in
partial satis-
faction of the requirements for the degree of Doctor of Philosophy in
Engineering-Materials
Science and Mineral Engineering in the Graduate Division of the University of
California,
Berkeley, submitted in the spring of 1999.
Other membranes, made from other material, may also be used. This specific
method relies
upon a buried nitride etch stop layer.
The buried nitride etch stop layer acts as an etchant stop during the
formation of nanometer
scale pores. The buried nitride etch stop layer facilitates three-dimensional
control of the pore
structure, and facilitates the formation of pores less than 50 nanometers in
diameter.
Moreover, these pores can be uniformly formed across the entire wafer.
Preferably, the first step in the fabrication protocol is to etch a support
ridge structure into a
substrate. The ridges provide mechanical rigidity to the subsequently formed
membrane
structure.
A low stress silicon nitride (LSN or nitride), which operates as an etch stop
layer, is then
deposited on the substrate using low pressure chemical vapor depositions
(LPCVD). In one
embodiment, 0.4 ~m of nitride was used. The resultant structure is shown in
Figure 1.
Figure 1 illustrates a substrate 20 with a nitride etch stop layer 22 formed
thereon.
The base structural layer (base layer) of the membrane is deposited on top of
the stop layer 22.
Since the etch stop layer 22 is thin, the structural layer is deposited down
into the support
ridges formed in the substrate 20. In one embodiment, 5 ~m of polysilicon is
used as the base
layer. Figure 2 illustrates the base layer 24 positioned on the etch stop
layer 22. Low stress
silicon nitride may also be used as the base layer, in which case it operates
as its own etch stop
layer.
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The next processing step is to etch holes in the base layer 24 to define the
shape of the pores.
Masks, such as those used in traditional semiconductor processing, may be used
to define the
pores. For example, the holes may be etched through the polysilicon by
chlorine plasma, with
a thermally grown oxide layer used as a mask.. In this step, it is important
to make sure the
5 etching goes completely through the base layer 24, so a 10-15% overetch is
preferably used. It
is useful to note that the buried nitride etch stop 22 acts as an etch stop
for the plasma etching
of a silicon base layer 24. Otherwise, if the plasma punched through the
nitride, tighter control
of the etch step would have to be exercised to prevent the complete removal of
the nitride
under the plug layer (to prevent removal in the final KOH etch). Figure 3
illustrates the result
10 of this processing. In particular, the figure illustrates holes 26 formed
in the base layer 24, but
terminating in the nitride etch stop layer 22.
Pore sacrificial oxide is subsequently grown on the base layer 24. Figure 4
illustrates a sacri-
ficial oxide 28 positioned on the base layer 24.
The sacrificial oxide thickness determines the pore size in the final
membrane, so control of
this step is critical to reproducible membranes. This is accomplished by the
thermal oxidation
of the base layer 24 (e.g., a growth temperature of between 850-950°C
for approximately one
hour with a ten minute anneal). Naturally, many techniques may be used to form
a controlled
thickness sacrificial layer. For example, a thermally evaporated tungsten film
may be used as a
sacrificial layer for polymer membranes and selectively removed with hydrogen
peroxide. The
basic requirement of the sacrificial layer is the ability to control the
thickness with high pre-
cision across the entire wafer. Thermal oxidation of both polysilicon and
nitride allows the
control of the sacrificial layer thickness of less than 5% across the entire
wafer. Limitations on
this control arise from local inhomogeneities in the base layer, such as the
initial thickness of
the native oxide (especially for polysilicon) the grain size or the density,
and the impurity con-
centrations.
To mechanically connect the base layer 24 with the plug layer (necessary to
maintain the pore
spacing between layers), anchor points were defined in the sacrificial oxide
layer 26. In the
present design, this is accomplished by using the same mask shifted from the
pore holes by
1 ~m diagonally. This produced anchors in one or two corners of each pore
hole, which pro-
vides the desired mechanical connection between the structural layers while
opening the pore
area as much as possible. Figure 5 illustrates anchors 30 formed via this
process.
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A plug structural layer is subsequently deposited to fill in the holes 26.
This step has been
implemented by depositing 1.5 ~m of polysilicon. The resultant plug layer 32
is shown in
Figure 6.
To open the pores at the surface, the plug layer 32 is planarized down to the
base layer, leaving
the final structure with the plug layer only in the pore hole openings, as
shown in Figure 7.
The method of planarization depends on the material used as the plug material.
For the hard
micro-fabrication materials (polysilicon and nitride), chemical mechanical
polishing was used
for planarization. The other materials studied were roughly planarized using a
plasma etch,
with a quick wet chemical smoothing. This technique has the advantage that,
assuming it is
not etched by the plasma used, the base layer is not affected, but has the
disadvantage of the
need for controlled etch timing to avoid completely etching the plugs
themselves.
At this point, the membrane is ready for release, so a protective layer 34 is
deposited on the
wafer (completely covering both sides of the wafer). The requirements of the
protective layer
34 are that it be impervious to the silicon etch (KOH for these studies) and
that it be removed
without removing the plug 32 or base 24 structural layers. For polysilicon and
nitride struc-
tural layers, a thin nitride layer is used as the protective layer (nitride is
not etched at all by
KOH and dissolves slowly in HF). For polymeric structural materials, silicon
is used as a pro-
tective layer, due to the processing temperature necessary for nitride
deposition (835° C).
The backside etch windows were etched in the protective layer, exposing the
silicon in desired
areas, and then the entire structure was placed in an 80°C KOH bath
until the silicon wafer
substrate 20 is etched up to the membrane base layer 24 (as evidenced by the
smooth buried
etch stop layer). Figure 8 illustrates the resultant aperture 36 formed in the
substrate 20.
At this point, the buried nitride layer 22, the sacrificial oxide layer 34,
and plug layer 32 are
removed by etching in HF or SF~/oxygen plasma. The resultant membrane 4 with
nanometer
scale pores is shown in Figure 9.
Characterization of membranes
The purpose of the membranes is to allow the analyte of interest (such as
glucose) to diffuse
through the membrane, while excluding large molecules (such as proteins).
Therefore, two
important characteristics of the membranes are glucose diffusion and albumin
diffusion. All
tests are carried out at room temperature (25°C).
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The following is a glucose diffusion test:
Diffusion of glucose is measured using a mini diffusion chamber constructed
around the
membranes. The diffusion chamber, fabricated out of acrylic, consists of two
compartments A
and B with fixed volumes of 2 ml, separated by the desired membrane, sealed
with o-rings,
and screwed together.
Glucose is measured on either side of the membrane using the diffusion chamber
by means of
a quantitative enzymatic assay (TRINDERT~f, SIGMA) and colorometric reading
via a spec-
trophotometer. Starting glucose concentrations for all tests were 6,666 mg/dl
and 0.0 mg/dl in
chambers A and B, respectively. Samples of 0.1 ml are taken from the diffusion
chamber and
10 p1 of that are added to 3 ml of glucose reagent in a cuvette, and mixed
gently by inversion.
Each tube is incubated for 18 minutes at room temperature and then readings
are taken at a
wavelength of 505 nm. The reagent is linear up to 750 mg/dl. The diffusion
chamber itself is
attached to a motor for stirring in order to minimize boundary layer effects
(diffusion resis-
tance at the liquid/membrane interface). In order to ensure wetting of the
pores, the receptor
cell is first filled with phosphate buffer saline (PBS) for fifteen minutes
before the filling of the
donor cell. The donor cell is filled with solutions of glucose in PBS in
varying concentrations.
The following is an albumin diffusion test:
Albumin is also measured on either side of the membrane using the same
diffusion chamber
as in the glucose diffusion test. Albumin diffusion and/or exclusion is first
measured and
quantified using Albumin I3CP (bromocresol purple, SIGMA). Starting albumin
concentra-
tions for all tests are 4 g/dl and 0.0 mg/dI in chambers A and B,
respectively. A sample of 0.1
ml is taken at time zero and at the end of the diffusion period (time = 330
minutes). An
aliquot of 300 u1 is then added to 3 ml of the reagent and absorbence is read
at 600 nm. Rea-
gent plus deionized water is used as the blank. The BCP assay is linear up to
6g/dl but is not
accurate below 1 g/dl. For the small concentration of albumin that might be
present in cham-
ber A, the presence of any protein in chamber B is measured using the Bradford
Method
(MICRO PROTEIN KIT, SIGMA). This method quantitates the binding of Coomassie
bril-
liant blue to an unknown protein and compares this binding to that of
different amounts of a
standard protein. Albumin is used as a standard protein. This method
quantifies 1 to 100
micrograms protein using a standard curve, with sensitivity down to 10 mg/dl
or 0.1 g/dl
protein. The absorbance is measured at 595 nm.
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Analysis of membranes
Diffusion of glucose was measured for three types of membranes: silicon
micromachined
membranes (average pore size = 0.0245 microns), WHATMAN ANODISC membranes
(aver-
age pore size = .02 microns), and MF-MILLIPORE mixed cellulose acetate and
nitrate mem-
brane (average pore size = 0.025 microns).
The results from the albumin test are shown in the table below.
WHATMAN MILLIPORE silicon (micro-
machined)
time albumin conc. albumin conc. albumin conc.
(g/dL) (g/dL) (g/dL)
0 0 0 0
420 MinØ25'd0.05 0.2'd0.01 OH0.001
The presence of albumin does not seem to impede passage of glucose through the
membranes,
nor slow down glucose transport. No detectable amounts of albumin diffuse
through the
micromachined membrane. The same membrane, however, shows glucose diffusion.
The
micromachined membranes are able to achieve complete exclusion of albumin (to
within the
limits of detection), while allowing glucose diffusion. Comparing diffusion
rates with that of
commercially available membranes, the micromachined membranes have glucose
diffusion
properties comparable to 1\IILLIPORE and alumina WHATMAN membranes with
similar
pore sizes.
The passage of albumin through the micromachined membrane is measured by
looking at the
change of albumin concentration in chamber A and chamber B over time. Using
the BCP
assay, there are no detectable traces of albumin in chamber B. However, the
amount of albu-
min in chamber B may have been below the limits of detectability of this assay
system. There-
fore, the Bradford Method was also employed. Using this microassay, again no
detectable
amounts of albumin were found in chamber B for the micromachined membrane, but
small
amounts of protein were found in chamber B using both the MILLIPORE and
WHATMAN
membranes. The amounts of albumin detected after 420 minutes in chamber B were
approximately 0.25 g/dI and 0.20 g/dI albumin for the MILLIPORE and WHATMAN
mem-
branes, respectively.
CA 02406814 2002-09-13
WO 01/69222 PCT/EPOI/03026
14
Glucose does diffuse through micromachined membranes at a rate comparable to
commer-
cially available membranes. At the same time, albumin is excluded from
passage. In mixed
solutions of glucose and albumin, only glucose diffuses through the
micromachined mem-
branes.
