U.S. patent application number 15/189372 was filed with the patent office on 2016-12-22 for sensors for long-term and continuous monitoring of biochemicals.
This patent application is currently assigned to THE TRUSTEES OF COLUMBIA UNIVERSITY IN THE CITY OF NEW YORK. The applicant listed for this patent is THE TRUSTEES OF COLUMBIA UNIVERSITY IN THE CITY OF NEW YORK. Invention is credited to Xian Huang, Qiao Lin.
Application Number | 20160370362 15/189372 |
Document ID | / |
Family ID | 43011383 |
Filed Date | 2016-12-22 |
United States Patent
Application |
20160370362 |
Kind Code |
A1 |
Lin; Qiao ; et al. |
December 22, 2016 |
SENSORS FOR LONG-TERM AND CONTINUOUS MONITORING OF BIOCHEMICALS
Abstract
The disclosed subject matter relates to a sensor or system for
monitoring a target analyte by using a polymer solution that is
capable of binding to the analyte. The sensor of the disclosed
subject matter includes a viscosity-based sensor or a
permittivity-based sensor. The viscosity-based sensor contains a
semi-permeable membrane, a substrate, and a microchamber including
a vibrational element. The permittivity-based sensor contains a
semi-permeable membrane, a substrate, and a microchamber. The
sensor discussed herein provides excellent reversibility and
stability as highly desired for long-term analyte monitoring.
Inventors: |
Lin; Qiao; (New York,
NY) ; Huang; Xian; (New York, NY) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
THE TRUSTEES OF COLUMBIA UNIVERSITY IN THE CITY OF NEW
YORK |
New York |
NY |
US |
|
|
Assignee: |
THE TRUSTEES OF COLUMBIA UNIVERSITY
IN THE CITY OF NEW YORK
New York
NY
|
Family ID: |
43011383 |
Appl. No.: |
15/189372 |
Filed: |
June 22, 2016 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
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13246404 |
Sep 27, 2011 |
9400233 |
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15189372 |
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PCT/US09/62891 |
Oct 30, 2009 |
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13246404 |
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61171338 |
Apr 21, 2009 |
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61225475 |
Jul 14, 2009 |
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61225473 |
Jul 14, 2009 |
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Current U.S.
Class: |
1/1 |
Current CPC
Class: |
G01N 33/54366 20130101;
G01N 27/3275 20130101; G01N 1/16 20130101; G01N 27/413 20130101;
G01N 27/40 20130101; G01N 33/66 20130101 |
International
Class: |
G01N 33/543 20060101
G01N033/543; G01N 27/413 20060101 G01N027/413; G01N 27/327 20060101
G01N027/327; G01N 33/66 20060101 G01N033/66; G01N 27/40 20060101
G01N027/40 |
Goverment Interests
STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH OR DEVELOPMENT
[0002] This invention was made with government support under NIH
grant number DK63068-05 awarded by National Institutes of Health,
and NSF grant number ECCS-0702101 awarded by National Science
Foundation. The United States Government has certain rights in the
invention.
Claims
1. A sensor for monitoring a target analyte by using a polymer
solution that is capable of binding to the analyte, comprising (a)
a semi-permeable membrane comprising a material permeable to the
analyte; (b) a substrate; and (c) a microchamber, formed between
the semi-permeable membrane and the substrate, and adapted to
receive the polymer solution, such that when the analyte is placed
on the semi-permeable membrane, at least a portion of the analyte
will permeate the semi-permeable membrane and bind to at least a
portion of the polymer solution to thereby cause a change in
permittivity of the polymer solution.
2. The sensor of claim 1, further comprising a top electrode and a
bottom electrode, each embedded to the microchamber.
3. The sensor of claim 1, wherein the microchamber comprises a gap
between the top electrode and the bottom electrode.
4. The sensor of claim 1, further comprising a detector, coupled to
the microchamber, for detecting a permittivity change, if any,
caused by binding between the analyte and the polymer.
5. The sensor of claim 4, wherein the detector comprises a
capacitive detector.
6. The sensor of claim 1, wherein the polymer comprises a polymer
that reversely binds to the analyte.
7. The sensor of claim 1, wherein the analyte comprises
glucose.
8. The sensor of claim 1, wherein the polymer comprises a plurality
of boronic acid moieties.
9. The sensor of claim 8, wherein the polymer comprises
poly(acrylamide-ran-3-acrylamidophenylboronic acid).
10. The sensor of claim 1, further comprising a temperature
sensor.
11. The sensor of claim 1, wherein the permittivity sensor
comprises a first thin-film electrode coated on the top chamber
wall and a second thin-film electrode coated on the bottom chamber
wall.
12. The sensor of claim 1, wherein the first thin-film electrode
and the second thin-film electrode comprise at least one of copper
and gold.
13. The sensor of claim 1, wherein each of the top chamber wall and
the bottom chamber wall comprise glass coverslips.
14. The sensor of claim 1, wherein the introduction channel
comprises a tube.
Description
CROSS REFERENCE TO RELATED APPLICATIONS
[0001] This application is a divisional application of U.S. patent
application Ser. No. 13/246,404, filed Sep. 27, 2011, which is a
continuation of International Application PCT/US09/062891, filed
Oct. 30, 2009, and which claims priority to U.S. Provisional
Application No. 61/171,338, filed on Apr. 21, 2009; 61/225,473,
filed Jul. 14, 2009; and 61/225,475, filed Jul. 14, 2009, the
disclosures of which are hereby incorporated by reference in their
entireties and from which priority is claimed.
BACKGROUND
[0003] Diabetes mellitus is a metabolic disease characterized by
persistent hyperglycemia (high blood sugar levels). Close
monitoring of daily physiological glucose levels reduces the risk
of complications caused by conditions such as hypoglycemia or
hyperglycemia. This can be achieved by continuous glucose
monitoring (CGM) systems, which involve either non-invasive or
minimally invasive detection of glucose. Currently, subcutaneously
implanted enzymatic electrochemical detection is a prevailing CGM
technique, and is the basis for a number of commercially available
sensors. These FDA approved commercial products detect glucose by
enzyme-catalyzed reactions.
[0004] Electrochemical methods are sensitive and specific for
glucose detection, but suffer from drawbacks. Firstly, the
irreversible consumption of glucose in electrochemical detection
induces a potential change in the equilibrium glucose concentration
in the tissue, and thus, affects the actual measured glucose level.
Furthermore, the rate of glucose consumption can be diffusion
limited. Any changes in diffusion layers due to biofouling (e.g.,
by protein adsorption, cell deposition, and capsule formation) on
the sensor surface can affect the diffusion rate, and, thus, the
sensor sensitivity. In addition, drift from hydrogen peroxide
production and interference from electrode-active chemicals can
cause erosion of the sensor electrodes and deactivation of
functional enzymes, compromising the sensor accuracy, reliability
and longevity. As a result, electrochemical CGM sensors can exhibit
large drifts over time, and require frequent calibration by finger
pricks. This lack of reliability has been severely hindering CGM
applications to practical diabetes management.
[0005] To overcome the drawbacks of electrochemical detection,
alternative glucose sensing techniques have been investigated.
Methods that use non-consumptive, competitive affinity binding of
glucose have been considered. One technique exploits the solution
of a polysaccharide (e.g., dextran) crosslinked by a
glucose-binding protein (e.g., concanavalin A, or Con A): glucose
binds competitively to Con A and causes reversible de-crosslinking
of the dextran-Con A complex, which can be detected via the
resulting changes in solution properties, such as fluorescence or
viscosity. As affinity sensing is based on equilibrium binding in
which glucose is not consumed, it is not susceptible to
electroactive interferents. Also, affinity sensing is considerably
more tolerant to biofouling. That is, the deposition of biological
material (e.g., cells and proteins) on the implanted affinity
sensor surface results only in an increased equilibration time
without any changes in measurement accuracy. Consequently, affinity
glucose sensors can be highly stable and low-drift.
[0006] Unfortunately, Con A is immunogenic and cytotoxic and
degrades with time. Although certain alternatives, such as ones
utilizing Microelectromechanical Systems (MEMS) technology have
been developed, they can suffer from the same or different
limitations associated with Con A, e.g., limited mechanical
reliability, poor reversibility, and significant drifts. Thus,
there remains a need in the art for a sensor for stable and
potentially implantable MEMS-based continuous glucose sensing.
SUMMARY
[0007] The presently disclosed subject matter provides a sensor for
monitoring a target analyte by using a polymer solution that is
capable of binding to the analyte, which sensor includes a
semi-permeable membrane which includes a material permeable to the
analyte, a substrate, and a microchamber including a vibrational
element therein, formed between the semi-permeable membrane and the
substrate, and adapted to receive the polymer solution, such that
when the analyte is placed on the semi-permeable membrane, at least
a portion of the analyte will permeate the semi-permeable membrane
and bind to at least a portion of the polymer solution to thereby
cause a change in vibration of the vibrational element. In one
embodiment, the change in vibration of the vibrational element of
the sensor is caused by a change in viscosity of the polymer
solution. In certain embodiments, the vibrational element includes
a vibrational cantilever or a vibrational diaphragm. In other
embodiments, the sensor further includes a detector which is
coupled to the microchamber for detecting a viscosity change, if
any, caused by binding between the analyte and the polymer. The
detector includes an optical lever or a capacitive detector.
[0008] In the disclosed subject matter, the polymer can reversely
bind to the analyte. In certain embodiments, the analyte can be
glucose. The polymer solution can include a polymer having a
plurality of boronic acid moieties. In certain embodiments, the
polymer includes poly(acrylamide-ran-3-acrylamidophenylboronic
acid) (PAA-ran-PAAPBA).
[0009] The vibrational element of the disclosed subject matter can
include Parylene. In certain embodiments, when the vibrational
element is a vibrational diaphragm, the sensor further includes a
top electrode embedded in the vibrational diaphragm within the
microchamber to thereby form a capacitor with a bottom electrode on
the substrate. In certain embodiments, the capacitor is adapted to
sense a change in the capacitance between the top electrode and the
bottom electrode caused by binding between the analyte and the
polymer. In other embodiments, the sensor can further include at
least one integrated permalloy film. In certain embodiments, the
permalloy film further includes a Parylene layer for
passivation.
[0010] The presently disclosed subject matter also provides a
sensor for monitoring a target analyte by using a polymer solution
that is capable of binding to the analyte, which sensor includes a
semi-permeable membrane having a material permeable to the analyte,
a substrate, and a microchamber which is formed between the
semi-permeable membrane and the substrate, and adapted to receive
the polymer solution, such that when the analyte is placed on the
semi-permeable membrane, at least a portion of the analyte will
permeate the semi-permeable membrane and bind to at least a portion
of the polymer solution to thereby cause a change in permittivity
of the polymer solution.
[0011] In certain embodiments, the sensor further includes a top
electrode and a bottom electrode, each coupled to the microchamber.
In certain embodiments, the microchamber is formed from the gap
between the top electrode and the bottom electrode. In other
embodiments, the sensor further includes a detector which is
coupled to the microchamber, for detecting a permittivity change,
if any, caused by binding between the analyte and the polymer. The
detector includes a capacitance-voltage transformation circuit. In
other embodiments, the sensor further includes a temperature
sensor.
BRIEF DESCRIPTION OF THE DRAWINGS
[0012] FIG. 1 is a diagram illustrating a MEMS affinity CGM sensor
design in accordance with the disclosed subject matter.
[0013] FIGS. 2A-2D are diagrams illustrating a fabrication process
in accordance with the disclosed subject matter: FIG. 2A etching
small anchoring cavities in silicon and depositing a Parylene
layer; FIG. 2B electroplating permalloy and passivating it with
Parylene; FIG. 2C patterning the Parylene and permalloy layers to
define a cantilever; and FIG. 2D etching silicon and SiO.sub.2 to
release the cantilever.
[0014] FIGS. 3A-3B illustrate an exemplary fabricated MEMS CGM
sensor: FIG. 3A before, and FIG. 3B after packaging.
[0015] FIG. 4 is a chart illustrating variation with time of the
cantilever vibration amplitude at 28 Hz upon introduction of a 108
mg/dL glucose solution to the sensor. Glucose was initially absent
from the solution of polymer in the microchamber.
[0016] FIG. 5 is a chart illustrating frequency dependent amplitude
of the cantilever vibration obtained after glucose permeation and
binding had reached equilibrium (squares), as compared to that from
a comparison experiment (diamonds) in which glucose concentrations
inside and outside the microchamber were pre-equilibrated.
[0017] FIG. 6 is a chart illustrating frequency dependent amplitude
of the cantilever vibration at physiologically relevant glucose
concentrations.
[0018] FIG. 7 is a chart illustrating frequency dependent phase lag
of the cantilever vibration at physiologically relevant glucose
concentrations.
[0019] FIG. 8 is a chart illustrating frequency-dependent amplitude
of the cantilever vibration in response to a sequence of glucose
concentrations.
[0020] FIG. 9 is a diagram illustrating an experimental setup for
characterization of a MEMS glucose sensor in accordance with the
disclosed subject matter.
[0021] FIG. 10 is a chart illustrating reversibility of the MEMS
sensor to glucose concentration changes. The noise shown reflects
environmental disturbances to the optical setup.
[0022] FIG. 11 is a chart illustrating evaluation of drift for the
cantilever-based MEMS glucose sensor.
[0023] FIGS. 12A-12B are images of a MEMS affinity glucose sensor
FIG. 12A before and FIG. 12B after packaging and putting into the
measurement system.
[0024] FIG. 13 is a chart illustrating viscosity responses of
polymers of different molecular weight and PAA-ran-PBA
percentage.
[0025] FIG. 14 is a chart illustrating reversibility of the
copolymer (4.7%) to glucose concentration changes.
[0026] FIGS. 15A-15B are diagrams illustrating an exemplary glucose
sensor design in accordance with the disclosed subject matter and
an image of a fabricated sensor.
[0027] FIG. 16 is a diagram illustrating an experiment setup in
accordance with the disclosed subject matter.
[0028] FIG. 17 is a chart illustrating sensor response (at 10 kHz)
to polymer solutions at various concentrations of glucose and
fructose (an unspecific analyte).
[0029] FIG. 18 is a chart illustrating frequency-dependent
capacitance changes of the polymer at various glucose
concentrations.
[0030] FIG. 19 is a chart illustrating sensor response to pure
glucose solutions (free of polymer) at varying glucose
concentrations.
[0031] FIG. 20 is a chart illustrating drift of the sensor output
(at 10 kHz) over a period of about 10 hours.
[0032] FIGS. 21A-21B are schematics of the MEMS capacitive glucose
sensor: FIG. 21A side-view of the capacitive glucose sensor; FIG.
21B top-view of the capacitive glucose sensor (Dimensions are given
in .mu.m).
[0033] FIGS. 22A-22F are a diagram illustrating an exemplary
fabrication process: FIG. 22A bottom gold electrode deposition and
sacrificial layer patterning; FIG. 22B Parylene deposition and top
gold electrode deposition; FIG. 22C permalloy electroplating and
additional Parylene layer deposition; FIG. 22D photoresist etching
holes patterning; FIG. 22E sacrificial layer removal and diaphragm
releasing; FIG. 22F membrane bonding and sensor packaging.
[0034] FIG. 23 is a schematic of the sensor outfitted with a flow
cell containing glucose solution.
[0035] FIG. 24 is a diagram illustrating exemplary capacitive
measurement circuitry.
[0036] FIG. 25 is a diagram illustrating an experimental setup for
characterization of the MEMS glucose sensor.
[0037] FIGS. 26A-26B are a chart illustrating frequency-dependent
behavior of the harmonically driven vibration of the sensor
diaphragm at physiologically relevant glucose concentrations: FIG.
26A amplitude, and FIG. 26B phase shift.
[0038] FIG. 27 is a chart illustrating 1-DOF mass-spring-damper
model fitted to the experimental data obtained at a glucose
concentration of 90 mg/dL.
[0039] FIG. 28 is a chart illustrating damping ratio obtained by
fitting the 1-DOF mass-spring-damper model to the experimental data
at varying glucose concentrations.
[0040] FIG. 29 is a chart illustrating time course of the diaphragm
vibration amplitude at 1000 Hz as the sensor responded to glucose
concentration changes from 90 to 120 mg/dL, which was then reversed
to 90 mg/dL.
[0041] FIG. 30 is a chart illustrating diaphragm vibration
amplitude a5 1000 Hz over an extended time duration as the glucose
concentration was held constant at 90 mg/dL.
[0042] FIGS. 31A-31C are diagrams illustrating a biocompatible,
glucose-sensitive polymer, PAA-ran-PAAPBA. FIG. 31A The polymer
composition and mechanism of interaction with glucose. FIG. 31B
Glucose-induced viscosity change of a 1.9% PAA-ran-PAAPBA solution
in PBS buffer (pH 7.4). FIG. 31CA Glucose-induced viscosity change
of a 5% PAA-ran-PAAPBA solution in PBS buffer (pH 7.4).
[0043] FIGS. 32A-C are diagrams of a MEMS device for demonstration
of dielectric affinity biosensing: FIG. 32A design schematic; and
images of a fabricated device FIG. 32B before, and FIG. 32C after
packaging.
[0044] FIGS. 33A-B depict an experimental setup FIG. 33A. The
equivalent capacitance of the device is measured by a
transformation circuit. FIG. 33B The output voltage of the
capacitance/voltage transformation circuit is proportional to the
sensor admittance.
[0045] FIGS. 34A-34B depict frequency dependence of the equivalent
device capacitance (PAAPBA content of the PAA-ran-PAAPBA polymer:
5%). FIG. 34A Equivalent capacitance when the polymer solution is
free of glucose. FIG. 34B Changes in equivalent capacitance at
physiologically relevant glucose concentrations with respect to the
glucose-free case.
[0046] FIG. 35 shows the equivalent capacitance of a device at 100
kHz plotted as a function of glucose concentration for different
polymer compositions.
[0047] FIG. 36 shows equivalent capacitance at 100 kHz plotted as a
function of concentrations of glucose, and the unspecific
monosaccharides galactose and fructose.
DETAILED DESCRIPTION
[0048] The disclosed subject matter addresses the need in the art
for a stable and implantable sensor that is capable of monitoring
target analytes. More specifically, the disclosed subject matter
provides for a sensor and system that can be used for continuous
analyte monitoring. In certain embodiments, the sensor is used for
continuous glucose monitoring (CGM).
[0049] The sensor of the presently disclosed subject matter
includes a sensor that can be a viscosity-based sensor or a
permittivity-based sensor. The viscosity-based sensor includes a
semi-permeable membrane permeable to a target analyte, a substrate,
and a microchamber formed between the membrane and substrate and is
adapted to receive a polymer solution. The microchamber includes a
vibrational element, which can be a vibrational cantilever or a
diaphragm-based sensor. In the viscosity-based sensor, the binding
between the analyte of interest and the polymer can produce a
viscosity change of the polymer which can be detected and
measured.
[0050] The permittivity-based sensor includes a semi-permeable
membrane permeable to a target analyte, a substrate, and a
microchamber formed between the membrane and the substrate and is
adapted to receive a polymer solution. In the permittivity-based
sensor, the binding between the analyte of interest and the polymer
can produce a change in the permittivity of the polymer which can
be detected and measured. The sensor can then be used to monitor
and detect target analytes of interest in a stable and accurate
manner. Such methods can be useful in patient monitoring,
diagnosis, and treatment.
[0051] Unless otherwise defined, all technical and scientific terms
used herein have the same meanings as commonly understood by one of
ordinary skill in the art to which the disclosed subject matter
belongs. Although methods and materials similar or equivalent to
those described herein can be used in its practice, suitable
methods and materials are described below.
[0052] It is to be noted that the term "a" entity or "an" entity
refers to one or more of that entity. As such, the terms "a", "an",
"one or more", and "at least one" can be used interchangeably
herein. The terms "comprising," "including," and "having" can also
be used interchangeably. In addition, the terms "amount" and
"level" are also interchangeable and can be used to describe a
concentration or a specific quantity. Furthermore, the term
"selected from the group consisting of" refers to one or more
members of the group in the list that follows, including mixtures
(i.e. combinations) of two or more members.
[0053] The term "about" or "approximately" means within an
acceptable error range for the particular value as determined by
one of ordinary skill in the art, which will depend in part on how
the value is measured or determined, i.e., the limitations of the
measurement system. For example, "about" can mean within 3 or more
than 3 standard deviations, per the practice in the art.
Alternatively, "about" can mean a range of up to +/-20%, up to
+/-10%, up to +/-5%, or alternatively up to +/-1% of a given value.
Alternatively, with respect to biological systems or processes, the
term can mean within an order of magnitude, preferably within
5-fold, and more preferably within 2-fold, of a value.
[0054] As used herein, the term "analyte" is a broad term and is
used in its ordinary sense and includes, without limitation, any
chemical species the presence or concentration of which is sought
in material sample by the sensors and systems disclosed herein. For
example, the analyte(s) include, but not are limited to, glucose,
ethanol, insulin, water, carbon dioxide, blood oxygen, cholesterol,
bilirubin, ketones, fatty acids, lipoproteins, albumin, urea,
creatinine, white blood cells, red blood cells, hemoglobin,
oxygenated hemoglobin, carboxyhemoglobin, organic molecules,
inorganic molecules, pharmaceuticals, cytochrome, various proteins
and chromophores, microcalcifications, electrolytes, sodium,
potassium, chloride, bicarbonate, and hormones. In one embodiment,
the analyte is glucose. In various embodiments, the analytes can be
other metabolites, such as lactate, fatty acids, cysteines and
homocysteines.
[0055] As used herein, "response time" or "time constant"
represents time consumption for analyte permeation into the
microchamber and equilibrium binding with the polymer. Response
time is time for a sensor to complete one measurement cycle. The
response time of current commercial CGM products is between about 5
and about 15 minutes. In various embodiments of the disclosed
subject matter, the response times of the sensor are less than
about 5 minutes, alternatively less than about 3 minutes, and
alternatively less tan about 1.5 minutes. In certain embodiments,
the response times are about 1.5 minutes or about 3 minutes.
[0056] As used herein, the term "vibrational element" refers to a
mechanical moving part, which is capable of vibrating. The
vibrational element as used in presently disclosed subject matter
includes, but is not limited to, a vibrational cantilever or a
vibrational diaphragm.
The Sensor
[0057] The disclosed subject matter provides a sensor for
monitoring a target analyte by using a polymer solution that is
capable of binding to the analyte. The sensor includes a
semi-permeable membrane that is made of materials permeable to the
analyte, a substrate, and a microchamber. The substrate can be
comprised of known applicable materials in the art including but
not limited to silicone. The microchamber is formed between the
semi-permeable membrane and the substrate. The analyte can permeate
the membrane and bind to at least a portion of the polymer
solution.
[0058] The sensor can be either a viscosity-based sensor or can be
a permittivity-based sensor. When the sensor is a viscosity-based
sensor, the microchamber can include either a vibrational
cantilever or a vibrational diaphragm.
Viscosity-Based Sensor
Vibrational Cantilever
[0059] In certain embodiments, the sensor is a viscosity-based
sensor that includes a vibrational cantilever, which is also
referred as "cantilever-based sensor." In these embodiments, the
cantilever can be made of any appropriate polymer known in the art,
including but not limited to the polymer Parylene. The cantilever
is situated inside a microchamber as shown in FIGS. 1 and 3A. The
microchamber is formed between a cavity etched into the substrate
and a semi-permeable membrane, and is filled with a solution of
polymer, for example, PAA-ran-PAAPBA. The semi-permeable membrane
is made of material capable of being permeable to the target
analyte. Such materials include but are not limited to cellulose
acetate (CA).
[0060] The Parylene cantilever is anchored onto the substrate at
one end, with its free end coated with a permalloy thin film, which
is protected by an additional Parylene layer to prevent the
permalloy from corrosion by the polymer solution. The environmental
analyte can permeate through the semi-permeable membrane and bind
with the polymer inside the microchamber. The semi-permeable
membrane prevents the polymer from escaping, while allowing the
analyte to diffuse into and out of the microchamber. While not part
of the sensor, a test cell can be incorporated on the other side of
the membrane for introduction of an analyte solution for sensor
characterization. The cross-linking between the analyte and the
polymer can increase the viscosity of the polymer in the
microchamber as well as the damping of the cantilever vibration,
causing a decrease in cantilever vibration amplitude and a shift in
vibration phase which can be and detected and measured, for
example, by using an optical lever technique.
[0061] As shown in FIG. 1, a cantilever-based sensor 1 includes a
semi-permeable membrane 2 which is permeable to a sample analyte 3,
a substrate 4, and a microchamber 5 which includes a Parylene
cantilever 6. The cantilever 6 has a fixed end 7 and a free end 8.
The cantilever 6 is anchored to the microchamber 5 at its fixed end
7, and its free end 8 is coated with a permalloy film 9. The
permalloy film 9 is passivated by a Parylene layer. The
microchamber 5 is filled with a solution of a biocompatible polymer
10 that binds with the analyte 3. The polymer includes, but is not
limited to a polymer having a plurality of boronic acid moieties,
for example, PAA-ran-PAAPBA. The semi-permeable membrane 2 which
prevents the polymer 10 from escaping, while allowing the analyte 3
to diffuse into and out of the microchamber 5.
[0062] The sensor is placed in two mutually orthogonal magnetic
fields. When the sensor is oriented horizontally, these include a
vertical electromagnetic (EM) field generated by a solenoid and a
horizontal magnetic field from a permanent magnet (PM). The PM
magnetizes the permalloy thin film, exciting a magnetic field in
the permalloy film along the cantilever length. A torque thus is
generated on this magnetized permalloy film attempting to align the
cantilever with the EM field. This torque is distributed along the
length of the cantilever, with a magnitude proportional to the
product of the permalloy volume, the EM field intensity, and the
magnetization of the permalloy, and causes the cantilever to bend.
Thus, a time-dependent EM field produces a time-dependent torque,
resulting in vibration of the cantilever. In addition, the
vibration-induced flow of the polymer solution in general imparts
hydrodynamic inertia and damping on the cantilever. Because of the
direct dependency of the flow-structure interaction on the
viscosity of the polymer, the viscosity of the polymer can be
obtained by measuring the vibration of the cantilever, allowing
monitoring the presence and property of the analyte, for example,
the concentration of the analyte.
[0063] The response time or time constant of the diffusion can be
measured using methods known in the art, including but not limited
to an estimate based on consideration of the diffusion of the
analyte into the sensor. In one embodiment, the timescale for the
analytes to diffuse through the semi-permeable membrane and the
microchamber can be estimated to be
t.sub.diff.about.(d.sub.m.sup.2/.lamda.+d.sub.f.sup.2)/D.sub.g
(1)
where D.sub.g is the diffusivity of the analyte in the polymer
solution, d.sub.m and .lamda. are respectively the thickness and
porosity of the semi-permeable membrane, and d.sub.f is the
effective height of the microchamber accounting for the deflection
of the membrane caused by sample loading in the test cell. In one
embodiment, the analyte is glucose, and the polymer is
PAA-ran-PAAPBA. It is estimated that D.sub.g is on the order of
3.times.10.sup.-11 m/s.sup.2 according to glucose diffusivity in
water (7.1.times.10.sup.-10 m/s.sup.2) scaled by the ratio of water
viscosity to the viscosity of the PAA-ran-PAAPBA solution at
relevant glucose concentrations. The estimator d.sub.m is
approximately 20 .mu.m, and .lamda. is approximately 0.6
("Cellulose Acetate Membrane Filters,"
http://www.advantecmfs.com/filtration/membranes/mb_ca.shtml), and
d.sub.f is approximately 100 .mu.m. The estimated diffusion
timescale can be approximately 6 minutes.
[0064] FIG. 2 illustrates an exemplary manufacturing process for
the cantilever-based sensor. A Parylene layer (5 .mu.m) 11 is
deposited by chemical vapor deposition onto a SiO.sub.2-coated
silicon wafer 12 into which small cavities are etched (see FIG.
2A). The small cavities allow the Parylene to be anchored over an
increased surface area for improved adhesion. A 100 nm copper seed
layer is then deposited on the Parylene, followed by the deposition
and patterning a 1.5 .mu.m photoresist layer defining the permalloy
deposition area (150 .mu.m long and 200 .mu.m wide). A permalloy
thin film 13 is then deposited by electroplating, followed by the
removal of the photoresist and unused copper, and the deposition of
a second Parylene layer for passivation, therefore is referred as
Parylene passivation layer 14 (2 .mu.m) (see FIG. 2B). The two
Parylene layers 11 and 14, along with the underlying SiO.sub.2
layer 12, are then patterned to define a cantilever (250 .mu.m in
both length and width) (see FIG. 2C). The cantilever is finally
released by gas-phase XeF.sub.2 etching of the silicon underneath
(forming a cavity approximately 500 .mu.m.times.500 .mu.m.times.250
.mu.m in dimensions) and removal of the SiO.sub.2 directly beneath
the cantilever (see FIG. 2D). Following wafer dicing, a chip bonds
to a poly(dimethylsiloxane) (PDMS) sheet in which a hole is
fabricated by replica molding to define a microchamber with inlet
and outlet channels (approximate 30 .mu.L in dimensions), which is
in turn bonded to a semi-permeable membrane (regenerated CA, with a
molecular weight cutoff of 3500 Da; Fisher Scientific) using an
adhesive (Devcon epoxy adhesive). Another PDMS sheet, in which a
test cell (for example, in a volume of 500 .mu.L) is fabricated
along with inlet and outlet channels and wells for introduction of
glucose solution, is finally adhesive bonded to the CA membrane. A
fabricated cantilever-based sensor is shown in FIG. 3.
Viscosity-Based Sensor
Vibrational Diaphragm
[0065] In a different embodiment of the disclosed subject matter,
the sensor is a viscosity-based sensor that includes a vibrational
diaphragm, which is also referred as "diaphragm-based sensor." In
these embodiments, the diaphragm can be made of any applicable
polymer known in the art, including but not limited to the polymer
Parylene. In specific embodiments, the diaphragm can be a Paralene
membrane. The vibrational Parylene diaphragm can be situated inside
a microchamber, as shown in FIGS. 12A and 21. The microchamber is
filled with the solution of a biocompatible polymer that binds with
the analyte, and is equipped with a cellulose acetate
semi-permeable membrane, which allows the analyte to permeate into
and out of the chamber while keeping the polymer from escaping. In
one embodiment, the diaphragm is embedded with a moving top
electrode, which forms a capacitor with a fixed bottom electrode.
Separating the electrodes is a sealed air gap. A set of permalloy
thin-film strips are also integrated on the diaphragm. The
permalloy and moving top electrode are passivated by additional
Parylene layer to avoid direct contact with the polymer solution.
The interaction of the polymer and the analyte leads to increased
vibrational damping of the Parylene diaphragm, thus producing a
measureable capacitance change across the metal electrodes. The
capacitance changes are defined as the difference between the
sensor capacitance values with and without the presence of the
analyte.
[0066] When the analyte permeates through the semi-permeable
membrane, it interacts with the polymer to result in a viscosity
change, which is detected via vibration measurements. Specifically,
an externally applied time varying magnetic field acts upon the
permalloy strips, which are magnetized along their length by a
permanent magnet. This results in a time varying moment in the
permalloy, directed in the in-plane direction perpendicular to the
strip length. Under the action of the moment, the diaphragm
vibrates, whose deflection is detected from the capacitance change
between the electrodes. As the viscous damping on the vibration
directly depends on the viscosity, the measured capacitance change
can be used to determine the viscosity change, and hence the
presence and property of the analyte, for example, the
concentration of the analyte.
[0067] In one embodiment, as shown in FIG. 21, a sensor 15 includes
a semi-permeable membrane 16 which is permeable to a sample analyte
17, a substrate 18, and a microchamber 19 which includes a Parylene
diaphragm 20. The microchamber 19 is filled with a solution of
polymer 21 that binds with the analyte 17. The semi-permeable
membrane 16 allows the analyte 17 to permeate into and out of the
microchamber 19 while keeping the polymer 21 from escaping. The
Parylene diaphragm 20 is embedded with a moving top electrode 22
within the microchamber 19, which forms a capacitor with a fixed
bottom electrode 23. The bottom electrode 23 is separated from the
Parylene diaphragm 20 and top electrode 22 by a sealed air gap 24.
A set of permalloy thin-film stripes 25 are integrated on the
Parylene diaphragm 20. The permalloy film stripes 25 and top
electrode 22 are passivated to avoid direct contact with the
solution of polymer 21. The polymer includes, but is not limited to
a polymer having a plurality of boronic acid moieties, for example,
PAA-ran-PAAPBA.
[0068] The diaphragm vibration is in general a complex physical
phenomenon involving the intimate coupling of the motion of the
continuously deflecting diaphragm and the flow of the viscous
polymer solution. Nonetheless, useful insight can be gained into
this phenomenon with a simplified analysis, in which the diaphragm
was represented as a one-degree-of-freedom (1-DOF)
mass-spring-damper system. In one embodiment, the permalloy strips
are collectively represented as a 1-DOF rigid plate that can rotate
about a fixed axis under a magnetically applied torque. The
diaphragm outside the plate region is assumed to have negligible
inertia while applying a linear elastic restoring torque on the
plate. The interaction of the plate and diaphragm motion with the
polymer solution can be represented as a linear viscous torque.
Then the equation governing the plate rotation, .theta., under a
torque, T(t), takes the form:
I{umlaut over (.theta.)}+D{umlaut over (.theta.)}+K{umlaut over
(.theta.)}=T(t) (2)
where I is the plate's moment of inertia, D the viscous damping
coefficient, and K the diaphragm's spring constant.
[0069] Consider the steady-state motion of the permalloy plate
under a harmonic magnetic excitation, T(t)=T.sub.me.sup.i.omega.t,
at frequency f=.omega./2.pi.. Equation (2) can be reformulated into
a form:
{umlaut over (.theta.)}+2.zeta..omega..sub.0{dot over
(.theta.)}+.omega..sub.0.sup.2.theta.=.theta..sub.me.sup.i.omega.t
(3)
where .omega..sub.0=2.pi.f.sub.0=(K/1).sup.1/2 is the natural
frequency, and .xi.=1/2D/IK).sup.1/2 is the dimensionless damping
ratio. In addition, .theta..sub.m=T.sub.m/K is the plate rotation
at zero excitation frequency.
[0070] The steady-state solution to Equation (3) is of the form
.theta.=Ae.sup.-ioe.sup.i.omega.t, where Ae.sup.-io is the complex
amplitude with the amplitude (A) and phase shift (o) of the plate
rotation respectively. By defining p=f/f.sub.0, these quantities
are given by
A.sub.x=A cos
.phi.=.theta..sub.m(1-p.sup.2)/[(1-p.sup.2).sup.2+4.zeta..sup.2p.sup.2]
(4)
A.sub.y=A sin
.phi.=.theta..sub.m(2.zeta.p)/[(1-p.sup.2).sup.2+4.zeta..sup.2p.sup.2]
(5)
[0071] Equations (4) and (5) can be fitted to the experimental
data, recognizing that the diaphragm rotation is proportional to
the sensor output. For example, when the analyte of interest is
glucose, a 1-DOF mass-spring-damper model fitted to the
experimental data obtained at a glucose concentration at 90 mg/dL
is shown in FIG. 27. It is important to note that to obtain
consistent results, it is not appropriate to fit A.sub.x or A.sub.y
(or equivalently A or o) independently to the data. Instead, it is
necessary to fit a vector-valued function {A.sub.x, A.sub.y} as a
function of f as given in Equations (4) and (5) to the experimental
data, yielding consistent estimates of the parameters
.theta..sub.m, f.sub.0, and .xi.. As shown in FIG. 27, a
least-square fit yields the following estimates: f.sub.0=1190 Hz
and .theta..sub.m=0.38 V, with .xi.=0.39, which shows the model
agrees with the experimental data well, considering that the
fitting involves a vector-valued function, i.e., the model must fit
two sets of experimental data with a single set of parameter
estimates. The model given by Equations (4) and (5) can also be
fitted to experimental data obtained at other glucose
concentrations as shown in FIG. 27. The value of .theta..sub.m is
consistently estimated to be almost constant at 0.38 V, with
variations less than 0.9% as the glucose concentration varies from
30 to 360 mg/dL. In addition, the estimated natural frequency
f.sub.m changes only by 0.6%, suggesting that inertial
contributions of viscous effects are insignificant. As shown in
FIG. 28, the damping ratio estimated from the fits increases
steadily, by 20% as the glucose concentration varies from 30 to 360
mg/dL. This is consistent with increased viscosity of the polymer
solution at elevated glucose concentrations.
[0072] FIG. 22 illustrates an exemplary manufacturing process for
the diaphragm-based sensor. The fabrication process begins with the
deposition and patterning of chrome (5 nm) and gold (100 nm) to
form the fixed bottom electrode 26 on the thermally grown SiO.sub.2
layer on a silicon wafer (420.times.420.times.0.1 .mu.m). A
sacrificial photoresist layer 27 (5 .mu.m) is then spin-coated and
patterned to define an electrode air gap 28 (see FIG. 22A),
followed by the deposition of a Parylene Diaphragm 29 (3 .mu.m). A
second layer of chrome (5 .mu.m) and gold (100 nm) are next
deposited for the moving top electrode 30 and permalloy seed layer
(see FIG. 22B). Subsequently, with the permalloy strips defined by
a photoresist mold (5 .mu.m), permalloy 31 (2 .mu.m) is
electroplated. This is followed by the removal of the photoresist
mold, patterning of the moving top electrode 30, and deposition of
an additional Parylene layer 32 (3 .mu.m) for passivation (see FIG.
22C). Two etching holes (500.times.500 .mu.m) are opened through
the two Parylene layers 29 and 32 by oxygen plasma to expose the
sacrificial photoresist layer (see FIG. 22D), which is subsequently
removed by acetone (80.degree. C.) to release the diaphragm 29.
These two etching holes are then sealed by epoxy 33 (Devcon) (see
FIG. 22E). After wafer dicing and wire bonding, a chip is bonded to
a polycarbonate sheet (thickness: 500 .mu.m), in which holes of
appropriate sizes are drilled to define the microchamber as well as
the inlet and outlet (each 10 .mu.L) for polymer solution handling.
The polycarbonate is in turn bonded to a semi-permeable membrane
(regenerated cellulose acetate with a molecular weight cutoff of
3500; Fisher Scientific) using epoxy (see FIG. 22F).
Permittivity-Based Sensor
[0073] In another embodiment of the disclosed subject matter, the
sensor for monitoring a target analyte is a "permittivity-based
sensor", which adopts permittivity measurements. The
permittivity-based sensor employs a solution of a biocompatible
polymer, for example, PAA-ran-PAAPBA which contains phenylboronic
acid moieties that specifically bind to the analyte. A chamber
between two parallel-plate electrodes, a top electrode and a bottom
electrode, is filled with a solution of the polymer. The electrodes
are imposed with an AC electric field which causes the polarization
of the polymer manifested as a permittivity. The binding between
the polymer and the analyte causes the polymer to crosslink,
thereby changing the polarization behavior and hence permittivity
of the polymer. Thus, measuring the capacitance between the
electrodes allows monitoring the presence and property of the
analyte, including, but not limited to, determining the
concentration of the analyte. In one embodiment, the
permittivity-based sensor consists of a pair of parallel-assembled
glass coverslips each coated with a thin-film copper electrode. The
gap between the electrodes, defined by a photoresist spacer layer,
is filled with PAA-ran-PAAPBA solution mixed with glucose, as shown
in FIG. 15A.
[0074] One embodiment of a permittivity based sensor is shown in
FIG. 15A, a sensor 34 has a top electrode 35 which is embedded with
a glass coverslip 36, a bottom electrode 37 which is embedded with
a glass coverslip 38, and a chamber 39 between the top electrode 35
and the bottom electrode 37 defined by a photoresist layer 40. The
chamber 39 is filled with a solution of polymer. The polymer
includes, but is not limited to a polymer having a plurality of
boronic acid moieties, for example, PAA-ran-PAAPBA.
[0075] Another embodiment of a permittivity based sensor is shown
in FIGS. 32A-32C. The device consisted of a microchamber 50, filled
with an aqueous solution. Two gold electrodes deposited on the top
51 and bottom 52 chamber walls were patterned to the chamber shape
and dimensions. A gold thin-film temperature sensor 53 was also
integrated on the bottom chamber wall. An AC E-field imposed on the
electrodes caused the polarization of the polymer polarization,
which directly depended on glucose binding. Thus, the permittivity
could be obtained to determine the glucose concentration.
[0076] Under an electric field (E-field), a dielectric material
dissolved or suspended in a liquid phase undergoes a charge
separation and molecular redistribution, resulting in electric
polarization, i.e., a regular rearrangement of charged particles
such as electrons, ions, and molecules. If the E-field is
periodically varying in time, the polarization can be frequency
dependent as the particle rearrangement does not respond
instantaneously to the field variations. The frequency-dependent
polarization is represented by the complex permittivity, whose real
part is related to the stored electric energy within the material,
and whose imaginary part represents the dissipation or loss of
energy within the material. The permittivity, as a macroscopic
property, is related to the molecular structure of the material
through the polarizability and the molecular dipole moments. In the
context of affinity biosensors, the dielectric material contains a
receptor group for a molecular target. When the receptor binds to
the target molecules, the material can undergo a structural change
as well as alteration of the overall dipole moments. This can be
manifested in terms of changes in the permittivity, which can be
used to determine the concentration of the bound molecules.
[0077] An exemplary manufacturing process for the
permittivity-based sensor, for example, a sensor as shown in FIG.
15A, begins with the deposition and patterning of copper electrodes
on two glass slides. Photoresist, for example, AZ P4620, is then
spin-coated on these two glass slides to prevent the direct contact
between the polymer solution and the electrodes. Another
photoresist layer is then coated and patterned on one of the glass
slide to create a chamber for introduction of the polymer solution.
Finally, these two glass slides are aligned and glued together by
photoresist reflowing.
[0078] In another embodiment, the manufacturing process of the
device shown in FIG. 32A includes a gold thin film 100 nm thick
deposited by thermal evaporation and pattered to form the
electrodes on the top and bottom glass slides, as well as the
temperature sensor 53 on the bottom slide. Passivation photoresist
layers can be spin-coated on the glass slides to prevent the direct
contact of the electrodes with the dielectric solution. Another
photoresist layer can then be deposited and patterned on the bottom
glass slide to define the microchamber along with an inlet and an
outlet for introduction and removal of the polymer solution. The
top glass slide is next placed over the bottom slide and aligned to
form the microchamber along with a parallel-plate capacitor
comprised of the two electrodes. The photoresist layers can be
baked at a set temperature for a set number of minutes on a
hotplate, causing the reflow of the photoresist and bonding of the
glass slides to complete the device fabrication. Images of a
fabricated device before and after packaging are shown in FIGS. 32B
and 32C.
Polymer Solution
[0079] The sensor as provided in the presently disclosed subject
matter can have many applications and therefore can be used to
monitor a number of target analytes. In selecting a target analyte,
an appropriate polymer solution must be used in the sensor.
[0080] In one embodiment, the biocompatible polymer can reversibly
and specifically binds to the analyte of interest. The binding
between the polymer and the analyte can result in changes of the
physical characteristics (e.g., the viscosity and/or permittivity)
of the polymer, which can be measured to extrapolate the presence
and amount of the analyte in the sample. The polymer specially
responds to the analyte of interest. For example, in one
embodiment, the analyte is glucose, through proper adjustment of
the composition percentage of the boronic acid moieties on the
polymer and polymer concentrations, the polymer can detect and
differentiate glucose from other monosaccharides and disaccharides.
Thus, the polymer can be highly specific response to glucose.
Applying this polymer to the sensor as disclosed herein will enable
highly reliable, continuous monitoring of glucose in ISF in
subcutaneous tissue.
[0081] In one embodiment, the binding between the polymer and the
analyte of interest is reversible. In another embodiment, the ester
bonds formed between the polymer and glucose can be broken to
substantially recover the polymer and/or the sample containing
glucose. For example, dialysis of the glucose cross-linked polymer
using water and passing through a semi-permeable membrane can
result in significant decreases of viscosity. This decrease in
viscosity is due to the loss of the ester bonds between the boronic
acid moieties of the polymeric backbone and the glucose in the
sample. As such, the polymer can be recovered and reused to test
other samples for the presence of glucose.
[0082] In one embodiment, a suitable polymer having boronic acid
moieties can be formed as a copolymer of at least two monomers,
where one of the monomers includes at least one boronic acid
functional group. A copolymer can be synthesized with these
monomers via classic free radical copolymerization processes. In
one embodiment, a copolymer is PAA-ran-PAAPBA. In various
embodiments, a suitable polymer includes, but is not limited to, a
polymer that contains boronic acid groups, or other receptor groups
that recognize the given analytes.
[0083] PAA-ran-PAAPBA is an amphiphilic copolymer containing two
components, hydrophilic polymer segment polyacrylamide (PAA) and
hydrophobic polymer segment poly(3-acrylamidophenylboronic acid)
(PAAPBA) (see FIG. 31A). PAAPBA is the glucose-sensitive component,
containing boronic acid groups which can form cyclic boronate
esters in aqueous media after binding with glucose. PAA, which is
water soluble, serves to improve the water solubility of the
hydrophobic PAAPBA segment. In addition to being a water soluble
component, PAA provides an added neighbor coordinating effect by
carbonyl oxygen and boron chelating which enhances the binding
between the boronic acid and carbohydrates (Li S. et al., J.
Diabetes Sci. Tech. (2008); 2(6):1066-1074). The polymer is
synthesized by a free radical polymerization process, e.g., as is
described in detail in Li S. et al. (2009); Li S. et al. (2008) (Li
S. et al., Biomacromolecules (2009); 10:113-118; Li S. et al.,
(2008)).
[0084] A solution of PAA-ran-PAAPBA undergoes a viscosity change
when interacting with glucose molecules. That is, when glucose is
added to the PAA-ran-PAAPBA solution, the phenylboronic acid
moieties in the polymer are able to reversibly form strong ester
bonds with the glucose at a ratio of two to one on the
PAA-ran-PAAPBA polymer backbone, resulting in the cross-linking of
the polymer and hence a conformation change. This leads to a
shorter distance between polymer chains and a further aggregation
of the polymer, thus resulting in an increase in the viscosity of
the polymer (see FIG. 31B-31C), which can be detected by the sensor
provided in the disclosed subject matter. The viscosity change is
reversible due to the reversibility of the polymer-glucose binding,
and is highly specific to glucose over other saccharides as
experimentally verified elsewhere (Li S. et al., (2009); Li S. et
al., (2008)).
[0085] In addition, a solution of PAA-ran-PAAPBA polymer undergoes
dielectric property changes when interacting with glucose
molecules. AAPBA segments bind with glucose at a two to one ratio,
leading to the reversible formation of cyclic esters of boronic
acid. That is, the binding of a glucose molecule results in the
elimination of two hydroxyl groups. Thus, given an equilibrium
dissociation constant on the order of 1.6 mM for the affinity
binding system, it can be estimated that up to about 11% of
permanent dipoles are lost over the glucose concentrations tested.
This directly reduces dipole reorientation effects. In addition,
due to an overall charge decrease on polymer segments as well as
changes in polymer conformations that alter the electric double
layer structure, the glucose binding reduces Maxwell-Wagner-Sillars
and counterion polarization effects. These lead to a decrease in
the permittivity of the polymer, which can be detected by the
sensor provided in the disclosed subject matter.
Applications of the Sensor
[0086] In various embodiments of the disclosed subject matter, the
sensor can be used to determine the level of an analyte in the
body, for example oxygen, lactase, insulin, hormones, cholesterol,
medicaments, viruses, or the like. The sensor can use any known
method to provide an output signal indicative of the concentration
of the target analyte. 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 doctor, for example.
[0087] In a specific embodiment of the disclosed subject matter,
the sensor is used to monitor glucose as the target analyte. In
these embodiments, the sensor can measure a concentration of
glucose or a substance indicative of the concentration or presence
of the glucose by using a specific polymer in the sensor. In one
embodiment, the polymer has boronic acid functional groups
extending from the polymeric backbone (i.e., the boronic acid
moieties) that allows glucose to reversibly form ester bonds with
the boronic acid moiety. These bonds can result in cross-linking of
the polymers in the system, which leads to an increase in the
solution viscosity. The viscosity of the polymer can become steady
within minutes upon changing glucose concentrations. The polymeric
composition can be configured such that a viscosity increase can be
observed when glucose concentration increases. For example, a
viscosity increase can be observed when glucose concentration
increases from about 0 to about 25 mM. In another embodiment, the
cross-linking of the polymers resulted from the binding between the
polymer and glucose leads to a change in the polarization behavior
of the polymer, and hence changes the permittivity of the
polymer.
[0088] The sensor can also be used for other applications. In
addition to diabetes, the proposed miniature CGM device can also be
used for glucose monitoring for other diseases (e.g., glycogen
storage disease and hyperinsulinaemic hypoglycaemia).
[0089] The method can be extended to other metabolites, such as
lactate, fatty acids, cysteines and homocysteines. For example, in
emergency medicine, lactate monitoring can be used to predict
possible organ failure of trauma patients, organ transplant
patients, and patients with other critical conditions.
[0090] Further, the methods disclosed herein can be used as a
reliable method for long-term monitoring of metabolites. Such
methods can have great military significance. For example, a
miniature device for glucose detection with fully electronic
readout would have significant applications in protecting armed
forces in the field. It would also provide a platform to enable the
delivery of drug treatments and nutritional supplements to protect
and enhance performance in military personnel.
[0091] Moreover, the disclosed method can be applied to the
diagnosis of disease. For example, the development of boronic acid
based glucose sensing systems can be potentially extended to other
analytes, such as human viruses and bacteria, since most of those
microorganisms carry glycoproteins on the exterior surface that can
be targeted by the boronic acid based binding motifs.
[0092] Finally, metabolic monitoring is of great utility to
environmental monitoring. Changes in the concentrations of
metabolites are the precursors and products of enzymatic activity,
and can be associated with biological function and regulation.
Metabolic monitoring hence can be used for environmental
monitoring, e.g., risk assessment of chemicals and diagnosis of
diseases in wild animals. It can also be used as a tool to better
understand the underlying mechanisms of action of toxic compounds
in the environment.
EXAMPLES
[0093] The disclosed subject matter will be better understood with
reference to the following Example, which is provided as exemplary
of the disclosed subject matter, and not by way of limitation.
Example 1
A MEMS Affinity Glucose Sensor Using a Biocompatible
Glucose-Responsive Polymer
[0094] In the present example, a sensor having a vibrational
cantilever is used and tested for monitoring the concentration of
glucose.
Experimental Method
[0095] The sensor consists of a Parylene cantilever which vibrates
under magnetic excitation inside a microchamber. The micochamber is
filled with a solution of PAA-ran-PAAPBA that specifically
recognizes glucose by affinity binding. The mircochamber is sealed
with a cellulose acetate semi-permeable membrane, which prevents
the polymer from escaping while allowing permeation of glucose into
and out of the chamber (FIG. 1). Affinity binding between the
polymer and glucose results in the cross-linking of the polymer and
an increase in the viscosity of the polymer solution. Thus, by
measuring the damped cantilever vibration in the solution, the
glucose concentration can be determined. The fabrication process
and experimental setup of the sensor are shown in FIGS. 2 and 9,
respectively. The cantilever vibration was measured with an optical
lever system as shown in FIG. 9.
[0096] Chemical reagents used in the experiments include
d-(+)-glucose (Sigma-Aldrich) and PAA-ran-PAAPBA that is
synthesized using a method described in Li S. et al., (2009).
Phosphate buffer saline (PBS), pH 7.4, was prepared from potassium
phosphate (20 mM), NaCl (150 mM) and NaN.sub.3 (0.05%). Glucose
stock solution (1 M) was prepared by dissolving glucose (1.8 g) in
PBS to 10 mL. A series of glucose solutions (27 mg/dL, 54 mg/dL,
108 mg/dL, 216 mg/dL, and 324 mg/dL) were prepared by further
diluting the stock solution with PBS. To prepare polymer solutions
that were mixed with glucose before being loaded into the
microchamber, PAA-ran-PAAPBA solution (1 mL, 0.45 mg/dL) and
glucose (6 .mu.L, 1 M) were mixed to obtain 108 mg/dL glucose
concentration in PAA-ran-PAAPBA solution. PAA-ran-PAAPBA copolymer
solution with 1.9%, 2.3%, 4.3%, 5%, and 7.3% of PAAPBA composition
was prepared using a method described in Li S. et al. (2008)).
PAA-ran-PAAPBA (284 mg, with 1.9% of PAAPBA in the polymer) was
dissolved in PBS (6 mL) to obtain the sensing solution.
[0097] All experiments were conducted at 37.degree. C. with
closed-loop temperature control by placing the sensor on an
ultra-thin kapton heater with temperature measured by a k-type
thermocouple. This was necessary so as to minimize
temperature-dependent viscosity changes in the fluids, and provide
a physiologically relevant glucose monitoring condition. The
thermocouple was connected to a multimeter (Agilent 34420A Nano
Volt/Micro Ohm meter) to obtain temperature measurements. These
values were then transmitted to a computer to control the voltage
output of the DC power supply (Agilent E3631A DC power supply)
connected to the heater. In all experiments unless otherwise noted,
the sensor's microchamber was filled with a glucose-free polymer
solution (1.9%), while glucose solutions of varying concentrations
were introduced into the test cell. Glucose permeated through the
semi-permeable membrane and bound to the polymer, until this
process reached equilibrium. Because the volume of the test cell
was 20 times than that of the chamber, it was reasonable to assume
that the equilibrium glucose concentration equaled the initial
glucose concentration in the test cell.
[0098] The cantilever vibration was driven by a hand-wound solenoid
(4000 turns of a 200 .mu.m diameter copper wire on a 2.5 cm
diameter plastic core), which under a 5 V.sub.rms driving voltage
generated an electromagnetic of approximately 0.8 kA/m
perpendicular to the cantilever surface. A permanent magnet with
field strength of 500 kA/m was placed parallel to the cantilever
surface to magnetize the permalloy film. The cantilever vibration
was detected by an optical-lever system as described in Zhao Y. et
al. (Zhao Y. et al., (2007)). Briefly, a laser beam was directed
onto and reflected off of the free end of the cantilever. Detection
of the reflected laser beam using a position sensitive detector
allowed determination of the cantilever deflection.
Results and Discussion
[0099] 1. Binding Ability of Copolymer
[0100] To investigate the binding ability of the copolymer with the
glucose molecules, a series of copolymer solutions with varying
PAAPBA composition (1.9% to 7.3%) were synthesized. An Ubbelohde
viscometer (CANNON.RTM. Instrument Company) was used to measure the
viscosity change in the copolymer solution after its interaction
with glucose molecules. As shown in FIG. 13, polymer 1 and 4 with
about 2% PAAPBA had similar responses to glucose concentration from
0 to 918 mg/dL, while polymer 1 showed a higher viscosity due to an
increased PAAPBA percentage. Comparably, copolymer solutions with a
PAAPBA moiety larger than 4.3% exhibited a dramatic change in
viscosity with increasing glucose concentration, indicating
significantly enhanced sensitivity to viscosity with higher PAAPBA
content. The control polymer 3 (using N-phenylacrylamide (NPAA)
instead of N-3-acrylamidophenylboronic acid (AAPBA) as the monomer)
had no boronic acid group in the polymer. It was highly inert to
the glucose concentration change, indicating that AAPBA was indeed
the glucose-reactive component in the copolymer. Since a higher
PAAPBA percentage would significantly increase the viscosity of the
polymer solution and the damping to the cantilever vibration, a
1.9% PAAPBA copolymer solution was more preferable in the following
MEMS sensor experiments to minimize the damping of the cantilever
vibration and ensure large cantilever response.
[0101] The reversibility of the PAA-ran-PAAPBA copolymer,
indicating the stability of this copolymer over time, was measured
using the Ubbelohde viscometer. A 4.7% copolymer solution retained
inside a semi-permeable dialysis tubing, was repeatedly immersed
into a PBS buffer and 450 mg/dL glucose solution. The blank
copolymer solution showed a viscosity of 6.4 cP. In the glucose
solution, the viscosity increased to 20.4 cP. After dialysis
against PBS buffer, its viscosity significantly dropped to 5.4 cP
(FIG. 14), indicating that removal of glucose dissociated the
crosslinking network and lowered the viscosity of the copolymer
solution. The viscosity after copolymer/glucose binding was
slightly different over time, which can be attributed to the loss
of polymer on the dialysis sensor.
[0102] 2. Response Time
[0103] To characterize the response time of the sensor to glucose
concentration change, the time constant of an experiment involving
glucose permeation through the membrane and binding to the polymer
was obtained. The chamber of the sensor was initially filled with
glucose-free polymer solution, while the test cell was filled with
PBS buffer. Glucose solution (108 mg/dL) was then introduced into
the test cell. The cantilever vibration amplitude, which was
proportional to the amplitude of the voltage output of the position
sensitive detector, at a fixed frequency (28 Hz) was obtained over
time (FIG. 4). It can be seen that the amplitude decreased
gradually with time, corresponding to a steady increase in the
damping on the cantilever vibration due to glucose binding-induced
viscosity increase. The cantilever vibration amplitude finally
saturated to a constant level, reflecting that the cantilever
vibration had reached steady state and the process of glucose
permeation and binding had reached equilibrium. The time constant
of this process was determined to be approximately 3 minutes. This
is appropriate for CGM applications (Reifman J. et al., J. Diabetes
Science & Tech. (2007); 1:478-486), considering from
approximately 5 to 15 minutes of response time for commercially
available systems and a approximately 5 minutes detection
requirement for general clinical treatment.
[0104] 3. Evaluation of Glucose Permeation and Equilibrium Binding
Through the Membrane
[0105] The saturated amplitude frequency response of the
cantilever, obtained from the methods above after glucose
permeation and binding had reached equilibrium, was compared with
results from a comparison experiment. In the comparison experiment,
the microchamber and test cell were both filled with a polymer
solution that was mixed with glucose at 108 mg/dL. Thus, the
glucose concentrations inside and outside the microchamber were
pre-equilibrated at the predetermined value. The cantilever
vibration amplitude frequency responses obtained from these two
experiments are shown in FIG. 5. The responses exhibited resonances
at almost identical frequencies (27.0 Hz and 27.2 Hz) with nearly
the same amplitudes (48 .mu.V and 47.6 .mu.V). The amplitudes at
other frequencies also agree within 6%.
[0106] The small discrepancies between the two responses can be
attributed to the required separate preparation of the samples used
in the two experiments. These experiments confirm that the process
of glucose permeation through the membrane and binding to the
polymer indeed achieved equilibrium, and the sensor would be
capable of accurately determining glucose concentrations in its
implanted environment.
[0107] 4. Steady-State Response at Varying Glucose
Concentrations
[0108] To investigate the dependence of cantilever vibration
characteristics on the excitation frequency, the steady-state
cantilever vibration was then measured at varying, physiologically
relevant glucose concentrations. The vibration exhibited resonance
behavior at all glucose concentrations tested (FIGS. 6-7). As the
glucose concentration increased from 27 mg/dL to 324 mg/dL, the
resonance peak of the vibration amplitude decreased consistently by
about 70% (FIG. 6). This was accompanied by a downward shift of the
resonance frequency by about 0.77 Hz (from 27.54 to 26.77 Hz, see
FIG. 6) and an attenuation of vibration Q-factor from 29 to 7.
These observations indicate a significant increase in vibrational
damping, which is consistent with increased viscosity of the
polymer solution at higher glucose concentrations. In addition,
there was a significant change in the phase lag between the
cantilever vibration and the magnetic excitation for the cantilever
vibration (FIG. 7). For example, at 10 Hz, the phase lag increased
from 2.2 degree at 27 mg/dL to 28.3 degree at 324 mg/dL (FIG. 7),
and at 15 Hz, the phase shift increased from 4 degree at 27 mg/dL
to 30 degree at 324 mg/dL (FIG. 7). Based on the resolution of the
phase measurements (0.01 degree), this implies that the sensor
would be able to resolve glucose concentrations at about 0.1 mg/dL
resolution.
[0109] 5. Simulated Glucose Variation Measurements
[0110] First, the reversibility of the sensor with respect to
glucose concentration changes was tested by alternatively measuring
0 and 108 mg/dL glucose solutions (FIG. 10). The measured vibration
amplitude at 28 Hz repeatedly alternated between 37 and 43
.mu.V.
[0111] Measurements of a time-dependent sequence of glucose
variations were made to simulate possible glucose concentration
changes in the interstitial fluid of a diabetes patient, and also
evaluate the reversibility and stability of the sensor. In these
measurements (FIG. 8), while a glucose concentration of 108 mg/dL
represented a stable daily glucose level, glucose concentrations of
54 mg/dL and 324 mg/dL were used respectively to simulate glucose
levels before and after intake of food. In addition, two
intermediate glucose concentrations, 162 mg/dL and 216 mg/dL, were
also measured. The measured vibration amplitude at 28 Hz varied
from 43 .mu.V at 54 mg/dL to 20 .mu.V at 324 mg/dL, and remained to
be 37 .mu.V at 108 mg/dL. Various noise observed in the data is
attributable to environmental disturbances to the optical setup. In
particular, when the sensor was exposed to a glucose concentration
after experiencing another sample that was either higher or lower
in concentration, virtually the same amplitude was consistently
obtained. For example, the average amplitudes at 108 mg/dL over the
two periods, approximately defined by the intervals of [110, 200]
and [800, 1100] min, were respectively 36.83 .mu.V and 36.36 .mu.V,
which agree within 1.3%. Similarly, the reversibility was within
1.06% and 1.13% for the measurement data at 162 and 216 mg/dL
glucose concentrations, respectively. This data indicates an
excellent reversibility of the sensor in response to glucose
concentration variations, indicating its ability for long-term
continuous monitoring of glucose in subcutaneous tissue without
need of recalibration.
[0112] Moreover, the drift in the sensor response was assessed by
exposing the sensor to a glucose concentration over an extended
period. A consistent vibration amplitude measurement of 37 .mu.V
was observed and there was virtually no drift of this measured
signal over a preliminary measurement period of 5 hours (FIG. 11).
In addition, in the aforementioned experiment as shown in FIG. 8,
for the [800, 1100] min period during which the glucose
concentration was maintained at 108 mg/dL, the vibration amplitude
was steady at 36.36 .mu.V with a standard deviation of 0.427 .mu.V.
That is, the drift in the sensor response was about 0.17 .mu.V, or,
0.5% per hour. The drift can be further reduced by minimizing
osmotic effects across the membrane, ultimately allow a stability
appropriate for long-term continuous glucose monitoring.
[0113] Experimental results have shown that this sensor responded
to glucose concentration variations at a time scale of
approximately 3 minutes which is shorter compared with time
responses of commercially available electrochemical CGM sensors.
Additionally, comparative experiments confirmed that the process of
glucose permeation through the membrane and binding to the polymer
was able to equilibrate the glucose concentrations inside and
outside the mirochamber, making sensor suitable for an implanted
setting. Moreover, the sensor response obtained at varying,
physiologically relevant glucose concentrations indicated that the
sensor was capable of resolving glucose concentration changes by
measurement of viscosity. Specifically, over a glucose
concentration range of 27-324 mg/dL, the phase frequency response
could allow a glucose concentration resolution of 0.1 mg/dL.
Finally, the sensor was tested with a time-dependent sequence of
glucose variations to simulate possible glucose concentration
changes in the interstitial fluid of a diabetes patient. The
measurement data indicated that the sensor response was highly
reversible (within 1.2%) and stable (within 0.5%/h). These results
demonstrate the potential of sensor for use as a subcutaneously
implanted device for stable and reliable continuous monitoring of
glucose in practical diabetes management.
Example 2
A MEMS Affinity Glucose Sensor Using Permittivity Measurements
[0114] In the present example, a sensor using permittivity
measurements is used and tested for monitoring the concentration of
glucose.
Experimental Method
[0115] The disclosed MEMS affinity glucose sensor uses permittivity
measurements This sensor employs a biocompatible glucose-specific
polymer, PAA-ran-PAAPBA, as the sensing solution. The polymer
solution is bound between two parallel-plate electrodes imposed
with an AC electric field, which causes the polarization of the
polymer manifested as a permittivity. Glucose binding causes the
polymer to crosslink, thereby changing the polymer's polarization
behavior and hence permittivity. Thus, measuring the capacitance
between the electrodes allows determination of glucose
concentration. The sensor consists of a pair of parallel-assembled
glass coverslips each coated with a thin-film copper electrode. The
gap between the electrodes, defined by a photoresist spacer layer,
is filled with PAA-ran-PAAPBA solution mixed with glucose, as shown
in FIG. 15A.
[0116] To fabricate the device, copper electrodes were first
deposited and patterned on two glass slides. AZ P4620 photoresist
was then spin-coated on both slides to prevent the direct contact
between polymer solution and the electrodes. Another photoresist
layer was then coated and patterned on one of the glass slide to
create a chamber for introduction of the polymer solution. Finally,
these two glass slides were aligned and glued together by
photoresist reflowing.
[0117] The capacitance between the electrodes was measured using a
capacitance/voltage transformation circuit (FIG. 16). As shown in
FIG. 16, the imaginary part of the output voltage, which was
proportional to the polymer solution's complex permittivity, was
then obtained. The specificity of device was investigated by
measuring polymer solution premixed with glucose and fructose. The
device was then characterized by obtaining the frequency-dependent
complex permittivity of the polymer solution at various
physiological-relevant glucose concentrations. In addition, the
device response to glucose solution (free of polymer) was also
demonstrated. Finally, the device stability was assessed over an
extended measuring period of about 10 hours to evaluate the
device's potential suitability for long-term, stable CGM
applications.
Results and Discussion
[0118] First, the device response (at 10 kHz) to varying glucose
and fructose concentrations was investigated (FIG. 17). As the
glucose concentration varied from 0 to 360 mg/dL, the complex
voltage amplitude of the circuit decreased monotonically by 10%
from 0.149 to 0.135 V while showing virtually no response to
fructose, suggesting glucose-specific detection. Next, the device
at additional frequencies was investigated (FIG. 18). Strong
frequency dependence of the device complex permittivity was
observed, indicating a frequency-dependent polymer polarization. At
any given frequency, the imaginary part of the capacitive circuit
output consistently decreased with glucose concentration.
Measurements with polymer-free glucose solutions showed no response
to glucose concentration changes, indicating that the polymer was
critical for dielectrically based glucose detection (FIG. 19).
Finally, measurements of a polymer solution with 90 mg/dL glucose
over 10 hours showed a minimal drift of 9 ppm/hr (FIG. 20),
suggesting an excellent stability, ideal for long-term CGM
applications.
Example 3
A Capacitive MEMS Viscometric Sensor for Affinity Detection of
Glucose
[0119] In the present example, a sensor having a vibrational
diaphragm is used and tested for monitoring the concentration of
glucose.
Experimental Method
[0120] The sensor consists of a Parylene diaphragm, which vibrates
under magnetic excitation inside a microchamber and whose
deflection is measured capacitively. The microchamber is filled
with a solution of PAA-ran-PAAPBA, and equipped with a cellulose
acetate semi-permeable membrane (FIG. 21A-21B). The semi-permeable
membrane prevents the polymer from escaping while allowing
permeation of glucose into and out of the microchamber. Affinity
binding between the polymer and glucose results in the crosslinking
of the polymer and an increase in the viscosity of the polymer
solution. Thus, by measuring the damped diaphragm vibration in the
solution, the glucose concentration can be determined. The
fabrication process and experimental setup of the sensor are shown
in FIGS. 22A-22F and 25, respectively.
[0121] The PAA-ran-PAAPBA polymer was synthesized by free radical
polymerization (Li S. et al., (2009); Li S. et al., (2008)), and
D-(+)-glucose was purchased from Sigma-Aldrich. Phosphate buffer
saline (PBS), pH 7.4, was prepared from potassium phosphate (20
mM), NaCl (150 mM) and NaN3 (0.05%). PAA-ran-PAAPBA (284 mg, with
5% of PAAPBA in the polymer) was dissolved in PBS (6 mL) to obtain
a solution. Glucose stock solution (1 M) was prepared by dissolving
glucose (1.8 g) in PBS to 10 mL. A series of glucose solutions (30
mg/dL, 60 mg/dL, 90 mg/dL, 120 mg/dL, 210 mg/dL, and 360 mg/dL)
were prepared by further diluting the stock solution with PBS.
[0122] All experiments were conducted at 37.degree. C. with
closed-loop temperature control to simulate a physiologically
relevant glucose monitoring condition and minimize
temperature-dependent viscosity changes. During testing, the
sensor's microchamber was filled with an initially glucose-free
solution of PAA-ran-PAAPBA (PAAPBA content: 5%). To facilitate
experimentation, a test cell (volume: 300 .mu.L) was constructed
from a polycarbonate sheet directly above the sensor (FIG. 23). A
glucose solution at a given concentration was introduced into the
test cell, where it was allowed to permeate through the sensor's
semi-permeable membrane to interact with PAA-ran-PAAPBA in the
microchamber. Because the volume of the test cell was 30 times that
of the microchamber, it was reasonably assumed that the glucose
concentration inside the microchamber equalized to the given
glucose concentration in the test cell when the glucose permeation
reached an equilibrium.
[0123] The sensor capacitance was measured using a
capacitance-voltage transformation circuit (FIG. 24). Given an
input sinusoidal wave, U.sub.i (1 V.sub.rms), the output voltage,
U.sub.0, had an amplitude that was proportional to the ratio of the
sensor capacitance, C.sub.x, to the standard reference capacitance,
C.sub.f This output was captured by a lock-in amplifier and
acquired by a personal computer. In the experimental setup (FIG.
25), the diaphragm vibration was driven by a home-made solenoid
(400 turns of a 250 .mu.m diameter copper wire on a plastic core),
which, under a driving voltage of 10 V.sub.pp, produced a magnetic
field strength of about 950 A/m at 1000 Hz perpendicular to the
cantilever surface. A permanent magnet bar with a field strength of
approximately 200 kA/m was placed parallel to the permalloy strips
to yield saturated magnetization of the permalloy. A fabricated,
not yet packaged sensor is shown in FIG. 12A, while a packaged
sensor in the experimental setup is shown in FIG. 12B.
Results and Discussion
[0124] The sensor's vibration characteristics at physiologically
relevant glucose concentrations under various excitation
frequencies were evaluated, and then the observed characteristics
with a simplified oscillator model were analyzed. The temporal
course of the diaphragm vibrations due to changes in glucose
concentration was observed to determine the sensor's response time
and its reversibility. Finally, the drift in the sensor response in
glucose measurements over an extended measuring period was
investigated to evaluate the sensor's potential suitability for
long-term, stable CGM applications.
[0125] 1. Measured Diaphragm Vibration Characteristics
[0126] The dependence of vibration characteristics of the sensor
diaphragm on the excitation frequency was first characterized. In
the experiment, the glucose concentration was first allowed to be
equilibrated at a physiologically relevant value (30, 60, 90, 120,
210, or 360 mg/dL). The diaphragm vibrated under the excitation of
a harmonically time-varying magnetic field, which had a
frequency-independent amplitude of approximately 110 A/m. The
steady-state amplitude and phase of the diaphragm vibration as a
function of the excitation frequency were obtained in terms of the
output voltage of the capacitive measurement circuit (FIG. 27). As
shown in amplitude frequency response (FIG. 26A), the diaphragm
vibration exhibited resonance behavior at all glucose
concentrations tested. The resonance peaks were relatively broad
because of significant damping from the highly viscous polymer
solution. As the glucose concentration increased from 30 to 360
mg/dL, the resonance peak decreased consistently by 53 mV (from 542
to 489 mV). This was accompanied by a downward shift of the
resonance frequency by 100 Hz (from 1000 to 900 Hz). These
observations indicate a significant increase in vibrational
damping, which is consistent with the increased viscosity of the
polymer solution at higher glucose concentrations.
[0127] In addition, from the phase frequency response (FIG. 26B),
it can be seen that at a given frequency, there was a significant
change in the phase shift between the diaphragm vibration and the
magnetic excitation. For example, at 400 Hz, the phase shift
decreased from 17.68 degree at 30 mg/dL to 7.69 degree at 360
mg/dL, which agrees with increased damping at higher glucose
concentrations. The phase shift curves at the different glucose
concentrations intersect at a single frequency of approximately
1200 Hz, at which the phase shift was 91.8 degree. This is
consistent with the behavior of a single-degree-of-freedom damped
harmonic oscillator, and suggests that the natural frequency of the
diaphragm, taking into account the added mass from the polymer
solution, was about 1200 Hz.
[0128] 2. Time-Resolved Measurements of Sensor Response to Glucose
Concentration Changes
[0129] Having systematically characterized the diaphragm vibration
characteristics, time-resolved measurements of the diaphragm
vibration in response to glucose concentration changes were
performed, and these measurements were applied to assess the
response time, reversibility and drift in the sensor response.
[0130] To characterize the sensor response time, the glucose
concentration was initially allowed to be equilibrated at 90 mg/dL
in the test cell and microchamber. Next, the solution in the test
cell was replaced with another glucose solution at 120 mg/dL. When
the glucose concentration inside the microchamber had equilibrated
to 120 mg/dL, the reverse process was initiated, in which the test
cell was refilled with a 90 mg/dL glucose concentration. Some of
the polymer-bound glucose molecules dissociated and permeated out
of the semi-permeable membrane, allowing the glucose concentration
inside the microchamber to equilibrate to 90 mg/dL. The process of
solution refilling of the test cell lasted about 10 seconds, which
was sufficiently fast when compared with the glucose concentration
equilibration. During the equilibration processes, the harmonic
vibration of the diaphragm, at a fixed frequency of 1000 Hz, was
measured as a function of time. A fixed frequency is used for a
much larger excitation magnetic field amplitude (250 A/m) than that
used above (110 A/m) when the frequency was varied, which was
limited by the frequency-dependence of the magnetic field generated
by the voltage-controlled solenoid.
[0131] As shown in FIG. 29, while the glucose concentration varied
from 90 to 120 mg/dL, the diaphragm vibration amplitude decreased
with time, corresponding to increased damping on the diaphragm
vibration due to glucose binding-induced viscosity increase. The
vibration amplitude finally saturated to a constant level,
reflecting that the process of glucose permeation and binding had
reached a dynamic equilibrium. The time constant of this process
was determined to be approximately 1.5 minutes. In the reverse
process where the glucose concentration in the test cell was
decreased from 120 mg/dL to 90 mg/dL, the vibration amplitude
increased with time due to reduced viscous damping from the polymer
solution. The time constant for the reverse process was
approximately 1.7 minutes. The longer reverse time constant could
be due to the smaller diffusivity of glucose molecules in the
initially more viscous polymer solution and needs to be
investigated in future work. Note that these time constants compare
favorably with response times of commercially available systems
that range from 5 to 15 minutes ("MiniMed Paradigm.RTM. REAL-Time
Insulin Pump and Continuous Glucose Monitoring System,"
http://www.minimed.com/products/insulinpumps/index.html; "FreeStyle
Navigator.RTM. Continuous Glucose Monitoring System,"
http://www.abbottdiabetescare.com; "DexCom.TM. STS.RTM. System,"
http://www.dexcom.com), and can be further reduced by shortening
the distance between the semi-permeable membrane and the diaphragm.
Thus, the experimentally determined glucose response time constants
(approximately 1.5 and approximately 1.7 minutes) are considerably
shorter than a theoretically estimated time constant (approximately
6 minutes), which does not yet include the time taking by
glucose-polymer binding kinetics. This suggests that the sensor
exhibited a faster response time than theoretically expected. It
was conjectured that this was caused by the possible active mixing
of glucose in the chamber as induced by the diaphragm vibration,
which could also have induced the vibration of the semi-permeable
membrane by fluid-structure interactions.
[0132] The reversibility of the sensor response, which can be
obtained by comparing differences in sensor output between two
separated measurements at the same glucose concentration were also
assessed. For example, as shown in FIG. 29, the sensor output at
1000 Hz varied from 0.598 (averaged over the period [0, 5] minutes)
to 0.588 V (averaged over [9, 14] minutes) as the glucose
concentration varied from 90 to 120 mg/dL. The sensor output then
returned to 0.598 V (averaged over [19, 24] minutes) when the
glucose concentration was reversed to 90 mg/dL. The difference
between the average sensor outputs over the two periods with the
glucose concentration at 90 mg/dL was only about 0.3 mV, or 60 ppm.
Thus, there is excellent reversibility in the sensor with respect
to glucose concentration variations.
[0133] In general, the measurement accuracy of the sensor is
primarily determined by three factors. That is, in addition to
reversibility considerations above, the sensor accuracy is also
limited by the repeatability of the sensor output from multiple
measurements at a certain glucose concentration, and the noise in
the measurement as the glucose concentration is held constant. In
terms of repeatability, for example, multiple measurements of
glucose samples at 90 mg/dL at 1000 Hz were performed, and the
measurements resulted in the sensor output differing by only about
90 ppm. The accuracy was more significantly influenced by the
measurement noise, which can be observed in FIGS. 29-30. This noise
is attributed to randomly present tiny air bubbles in the polymer
solution that influenced the diaphragm vibration, as well as small
temperature fluctuations in the chamber due to limitations in
temperature control. The noise, characterized by standard
deviations from FIG. 29, was about 0.32 mV, or 3% of the sensor
output change (10.7 mV) as the glucose concentration was varied
from 90 to 120 mg/dL. This translates into a glucose measurement
resolution of about 1.8 mg/dL at 90 mg/dL, which is considered
excellent in the context of practical applications.
[0134] Finally, the drift of the sensor output by exposing it to
constant glucose concentrations over long periods was studied. For
example, the harmonic vibration amplitude at 1000 Hz as the glucose
concentration was held constant at 90 mg/dL (FIG. 30). It can be
seen that the sensor output was steady at 0.596 V over a period of
about 10 hours, with a drift rate of less than 10 .mu.V/hr. The
drift can be further reduced by measures such as minimization of
osmotic effects across the membrane, optimization of the parameters
in closed-loop temperature control, and development of improved
capacitance measurement circuitry. The low drift demonstrates that
the sensor holds potential to offer highly stable measurements for
long-term continuous glucose monitoring.
[0135] The MEMS sensor vibration characteristics obtained at
several selected glucose concentrations indicated that the sensor
was capable of resolving glucose concentration changes by viscosity
measurements. By fitting the experimental data to a simple 1-DOF
mass-spring-damper model, the sensor's natural frequency was
estimated to be 1190 Hz, while the dimensionless damping ratio
increased from 0.37 to 0.44 as the glucose concentration varied
from 30 to 360 mg/dL, reflecting the steady increase of the
viscosity of the polymer solution with glucose concentration.
Experimental results have also shown that the sensor responded
quite rapidly to glucose concentration variations with a time
constant of approximately 1.5 minutes which is shorter than the
time responses of commercially available electrochemical CGM
sensors. Additionally, it was also observed from the experimental
data that the sensor response to glucose concentration changes was
highly reversible; for example, as the glucose concentration was
changed from 90 to 120 mg/dL and then reversed 90 mg/dL, the
deviation in the diaphragm vibration amplitude was only 60 ppm.
Finally, it was demonstrated that the sensor response was highly
stable. For example, as the glucose concentration was held constant
at 90 mg/dL, the drift rate in the diaphragm vibration amplitude
was only 0.17 ppm/hr. These results demonstrate that the sensor
holds the potential to be used as a subcutaneously implanted device
for long-term, stable and reliable continuous monitoring of glucose
in practical diabetes management.
Example 4
MEMS-Based Dielectric Affinity Biosensing
[0136] In the present example, a sensor using permittivity
measurements is used and tested for monitoring the concentration of
glucose.
Experimental Methods and Materials
[0137] The sensor consists of a microchamber 50 filled with an
aqueous solution of PAA-ran-PAAPBA, mixed with glucose at
physiologically relevant concentrations. Two gold electrodes
deposited on the top 51 and bottom 52 chamber walls were patterned
to the chamber shape and dimensions. A gold thin-film temperature
sensor 53 was also integrated on the bottom chamber wall. An AC
electromagnetic (EM) field imposed on the electrodes caused the
polarization of the polymer polarization, which was directly
depended on glucose binding (FIGS. 32A-32C). Thus, the permittivity
can be obtained to determine the glucose concentration.
[0138] To fabricate the device, a gold thin film 100 nm thick was
deposited by thermal evaporation and pattered to form the
electrodes (each 1.times.1 mm.sup.2 in area) on the top and bottom
glass slides, as well as the temperature sensor (280.times.200
.mu.m.sup.2 in area, and 40 .mu.m in line width) on the bottom
slide. Thereafter, passivation photoresist layers (500 nm in
thickness) were spin-coated on the glass slides to prevent the
direct contact of the electrodes with the PAA-ran-PAAPBA dielectric
solution. Another photoresist layer 55 (thickness: 3 .mu.m) was
then deposited and patterned on the bottom glass slide to define
the microchamber (over a 1.5.times.1.5 mm.sup.2 area on the slide)
along with an inlet and an outlet for introduction and removal of
the polymer solution. The top glass slide was next placed over the
bottom slide and aligned to form the microchamber along with a
parallel-plate capacitor comprised of the two electrodes. The
photoresist layers were finally baked at 180.degree. C. for 10
minutes on a hotplate, causing the reflow of the photoresist and
bonding of the glass slides to complete the device fabrication.
Images of a fabricated device before and after packaging are shown
in FIGS. 32B and 32C.
[0139] The sensor was filled with a sample consisting of the
polymer solution mixed with glucose at varying concentrations, and
measured as shown in FIG. 33A. The temperature of the device was
fixed at 37.degree. C. using closed-loop control, in which the
device was heated by a Peltier heater (Melcor, CP14) according to
feedback from the integrated temperature sensor. The sensor was
coupled to a capacitance/voltage transformation circuit (FIG. 33B)
driven by a sinusoidal input of angular frequency .omega.=2.pi.f
from a function generator (Agilent, 33220A). All experiments were
conducted at frequencies up to 100 kHz as allowed by the lock-in
amplifier. When the device was inserted into the circuit (with the
switch "T" connected to the position "S" in FIG. 33B), an input
voltage U.sub.1Se.sup.j.omega.t from the function generator yielded
an output voltage U.sub.2Se.sup.j.omega.t which was measured using
a lock-in amplifier (Stanford Research Systems, SR844). This was
immediately followed by switching the terminal "T" to a reference
capacitor C.sub.R (position "R" in FIG. 33B), with an input voltage
U.sub.1Re.sup.j.omega.t yielding an output voltage
U.sub.2Re.sup.j.omega.t measured similarly. These allowed the
determination of the device's complex admittance as follows:
Y.sub.S=G.sub.S+j.omega.C.sub.S=j.omega.C.sub.R(U*.sub.2SU*.sub.1R)/(U*.-
sub.1SU*.sub.2R) (6)
where C.sub.S and G.sub.S are the equivalent capacitance and
conductance, respectively.
[0140] PAA-ran-PAAPBA polymers with varying compositions were
synthesized by free radical polymerization (Li S et al., (2008).
Three polymer compositions, respectively with an acrylamide (AA) to
3-acrylamidophenylboronic acid (AAPBA) molar ratio of 50, 20 and
12.5 (or approximately 2%, 5% and 7% in PAAPBA content) were used
in the experiments. Corresponding to these compositions, the
polymers had molecular weights of 176,800, 170,700 and 71,700 Da,
respectively. To prepare solutions of the polymers, 284 mg of each
polymer was dissolved separately in 6 mL of phosphate buffered
saline (PBS;). The PBS buffer (pH 7.4) was prepared from potassium
phosphate (20 M NaCl (150 mM) and NaN.sub.3 (0.05%). D-(+)-glucose,
D-(+)-galactose and D-fructose, used respectively as target and
interfering sugars during dielectric affinity detection, were
purchased from Sigma-Aldrich. Glucose stock solution (1 M) was
prepared by dissolving glucose (1.8 g) in PBS to 10 mL. A series of
polymer solutions mixed with glucose at varying concentrations (30,
60, 90, 120, 210, and 360 mg/dL) were prepared. Polymer solutions
were also mixed with galactose and fructose.
Results
[0141] Regarding the device response to varying glucose
concentrations at frequencies from 1 to 100 kHz, the microchamber
was filled with a polymer solution premixed with glucose at
physiologically relevant concentrations. The device's equivalent
capacitance as a function of frequency is shown in FIGS. 34A and B
for a typical PAA-ran-PAAPBA polymer composition, which had a
PAAPBA content of 5%. The sensor capacitance decreased consistently
with the frequency (FIG. 34A), showing a relaxation of the polymer
solution's dielectric behavior (below). In addition, at a given
frequency, the sensor capacitance decreased consistently with
increasing glucose concentration (FIG. 34B). The permittivity of
the solution decreases due to the binding of the polymer with
glucose. In addition, given the voltage measurement resolution of
the experimental setup (.about.70 .mu.V), the data is used along
with a differential form of Eq. (6) to determine that the device is
capable of resolving glucose concentration changes down to 0.5
mg/dL, which would be sufficiently accurate for practical
continuous glucose monitoring applications.
[0142] The dependence of the measured equivalent capacitance on
glucose concentration at a fixed frequency (100 kHz) when the
device was filled with PAA-ran-PAAPBA polymers with different
PAAPBA contents (2%, 5% and 7%) is shown in FIG. 35. In the absence
of glucose, the equivalent capacitance increased with the PAAPBA
content of the polymer. Additionally, a higher PAAPBA content
offers a higher sensitivity of the equivalent capacitance, and
hence the solution permittivity, to the glucose concentration.
[0143] The polymers with 2% and 5% PAAPBA contents were more
sensitive at the low glucose concentrations, while showing a
somewhat saturated response as the glucose concentration increased.
The polymer with 7% PAAPBA content was highly sensitive at the
elevated glucose concentrations, although there was a significance
decrease in sensitivity at the low end of the glucose concentration
range tested. This significantly different trend can be
attributable to a transition from a liquid state to a gel-like
structure of the 7%-PAAPBA polymer solution at elevated glucose
concentrations, a phenomenon not observed for the other polymer
compositions.
[0144] To investigate the specificity of the device to glucose, the
device filled with a 5%-PAAPBA polymer was tested with unspecific
monosaccharides such as fructose and galactose (FIG. 36), which
represented potential interferents with glucose measurements. In
all concentrations tested, the device responses to fructose and
galactose were less than 0.4% and 5% of the device response to
glucose. These unspecific responses can be considered negligible,
given that fructose and galactose exist in interstitial fluid at
concentrations about 1000 times lower than glucose. These results
show that the dielectrically based affinity glucose sensing
approach can be highly specific.
[0145] The foregoing merely illustrates the principles of the
disclosed subject matter. Various modifications and alterations to
the described embodiments will be apparent to those skilled in the
art in view of the inventors' teachings herein. Features of
existing methods can be seamlessly integrated into the methods of
the exemplary embodiments of the disclosed subject matter or a
similar method. It will thus be appreciated that those skilled in
the art will be able to devise numerous methods which, although not
explicitly shown or described herein, embody the principles of the
disclosed subject matter and are thus within its spirit and
scope.
[0146] Patents, patent applications, publications, procedures, and
the like are cited throughout this application, the disclosures of
which are incorporated herein by reference in their entireties.
* * * * *
References