U.S. patent application number 17/650056 was filed with the patent office on 2022-05-19 for microneedle arrays for biosensing and drug delivery.
The applicant listed for this patent is National Technology & Engineering Solutions of Sandia, LLC, North Carolina State University, The Regents of the University of California. Invention is credited to Thayne L. Edwards, Phillip Miller, Roger Narayan, Ronen Polsky, Joseph Wang, Joshua Ray Windmiller.
Application Number | 20220151516 17/650056 |
Document ID | / |
Family ID | 1000006125290 |
Filed Date | 2022-05-19 |
United States Patent
Application |
20220151516 |
Kind Code |
A1 |
Wang; Joseph ; et
al. |
May 19, 2022 |
MICRONEEDLE ARRAYS FOR BIOSENSING AND DRUG DELIVERY
Abstract
Methods, structures, and systems are disclosed for biosensing
and drug delivery techniques. In one aspect, a device for detecting
an analyte and/or releasing a biochemical into a biological fluid
can include an array of hollowed needles, in which each needle
includes a protruded needle structure including an exterior wall
forming a hollow interior and an opening at a terminal end of the
protruded needle structure that exposes the hollow interior, and a
probe inside the exterior wall to interact with one or more
chemical or biological substances that come in contact with the
probe via the opening to produce a probe sensing signal, and an
array of wires that are coupled to probes of the array of hollowed
needles, respectively, each wire being electrically conductive to
transmit the probe sensing signal produced by a respective
probe.
Inventors: |
Wang; Joseph; (San Diego,
CA) ; Windmiller; Joshua Ray; (Del Mar, CA) ;
Narayan; Roger; (Raleigh, NC) ; Miller; Phillip;
(Greensboro, NC) ; Polsky; Ronen; (Albuquerque,
NM) ; Edwards; Thayne L.; (Albuquerque, NM) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
The Regents of the University of California
North Carolina State University
National Technology & Engineering Solutions of Sandia,
LLC |
Oakland
Raleigh
Albuquerque |
CA
NC
NM |
US
US
US |
|
|
Family ID: |
1000006125290 |
Appl. No.: |
17/650056 |
Filed: |
February 4, 2022 |
Related U.S. Patent Documents
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Application
Number |
Filing Date |
Patent Number |
|
|
16169939 |
Oct 24, 2018 |
|
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17650056 |
|
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|
15687145 |
Aug 25, 2017 |
10136846 |
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16169939 |
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|
|
|
14965755 |
Dec 10, 2015 |
9743870 |
|
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15687145 |
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14342536 |
Jul 30, 2014 |
9737247 |
|
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PCT/US2012/053544 |
Aug 31, 2012 |
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14965755 |
|
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61530927 |
Sep 2, 2011 |
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Current U.S.
Class: |
1/1 |
Current CPC
Class: |
A61B 10/0045 20130101;
A61M 5/158 20130101; A61M 2037/0023 20130101; A61B 5/14546
20130101; A61B 5/6833 20130101; A61B 5/157 20130101; A61M 2037/003
20130101; A61B 2010/008 20130101; A61B 2562/046 20130101; A61B
2562/028 20130101; A61B 5/4845 20130101; A61B 5/1473 20130101; A61B
5/14865 20130101; A61B 5/7282 20130101; G01N 33/48785 20130101;
A61M 2005/1726 20130101; A61B 5/685 20130101; A61B 5/14532
20130101; A61B 5/14514 20130101; A61B 5/150984 20130101; A61M
5/1723 20130101 |
International
Class: |
A61B 5/1473 20060101
A61B005/1473; G01N 33/487 20060101 G01N033/487; A61M 5/158 20060101
A61M005/158; A61B 5/00 20060101 A61B005/00; A61B 5/145 20060101
A61B005/145; A61B 5/1486 20060101 A61B005/1486; A61B 5/15 20060101
A61B005/15; A61B 5/157 20060101 A61B005/157; A61B 10/00 20060101
A61B010/00; A61M 5/172 20060101 A61M005/172 |
Goverment Interests
STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH OR DEVELOPMENT
[0002] This invention was made with government support under grants
no. N00014-08-1-1202 awarded by the Office of Naval Research (ONR),
grant no. DE-AC04-94AL85000 awarded by the U.S. Department of
Energy to Sandia Corporation, and grant no. 151337 awarded by
Sandia National Laboratories Laboratory Directed Research and
Development (LDRD). The government has certain rights in the
invention.
Claims
1-20. (canceled)
21. An analyte sensor for measuring an analyte in a biological
fluid of a user, the analyte sensor comprising: a substrate
comprising atop skin-facing surface, a bottom surface opposite the
top surface, and a thickness therebetween; an array of
electrically-conductive probes, each electrically-conductive probe
extending perpendicularly from the bottom surface of the substrate
through the thickness of the substrate and past the top surface of
the substrate; an array of conductors, wherein each conductor is
coupled to a corresponding electrically-conductive probe at the
bottom surface of the substrate, and wherein at least one
electrically-conductive probe includes an electrode comprising a
coating configured to interact with the analyte and produce an
electrical sensing signal that transfers through the at least one
electrically-conductive probe to a corresponding conductor.
22. The analyte sensor of claim 21, wherein the analyte sensor is
integrated into an adhesive patch for placement on skin.
23. The analyte sensor of claim 22, wherein the adhesive patch is
integrated with electronics configured for communication.
24. The analyte sensor of claim 21, wherein the biological fluid is
transdermal fluid, extracellular fluid, interstitial fluid, or
blood.
25. The analyte sensor of claim 21, wherein an electrochemical
interaction between the analyte and electrode is detectable using
amperometry, voltammetry, or potentiometry.
26. The analyte sensor of claim 21, wherein a first
electrically-conductive probe of the array is configured to detect
a first analyte and a second electrically-conductive probe of the
array is configured to detect a second, different analyte.
27. The analyte sensor of claim 21, wherein the coating includes an
entrapped biocatalyst.
28. The analyte sensor of claim 27, wherein the biocatalyst is
glucose oxidase.
29. The analyte sensor of claim 21, wherein each of the
electrically-conductive probes of the array is individually
addressable.
30. The analyte sensor of claim 21, wherein each conductor is
configured to transmit the sensing signal produced by the
respective electrically-conductive probe to a sensor circuit for
processing.
31. The analyte sensor of claim 21, wherein each of the
electrically-conductive probes of the array is solid.
32. The analyte sensor of claim 21, wherein each of the
electrically-conductive probes of the array comprises silicon,
glass, metal, or a resorbable polymer.
33. The analyte sensor of claim 21, wherein the coating comprises a
conducting polymer.
34. The analyte sensor of claim 21, wherein the array of
electrically-conductive probes includes at least one
electrically-conductive probe comprising a counter electrode.
35. The analyte sensor of claim 21, wherein the array of
electrically-conductive probes includes a subset of
electrically-conductive probes comprising a reference
electrode.
36. The analyte sensor of claim 21, wherein the analyte includes at
least one selected from: a biochemical, a metabolite, an
electrolyte, an ion, a pathogen, and a microorganism.
37. The analyte sensor of claim 21, wherein the electrode comprises
platinum.
38. The analyte sensor of claim 21, wherein a periphery of the at
least one electrically-conductive probe has a non-conductive
material.
39. The analyte sensor of claim 21, wherein the at least one
electrically-conductive probe is insulated within the
substrate.
40. An analyte sensor for transdermally measuring an analyte in an
interstitial fluid of a user, the analyte sensor comprising: a
substrate comprising atop skin-facing surface, a bottom surface
opposite the top surface, and a thickness therebetween; an array of
solid, electrically-conductive probes, each electrically-conducting
probe extending perpendicularly from the bottom surface of the
substrate through the thickness of the substrate and past the top
surface of the substrate; an array of conductors, wherein each
conductor is coupled to a corresponding electrically-conductive
probe at the bottom surface of the substrate, and wherein at least
one electrically-conductive probe includes a platinum electrode
disposed on a surface of the at least one electrically-conductive
probe, wherein the platinum electrode is configured to interact
with the analyte and produce an electrical sensing signal that
transfers through the at least one electrically-conductive probe to
a corresponding conductor, and wherein the platinum electrode
comprises a biocatalyst entrapped within a conducting polymer.
41. The analyte sensor of claim 40, wherein the analyte sensor is
integrated into an adhesive patch for placement on skin of the
user.
42. The analyte sensor of claim 41, wherein the adhesive patch is
integrated with electronics configured for communication.
43. The analyte sensor of claim 40, wherein the at least one
electrically-conductive probe is insulated within the
substrate.
44. A method for measuring an analyte within a biological fluid
comprising: providing an analyte sensor integrated with an adhesive
patch, the analyte sensor comprising: a substrate comprising atop
skin-facing surface, a bottom surface opposite the top surface, and
a thickness therebetween; an array of electrically-conductive
probes, each electrically-conductive probe extending
perpendicularly from the bottom surface of the substrate through
the thickness of the substrate and past the top surface; an array
of conductors, wherein each conductor is coupled to a corresponding
electrically-conductive probe at the bottom surface of the
substrate, and wherein at least one electrically-conductive probe
includes an electrode comprising a coating configured to interact
with the analyte and produce an electrical sensing signal that
transfers through the at least one electrically-conductive probe to
a corresponding conductor; placing the adhesive patch on skin to
transdermally contact the array of electrically-conductive probes
of the analyte sensor with the biological fluid; applying an
electrical stimulus signal to the at least one
electrically-conductive probe; measuring a resultant sensing signal
arising by an interaction between the coating on the electrode and
the analyte in the biological fluid; and determining a
concentration of the analyte based on the sensing signal, wherein
the sensing signal is transferred through the at least one
electrically-conductive probe to the corresponding conductor.
45. The method of claim 44, wherein the biological fluid is
interstitial fluid.
46. The method of claim 45, wherein the coating comprises a
conducting polymer with an entrapped biocatalyst.
47. The method of claim 46, wherein the at least one
electrically-conductive probe is insulated within the substrate.
Description
CROSS REFERENCE TO RELATED APPLICATIONS
[0001] This patent application is a continuation of, and claims
priority and benefits of, U.S. patent application Ser. No.
16/169,939 filed Oct. 24, 2018, which is a continuation of U.S.
patent application Ser. No. 15/687,145 filed Aug. 25, 2017, now
issued U.S. Pat. No. 10,136,846, issued Nov. 27, 2018, which is a
continuation of U.S. patent application Ser. No. 14/965,755 filed
Dec. 10, 2015, now U.S. Pat. No. 9,743,870, issued Aug. 29, 2017,
which is a continuation of U.S. patent application Ser. No.
14/342,536 filed Jul. 30, 2014, now U.S. Pat. No. 9,737,247, issued
Aug. 22, 2017, which is a 35 U.S.C. .sctn. 371 National Stage
application of International Application No. PCT/US2012/053544
filed Aug. 31, 2012, which further claims the benefit of priority
of U.S. Provisional Application No. 61/530,927, filed on Sep. 2,
2011. The entire content of the before-mentioned patent
applications is incorporated by reference as part of the disclosure
of this application.
TECHNICAL FIELD
[0003] This patent document relates to biosensors and drug delivery
devices.
BACKGROUND
[0004] Sensing biological events in vitro and in vivo can provide
real-time detection of physiologically relevant compounds, such as
monitoring of metabolites, electrolytes, biochemicals,
neurotransmitters, medically relevant molecules, cancer biomarkers,
and pathogenic microorganisms. Devices that perform such biological
event sensing are known as biosensors, which can provide real-time
detection of physiological substances and processes in living
things. A biosensor is an analytical tool that can detect a
chemical, substance, or organism using a biologically sensitive
component coupled with a transducing element to convert a detection
event into a signal for processing and/or display. Biosensors can
use biological materials as the biologically sensitive component,
e.g., such as biomolecules including enzymes, antibodies, nucleic
acids, etc., as well as living cells. For example, molecular
biosensors can be configured to use specific chemical properties or
molecular recognition mechanisms to identify target agents.
Examples can include evaluating physiologic and pathologic activity
within a tissue, as well as drug discovery and drug screening.
Biosensors can use the transducer element to transform a signal
resulting from the detection of an analyte by the biologically
sensitive component into a different signal that can be addressed
by optical, electronic or other means. For example, the
transduction mechanisms can include physicochemical,
electrochemical, optical, piezoelectric, as well as other
transduction means.
SUMMARY
[0005] Techniques, systems, and devices are disclosed for
biosensing and therapeutic interventions.
[0006] In one aspect of the disclosed technology, a device includes
an array of hollowed needles, in which each needle includes a
protruded needle structure including an exterior wall forming a
hollow interior and an opening at a terminal end of the protruded
needle structure exposing the hollow interior, and a probe inside
the exterior wall to interact with one or more chemical or
biological substances that come in contact with the probe via the
opening to produce a probe sensing signal, and an array of wires
that are coupled to the probes of the array of hollowed needles,
respectively, each wire being electrically conductive to transmit
the probe sensing signal produced by a respective probe.
[0007] Implementations can optionally include one or more of the
following features. For example, the one or more of the probes can
include a functionalized coating configured to interact with an
analyte within a fluid. An electrochemical interaction between the
analyte and the coating on one of the one or more functionalized
probes can be detected using at least one of amperometry,
voltammetry, or potentiometry. The device can further include a
processing unit in communication with the array of wires that
receives the probe sensing signals and uses the probe sensing
signals as data. The processing unit can compare the data to a
threshold value to determine whether the analyte concentration
reflects a healthy or disease state. The processing unit can
determine a pattern in the data that indicates whether the analyte
concentration reflects a healthy or disease state. The processing
unit can multiplex the received probe sensing signals from the
probes. The device can be integrated into an adhesive patch for
placement on skin to detect the analyte residing in transdermal
fluid.
[0008] In another aspect of the disclosed technology, a device
includes a substrate that includes a microneedle with a hollowed
interior located on one side of the substrate, in which the
microneedle includes a wall with an opening to the hollowed
interior, an electrode including a probe, in which the probe is
disposed inside the hollowed interior, and a wire that is connected
to the probe, in which the electrode is functionalized by a coating
over the probe to interact with an analyte to produce an electrical
signal.
[0009] Implementations can optionally include one or more of the
following features. For example, an electrochemical interaction
between the analyte and the coating on the functionalized electrode
can be detected using at least one of amperometry, voltammetry, or
potentiometry. The device can further include a processing unit in
communication with the wire that receives the electrical signal and
uses the electrical signal as data. The processing unit can compare
the data to a threshold value to determines whether the analyte
concentration reflects a healthy or disease state. The processing
unit can determine a pattern in the data that indicates whether the
analyte concentration reflects a healthy or disease state. The
device can be integrated into an adhesive patch for placement on
skin to detect the analyte residing in transdermal fluid. The
device can further can include a polymer film having pores of a
reversibly tunable porosity, in which the polymer film is attached
to an opposite side of the substrate, a protrusion structure
configured on the one side of the substrate, in which the
protrusion structure has a channel between an opening in the
substrate exposing the polymer film and an opening at a terminal
end of the protrusion structure, a containment structure that
contains a chemical substance, in which the containment structure
includes one or more openings attached to the polymer film
positioned above the protrusion structure, and an electrode
attached to the polymer film, in which the electrode provides an
electrical stimulus to trigger an expansion of the pores of the
polymer film to an open state or a contraction the pores of the
polymer film to a closed state. The processing unit can be in
communication with the wire that receives the electrical signal to
use as data and in communication with the electrode to generate the
electrical stimulus. The processing unit can process the data to
determine whether the analyte concentration reflects a healthy or
disease state. The processing unit can actuate the electrode to
apply an electrical stimulus to the polymer film to alter its
permeability from the closed state to the open state, thereby
releasing the chemical substance from the device. The processing
unit can multiplex the received electrical signals and the
actuation of the electrical stimuli.
[0010] In another aspect, a method to sense an analyte and deliver
a therapeutic agent includes detecting a signal produced by an
analyte at an interface with a chemically functionalized probe
configured to electrochemically interact with the analyte within a
biological fluid, in which the signal is transduced to an
electrical signal by the chemically functionalized probe,
processing the electrical signal to determine a parameter of the
analyte, and based on the determined parameter, applying an
electrical stimulus to a valve comprising a porous polymer film
having pores of a reversibly tunable porosity, the valve attached
to a container containing a therapeutic agent, in which the
electrical stimulus alters the permeability of the pores from a
closed state to an open state, thereby releasing the therapeutic
agent into the biological fluid.
[0011] In another aspect, a device includes a substrate that
includes a plurality of microneedles with a hollowed interior
located on one side of the substrate, in which each of the
microneedles includes a wall with an opening to the hollowed
interior, a biosensor module, an actuator module, and a processing
unit in communication with the plurality of wires to receive the
electrical signal and use the received electrical signal as data,
in which the processing unit is in communication with the actuator
electrode to generate the electrical stimulus based on the data.
The biosensor module includes a plurality of sensing electrodes
disposed inside the hollowed interior of a first group of the
plurality of microneedles, the sensing electrodes including a
probe, in which the probe includes a functionalized coating
configured to interact with an analyte within a fluid to produce an
electrical signal, and a plurality of wires, in which one wire of
the plurality of wires is connected to the probe of the sensing
electrodes. The actuator module includes a polymer film having
pores of a reversibly tunable porosity, in which the polymer film
is attached to an opposite side of the substrate, a plurality of
protrusion structures disposed inside the hollowed interior of a
second group of the plurality of microneedles, in which the
protrusion structures includes a channel between an opening in the
substrate exposing the polymer film and an opening at a terminal
end of the protrusion structure, a containment structure that
contains a chemical substance positioned above the polymer film, in
which the containment structure includes one or more openings
coupled to the polymer film positioned above the protrusion
structure, and an actuator electrode attached to the polymer film,
in which the actuator electrode provides an electrical stimulus to
trigger an expansion of the pores of the polymer film to an open
state or a contraction the pores of the polymer film to a closed
state.
[0012] Implementations can optionally include one or more of the
following features. For example, the processing unit can compare
the data to a threshold value to determines whether the analyte
concentration reflects a healthy or disease state. The processing
unit can determine a pattern in the data that indicates whether the
analyte concentration reflects a healthy or disease state. The
processing unit can actuate the actuator electrode to apply the
electrical stimulus to the polymer film to alter its permeability
from the closed state to the open state, thereby releasing the
chemical substance into the fluid. The processing unit can
multiplex the received electrical signals from the probe and the
actuation of the electrical stimuli to the actuator electrode. The
processing unit can include logic gates configured on the
substrate. The device can be integrated into an adhesive patch for
placement on skin to detect the analyte residing in transdermal
fluid.
[0013] The subject matter described in this patent document can be
implemented in specific ways that provide one or more of the
following features. Microneedle array devices and techniques are
described for performing multiplexed sensing applications and/or
drug delivery in an autonomous, minimally-invasive, and controlled
manner. For example, the disclosed technology can be implemented to
detect analytes in living things via electrochemical methods using
microneedle arrays that can be integrated into a patch and applied
to the skin. Biosensing can be implemented directly at the
microneedle-transdermal interface without the uptake and subsequent
processing of biological fluids. Potentiometric, voltammetric, and
amperometric techniques can be used to transduce physiological and
biochemical information using the microneedle array platform, which
can be integrated into one, all-inclusive platform to enable direct
biosensing of multiple analytes in bodily fluids. Additionally, the
biosensing functionality can be coupled with actuation
functionality. For example, a therapeutic agent (e.g., drugs,
vaccines, insulin, hormones, vitamins, anti-oxidants, and other
pharmacological agents) delivery feature can be initiated by
stimuli-responsive conducting polymer nanoactuators. The
biosensor-actuator platform can be integrated on an adhesive patch
to monitor key physiological/biochemical parameters and/or deliver
a therapeutic intervention on demand. The adhesive patch can be
integrated with electronics to allow signal transduction and
communication. The technology can be used as a "sense" constituent
and as a "treat" constituent in an exemplary "Sense-Act-Treat"
feedback loop process, which can be utilized in a variety of
applications that can include, at least, wireless healthcare,
personalized medicine, health profiling, performance/health
monitoring, and athlete/warfighter monitoring.
[0014] For example, the disclosed technology can have wide-ranging
applications within a multitude of fields and disciplines where the
assessment of health in real-time is desired. For example, the
technology can be easily applied for use in the generalized
healthcare, fitness, sport, remote monitoring, wireless healthcare,
personalized medicine, performance/health monitoring, and
warfighter monitoring domains. The minimally-invasive nature of the
technology, combined with its robust architecture, can make the
technology well-suited for diverse biomedical monitoring
applications, e.g., obtaining biomarker signatures for health
profiling, or patterns of bioanalytes as a measure of
performance/fitness. As another example, cancer cells, such as
melanoma, are known to undergo increased levels of glycolysis which
cause localized environments of decreased pH and glucose
concentrations, and increased lactate concentrations; thus the
disclosed technology can simultaneously and locally detect glucose,
lactate, and pH, thereby can be used as a point-of-care clinical
diagnostic device to determine if skin cells are cancerous and give
immediate data before a lengthy biopsy can be performed. For
example, when the technology is used as the "sense" constituent in
the exemplary `Sense-Act-Treat` feedback loop, the technology can
be employed as an element of a smart patch that is able to trend
pertinent physiological/biochemical information for high-risk
patients (e.g., stroke, cardiac, etc.). Moreover, for example, this
feature of the technology can be adapted as a "battlefield
hospital-on-a-patch" that is able to determine the occurrence of
acute injury/trauma and alert the appropriate personnel to
instigate a rapid evacuation of the individual and begin a targeted
treatment regimen. When the technology is used as the "treat"
constituent in the exemplary `Sense-Act-Treat` feedback loop, the
technology can be employed as an element of the smart patch that is
able to provide a targeted therapy for acute events experienced by
high-risk patients (stroke, cardiac, etc.). These exemplary
features can be adapted as a "battlefield hospital-on-a-patch" that
is able to begin a treatment regimen in combat situations where the
rapid evacuation and treatment of injured personnel is not
feasible.
BRIEF DESCRIPTION OF THE DRAWINGS
[0015] FIGS. 1A-1C show schematics of an exemplary microneedle
array device.
[0016] FIGS. 1D-1F show images of an exemplary microneedle and
microneedle array.
[0017] FIGS. 1G-1I show images of an exemplary microneedle array
device on an adhesive patch implemented on a living being.
[0018] FIG. 1J shows an exemplary diagram of a `Sense-Act-Treat`
feedback loop.
[0019] FIG. 1K shows a diagram of an exemplary microneedle array
sensor-actuator featuring individually-addressable microneedles to
monitor for different dysfunctions and individually-addressable
microneedles to deliver a therapeutic agent in response to
detection of the dysfunctions.
[0020] FIG. 1L shows an illustration of an exemplary multiplexed
controlled release of targeted therapeutic agents.
[0021] FIGS. 2A-2C show schematics of an exemplary microneedle
array strip system.
[0022] FIGS. 3A-3C show data plots of exemplary results of
multiplexed detection of various biomarkers on an electrochemical
array strip using individual microneedles.
[0023] FIGS. 4A and 4B show exemplary microneedle arrays fabricated
using screen printing and stencil processes to create a
pattern.
[0024] FIG. 5A shows an illustration of an exemplary process to
fabricate an exemplary bicomponent microneedle electrode array
using solid and hollow microneedle constituents.
[0025] FIG. 5B shows a schematic of the exemplary bicomponent
microneedle electrode array fully-assembled.
[0026] FIG. 5C shows an illustration of an exemplary process to
grow a glutamate oxidase (GluOx)-functionalized
poly(o-phenylenediamine) (PPD) film at the solid microneedle
surface.
[0027] FIG. 5D shows an illustration of the biocatalytic behavior
of the exemplary electropolymerized glutamate
oxidase-poly(o-phenylenediamine) film.
[0028] FIGS. 6A and 6B show scanning electron microscopy (SEM)
images of exemplary solid and hollow microneedle arrays,
respectively.
[0029] FIG. 7 shows a hydrodynamic voltammogram data plot of
exemplary glutamate bicomponent microneedle array electrodes.
[0030] FIGS. 8A and 8B show exemplary chronoamperogram data plots
recorded for increasing concentrations of glutamate.
[0031] FIG. 9 shows exemplary chronoamperogram data plots recorded
in 0.1 M phosphate buffer.
[0032] FIG. 10 shows exemplary data plots showing the stability of
the glutamate response over extended time periods.
[0033] FIGS. 11A and 11B show data plots showing sensitivity of an
exemplary glucose microneedle biosensor.
[0034] FIGS. 12A and 12B show a schematic illustration of an
exemplary microneedle-based multi-channel, multiplexed drug
delivery actuator device.
[0035] FIGS. 13A and 13B show SEM images of the surface morphology
of an exemplary hollow microneedle array.
[0036] FIG. 14 shows images of triggered release of methylene green
from the individually-addressable reservoirs of an exemplary
microneedle-based drug delivery actuator device.
[0037] FIG. 15 shows time-lapse still frame images of the release
of methylene green dye from a single microneedle of an exemplary
microneedle-based drug delivery actuator device.
[0038] FIG. 16 shows an exemplary UV-Vis spectrum data plot
illustrating the absorbance for the release of methylene green dye
from an exemplary microneedle.
[0039] FIG. 17 shows a schematic of an exemplary microneedle during
drug delivery.
[0040] FIGS. 18A-18D show illustrative schematics showing
processing steps for the assembly of an exemplary microneedle array
device.
[0041] FIGS. 19A and 19B show optical images of an array of carbon
fiber electrodes and a single carbon fiber electrode.
[0042] FIGS. 20A and 20B show SEM images of an exemplary hollow
microneedle array.
[0043] FIG. 21A shows an image of skin after application of an
exemplary microneedle array.
[0044] FIG. 21B and FIG. 21C show optical micrographs of hollow
microneedles before and after insertion into skin.
[0045] FIGS. 22A and 22B show SEM images of an exemplary hollow
microneedle array.
[0046] FIG. 23 shows a data plot of a cyclic voltammetric scan of
ferricyanide in KCl versus Ag/AgCl reference and Pt counter
electrodes.
[0047] FIG. 24 shows a data plot of cyclic voltammetric scans of
hydrogen peroxide versus Ag/AgCl reference and Pt counter
electrodes.
[0048] FIG. 25 shows a data plot of voltammetric scans of buffer
solution and ascorbic acid in buffer versus Ag/AgCl reference and
Pt counter electrodes.
[0049] FIGS. 26A and 26B show optical micrographs of the unpacked
and Rh-carbon paste packed microneedle array.
[0050] FIGS. 27A and 27B show SEM images of the unpacked and
Rh-carbon paste packed microneedle constituent of the array.
[0051] FIG. 28A shows plots of hydrodynamic voltammograms of buffer
and hydrogen peroxide at the rhodium-dispersed carbon paste
microneedle electrode.
[0052] FIG. 28B shows plots of chronoamperograms obtained using the
exemplary rhodium-dispersed carbon paste microneedle electrode.
[0053] FIG. 29 shows a plot of a calibration curve obtained for
hydrogen peroxide concentrations.
[0054] FIG. 30A shows plots of chronoamperograms obtained for
lactate.
[0055] FIG. 30B shows a plot of a calibration curve obtained for
lactate concentrations.
[0056] FIG. 31 shows plots of chronoamperograms showing the effect
of physiologically-relevant electroactive interferents.
[0057] FIG. 32 shows a data plot showing the stability of the
electrochemical response of an exemplary microneedle array used for
lactate detection.
DETAILED DESCRIPTION
[0058] Techniques, systems, and devices are disclosed for the
detection of analytes and delivery of therapeutic compounds in
living things using microneedle array based biosensors and
actuators.
[0059] In one aspect, the present technology includes a device
using an array of microscale structures that penetrate the surface
of a biological tissue to detect fluctuations in certain biomarkers
in tissue fluids and/or extracellular fluids. By detecting such
fluctuations, the devices can be used to monitor progression of
diseases, illnesses, and acute injuries, among other conditions.
For example, this can be implemented by loading the microstructures
with electrochemical transducers in the form of microscale needles,
e.g., also referred to as microneedles, microprobes, electrodes, or
probes, which can have different chemical functionalities towards
biochemical and physiological analytes, e.g., such as a
biochemical, metabolite, electrolyte, ion, pathogen, microorganism,
etc. For example, the device can employ various electrochemical
techniques to perform electrochemical reactions directly at the
microneedle/fluid interface and transduce that biochemical
information into an electrical signal (e.g., voltammetric,
potentiometric, amperometric, conductometric, and/or impedimetric),
which can be further processed. For example, each microneedle in an
array can be configured to detect a different analyte and
multiplexed to be addressed by one or more transduction modalities,
e.g., such as an N-array of microneedle elements, in which each
element N1 senses glucose, N2 senses lactate, element N3 senses
fatty acids, etc., and each analyte is transduced using any or all
of voltammetric, potentiometric, amperometric, conductometric,
and/or impedimetric techniques.
[0060] The device can also be used to implement a therapeutic
intervention that utilizes the microneedles to release a chemical
agent (e.g., a drug) into the fluid in a controlled manner at a
particular localized area in which the microneedle has been
applied. The delivery of a targeted therapeutic intervention can be
implemented in response to an acute event or based on a chronic
condition, e.g., monitored by the sensing contingent of the device.
The microneedle array can be configured in conjunction with a
permeability-tunable conducting polymer material to control the
porosity of the polymer. For example, the device can be configured
to include one or more reservoirs containing the chemical agents(s)
coupled to the permeability-tunable conducting polymer material
that is positioned between the reservoir(s) and the substrate of
the microneedle array. Under certain electrochemical stimuli, the
polymer material can selectively be made porous (e.g., change the
porosity of the polymer material), which can effectively act as a
valve that can be selectively opened or closed to transport the
chemical agent from the reservoir through the microneedle (lumen)
and subsequently into the tissue fluid. The chemical agent release
mechanism is electrochemically enabled, e.g., without moveable
parts or microelectromechanical (MEMS) components.
[0061] The actuation of the therapeutic contingent of the exemplary
device can be controlled using an integrated logic system or
processing unit, which can provide an electrical stimulus based on
feedback from the sensing contingent of the exemplary device. For
example, the described sensing contingent of the exemplary device
can continuously monitor a concentration level (e.g., based on
fluctuations in an average level, a maximum or minimum threshold
level, etc.) of a particular analyte associated with normal
function or dysfunction indicative of diseases, illnesses, and
acute injuries or other conditions and transduce the detected
biochemical information associated with the analyte into an
electrical signal. The electrical signal can be processed using a
processing unit, which can include, at least, a processor and a
memory coupled to the processor. For example, the memory may encode
one or more programs that cause the processor to perform one or
more of the method acts described in this patent document, e.g.,
including storing the detected signals, analyzing the detected
and/or stored signals against other stored values (e.g., such as
analyte threshold values indicative of a healthy or dysfunctional
state) and/or determining whether or not to release a chemical
agent using the therapeutic contingent of the exemplary device. For
example, the processing unit may determine the detected analyte
level has exceeded a threshold value stored in the memory, and
subsequently activate the described actuator to release a drug in
response to the determined analyte level. For example, this can be
performed by applying a suitable redox potential, the exemplary
device can "open" and "close" the described polymer in a reversible
manner by changing the intrinsic porosity of the matrix, thus
triggering the flux of medication from an on-body reservoir
directly into the transdermal fluid. Multivariate/multiplexed drug
delivery can be used to implement a therapy in a unique manner,
where drugs can be delivered at each microneedle constituent of the
array. For example, owing to the arrayed nature of the microneedle
structures, multivariate/multiplexed drug delivery can be realized,
and a unique analyte can be detected at each microneedle
constituent of the array or multiple analytes can be detected at
each individual needle through an array of electrodes within a
needle. Moreover, the arrayed nature of an exemplary system can
enable the device to tailor the cocktail of drugs to mitigate
various forms of injury/trauma. Furthermore, the ability to
selectively control the porosity of the membrane by adjusting the
applied redox potential can imply that the flux rate, and hence the
dosage, can be controlled as needed via the integrated biosensor
and/or the logic-gate sensing and processing unit.
[0062] In some examples, the described microneedle
biosensor-actuator technology can be implemented transdermally by
applying an exemplary microneedle device to the skin. In other
examples, the described microneedle biosensor-actuator technology
can be implemented in vivo to other organs within the body, e.g.,
including the liver, the sclera of the eye, etc. In an example of
the biosensing functionality of the disclosed technology, a device
can include an array of microneedles sensor-actuators integrated on
a skin adhesive patch and applied to the skin of a living being to
transdermally monitor physiological and biochemical parameters
(e.g., glucose). In an example of blood glucose monitoring, the
microneedles can be functionalized with glucose oxidase enzyme (a
biocatalyst) that is entrapped within a conducting polymer, e.g.,
in which the electrode component is conductive and functionalized
(e.g., coated) to include the biocatalyst. Upon application of the
patch to the skin, the microneedles penetrate the skin so that
extracellular fluid (e.g., blood) can diffuse into the microneedle.
The biocatalyst, as glucose diffuses into, can convert the glucose
substrate into gluconic acid. In the meantime, since the conversion
of glucose into gluconic acid is an oxidation-reduction (redox)
reaction (e.g., oxidizing glucose), reduction also occurs to oxygen
and water naturally present in the blood to form hydrogen peroxide.
Hydrogen peroxide is an electrochemically active species, which can
be oxidized or reduced at certain potentials at the electrode. For
example, this can be done amperometrically, in which a potential is
applied and current is monitored, or voltammetrically, in which the
potential is changed and current change is monitored. For example,
as the hydrogen peroxide changes at the electrode in which a
potential is applied, the corresponding current change is monitored
as an electrical signal that can be further processed using signal
processing techniques.
[0063] In one embodiment, a minimally-invasive multi-component
microneedle device for detecting an analyte and delivering a
therapeutic compound can include a microneedle array in conjunction
with electrodes, e.g., which can be chemically-functionalized,
enzyme-functionalized, and/or ion-selective electrodes, to perform
multiplexed sensing and actuating applications in an autonomous,
minimally-invasive, and controlled manner.
[0064] FIGS. 1A and 1B show schematics of an exemplary device based
on hollowed needles with probes. FIG. 1A shows a schematic of an
exemplary microneedle array device 100, and FIG. 1B shows an
exemplary schematic of the device unassembled. The microneedle
array 100 includes an array of hollowed microscale-sized needles
101, in which each needle 101 comprises a protruded needle
structure having an exterior wall forming a hollow interior and an
opening at the terminal end of the protruded needle structure to
expose the hollow interior, and a probe 102 formed inside the
exterior wall to interact with one or more chemical or biological
substances that come in contact with the probe 102 via the opening
to produce a probe signal (e.g., such as a sensing signal).
[0065] FIG. 1C shows the exemplary microneedle array 100 including
an array of wires 103 that are coupled to corresponding probes 102
of the array of hollowed needles 101, respectively, e.g., which can
provide an array of individually addressable microneedle sensing
electrodes. The array of wires 103 can be configured within a
substrate 105, e.g., such as an insulative material, which in some
examples can be flexible and adhesive to biological tissue. Each
wire of the array of wires 103 is electrically conductive to
transmit the probe sensing signal produced by a respective probe to
a sensor circuit, in which the probe sensing signals are
processed.
[0066] FIG. 1D shows an image of an exemplary microneedle imaged by
scanning electron microscopy. FIG. 1E shows an image of an
exemplary microneedle array near objects, such as a penny or an
electronic circuit on a printed circuit board, to provide size
scaling of the exemplary microneedle array. FIG. 1F shows a zoomed
view of the image of the exemplary microneedle array.
[0067] The exemplary microneedle array based sensor actuator device
can involve techniques in microfabrication, electrochemistry,
enzyme-immobilized electrodes, and ion-selective electrodes.
Potentiometric, voltammetric, amperometric, conductometric, and/or
impedimetric detection methodologies can be integrated into one
all-inclusive platform, e.g., in order to enable the direct
biosensing of multiple analytes residing in bodily fluids (e.g.,
such as key biomarkers occupying the transdermal fluid). For
example, the microneedle array platform can be integrated on an
adhesive patch that is placed on the skin in order to monitor key
physiological and biochemical parameters transdermally. The
exemplary adhesive patch can further be integrated with electronics
to allow communication and signal transduction. For example,
because the chemical information can be converted to the electrical
domain via electrochemistry, the device can be interfaced with
electronic readout, e.g., which can be analogous to
continuous-monitoring blood glucose devices. In this fashion, for
example, the disclosed technology can miniaturize and integrate
multiple laboratory-based tests into a single arrayed microneedle
sensing platform and provide the ability to deliver an autonomous
therapeutic intervention in a controlled and minimally-invasive
fashion, as well as to tailor a cocktail of drugs for different
forms of injury/trauma.
[0068] FIGS. 1G-1I show images of an exemplary microneedle array
device on an adhesive patch implemented on living beings. FIG. 1G
shows an image of an exemplary adhesive patch employing the
microneedle array sensor-actuator device being worn on a human arm.
FIG. 1H shows an image of another exemplary adhesive patch
employing the microneedle array sensor-actuator device being worn
on an animal. FIG. 1I shows an enlarged image of the exemplary
device after the adhesive patch has been removed from the animal's
skin, e.g., showing the exemplary microneedles intact.
[0069] The disclosed biosensor-actuator technology can be used to
extract relevant physiological information and provide a controlled
therapeutic response based on the detected physiological and
biochemical information. FIG. 1J shows an exemplary diagram of a
"Sense-Act-Treat" feedback loop where the sensed information is
used to control the actuator to adjust the drug delivery. The
information from the sensing operation enables the drug delivery to
be tailored according to the sensed information. A
biosensor-actuator device 180 in FIG. 1J can include a multiplexed
array of microneedles (e.g., which can be configured in a manner as
the exemplary microneedle array 100), in which some microneedles of
the array are configured for sensing and other microneedles of the
array are configured for therapeutic intervention.
[0070] The "Sense-Act-Treat" feedback loop, as shown in FIG. 1J,
includes the sensing contingent of the biosensor-actuator device
180 to extract the physiological information of an analyte from a
biological fluid (e.g., such as transdermal fluid). The exemplary
sensing feature can include individually addressable microneedles
functionalized as an electrochemical transducer that can detect
patterns of biomarker changes, e.g., such as acute conditions or
chronic disease conditions. A unique analyte can be detected at
each microneedle within the sensing array, or multiple analytes can
be detected by several electrodes housed in one microneedle. For
example, various catalysts, biocatalysts, substrates, reagents,
cofactors, and/or coreagents can be immobilized within the
transducers to impart selectivity towards the analyte of interest.
Likewise, ion-selective membranes (or solid state ion selective
components) can be employed with electrochemical measurements to
impart selectivity towards the ions of interest. In some examples,
the sensing contingent of the biosensor-actuator device 180 can
also include analyte logic-gate sensing for direct processing of
the sensed analyte information, as shown in an exemplary diagram
170. For example, the diagram 170 shows two exemplary inputs (e.g.,
Input 1 and Input 2) of the analyte sensing logic, in which the
Input 1 is a detected signal generated at a microneedle probe
configured to detect a first reaction, and the Input 2 is a
detected signal generated at another microneedle probe configured
to detect a second reaction. The diagram 170 shows the two
exemplary input signals passed through a logic gate (e.g., a single
NAND gate in this example), in which the output of the logic gate
can be used as a signal to control a microneedle actuator (e.g., in
which the logic gate output signal is interfaced with an exemplary
conducting polymer to control the porosity of a drug stored in a
reservoir). In other examples, the sensed analyte information can
be processed by a processing unit.
[0071] For example, the input biomarkers for a soft tissue injury
(STI) can include creatine kinase (CK) and lactate dehydrogenase
(LDH), which are incident on a biocatalytic cascade, and can be
representative of the Input 1 and Input 2, respectively, as shown
in the diagram 170. For example, CK converts the creatine substrate
into phosphocreatine, which simultaneously causes the compound to
convert ATP to ADP. In the presence of phosphoenolpyruvate (PEP),
pyruvate kinase (PK) can give rise to pyruvate. If lactate
dehydrogenase is present, the pyruvate can be converted to lactate
while NADH is simultaneously oxidized to NAD+. Thus, the decrease
in NADH can be monitored with respect to time in an amperometric
fashion. For example, since only the presence of both CK and LDH
causes a concomitant decrease in NADH, monitoring Input 1 and Input
2 using the exemplary biosensor-actuator device 180 can effectively
function as a NAND Boolean logic gate.
[0072] The "Sense-Act-Treat" feedback loop, as shown in FIG. 1J,
includes an exemplary image of a processing unit 175 to process the
sensed analyte information as data and employ logic and/or
instructions to control the actuator contingent of the
biosensor-actuator device 180 to release, not release, or adjust
the release of a therapeutic agent. For example, the processing
unit 175 can include a processor and a memory coupled to the
processor. The processing unit 175 can include a power supply,
e.g., including battery sources, renewable energy sources (e.g.,
solar power sources), or self-powering sources (e.g., motion
feedback power sources). The processing unit 175 can include an
input/output (I/O) unit, coupled to the processor and memory, which
can also be connected to an external interface, source of data
storage, or display device. Various types of wired or wireless
interfaces compatible with typical data communication standards,
e.g., including, but not limited to Universal Serial Bus (USB),
IEEE 1394 (FireWire), Bluetooth, IEEE 802.111, Wireless Local Area
Network (WLAN), Wireless Personal Area Network (WPAN), Wireless
Wide Area Network (WWAN), WiMAX, IEEE 802.16 (Worldwide
Interoperability for Microwave Access (WiMAX)), and parallel
interfaces, can be used to implement the I/O unit. For examples,
the I/O unit of the processing unit 175 can be in communication
with the biosensor-actuator device 180 using a wired configuration.
In other examples, the I/O unit of the processing unit 175 can
include wireless communication functionalities to receive data from
the sensing contingent and transmit control data to the actuator
contingent of the biosensor-actuator device 180. In such examples,
the biosensor-actuator device 180 can include a wireless
transmitter/receiver on or remotely tethered (e.g., using wires) to
the substrate facilitating the sensor-actuator microneedle arrays.
In such examples, the wireless transmitter/receiver can be
interfaced with multiplexing capabilities to multiplex the sensing
signals and control signals that are transmitted and received.
[0073] The "Sense-Act-Treat" feedback loop, as shown in FIG. 1J,
includes the actuator contingent of the biosensor-actuator device
180 to deliver one or more drugs to the region penetrated by the
microneedles based on the processed analyte information (sensed by
the sensing contingent). The exemplary drug delivery feature can
enable the autonomous delivery of a targeted therapeutic
intervention in response to the detected acute or chronic
condition. For example, the permeability of the conducting polymer
nanoactuators can be tunable through an autonomous porosity change
controlled by the integrated sensing or enzyme logic system (e.g.,
processed by the processing unit 175), which in turn can control
release of the drug, as illustrated in a diagram 185 of the drug
delivery actuator in FIG. 1J. The diagram 185 shows that a
processed signal represented by 0 does not actuate the release of
the drug (e.g., the porosity of the exemplary conducting polymer
remains in an effectively closed stat). The diagram 185 also shows
that a processed signal represented by 1 does actuate the release
of the drug (e.g., the porosity of the exemplary conducting polymer
is triggered to be in an open state, thereby allowing the drug to
pass through the pores of the polymer and exit the microneedles
into the tissue fluid). The arrayed microneedle structure can allow
multivariate/multiplexed drug delivery, and a unique therapy can be
delivered at each microneedle constituent of the array, as shown in
FIGS. 1K and 1L.
[0074] FIG. 1K shows a diagram of the exemplary microneedle array
sensor-actuator 180 featuring individually-addressable microneedles
to monitor for different dysfunctions and individually-addressable
microneedles to deliver a therapeutic agent in response to
detection of the dysfunctions. The diagram of the
biosensor-actuator device 180, as shown in FIG. 1K, includes
microneedles of the array that are configured for sensing (e.g.,
microneedles 181a, 181b, and 181c) and other microneedles of the
array that are configured for therapeutic intervention (e.g.,
microneedles 181d, 181e, and 181f). In this example, the
microneedle 181a is configured for sensing an analyte associated
with soft tissue injury (STI), the microneedle 181b is configured
for sensing an analyte associated with traumatic brain injury
(TBI), and the microneedle 181c is configured for sensing an
analyte associated with abdominal trauma (ABT). Also in this
example, the microneedle 181d is configured for delivering a drug
associated with treating STI, the microneedle 181e is configured
for delivering a drug associated with treating TBI, and the
microneedle 181f is configured for delivering a drug associated
with treating ABT.
[0075] For example, exemplary analytes associated with STI that can
be detected (e.g., using the sensor-actuator 180) include creatine
kinase, lactate, and lactate dehydrogenase; e.g., ameliorated with
glucocorticoids, NSAIDs. Exemplary analytes associated with TBI
that can be detected include glutamate, ceruloplasmin; e.g.,
ameliorated with acetaminophen. Exemplary analytes associated with
ABT that can be detected include lactate, lactate dehydrogenase;
e.g., ameliorated with acetylsalicylic acid or
iso-butyl-propanoic-phenolic acid.
[0076] FIG. 1L shows an illustrative diagram of a multiplexed
controlled release of a targeted therapeutic cocktail, e.g., where
the polymer actuator(s) is individually addressable for on-demand
release for targeted therapeutic intervention. The diagram of FIG.
1L shows exemplary microneedles of the actuator contingent (e.g.,
microneedles 182d, 182e, and 182f) that are configured for
therapeutic intervention. In this example, the microneedle 182d is
configured for controlled release of a drug 183d that is stored in
Reservoir 1, the microneedle 182e is configured for controlled
release of a drug 183e that is stored in Reservoir 2, and the
microneedle 182f is configured for controlled release of a drug
183f that is stored in Reservoir 3. For example, based on a control
signal received from the processing unit (e.g., like that shown in
the diagram 185 in FIG. 1L), the release of any or all of drugs
183d, 183e, and/or 183f can be controlled (e.g., using
multiplexing) to produce a targeted therapeutic cocktail, e.g., in
which the individually-addressable polymer actuator(s) are actuated
in a manner that can control the size of the porosity (e.g., and
thereby the flow), as well as the duration of the open state,
controlling concentration of each of the released drugs.
[0077] FIGS. 2A-2C show schematics of an exemplary microneedle
array strip system 200. In this example, the system can include a
flat flex cable (FFC) 210 that includes a plurality of conductors
211 (e.g., ten copper conductors) that interconnect the microneedle
array to a connector region 218, e.g., in which the conductors 211
can be 1.5'' length and interface (e.g., right end (re)connected)
to a circuit board, e.g., via a zero insertion force (ZIF)
connector. For example, the exemplary microneedle array strip
system can be configured on a FFC sized to be 11.0 mm by 38.1 mm.
Holes can be opened using laser ablation at the left end to expose
underlying traces. Metal electrodes, e.g., four working electrodes
202 and one counter electrode 207 and one or more reference
electrode(s) 206, can be sputter deposited on the surface over the
openings. Four vented fluidic chambers 215 can be made from laser
ablated Mylar with a pressure sensitive adhesive. The adhesive
layer can also bond microneedle array(s) to the FFC. FIG. 2A shows
a top view of the schematic of the microneedle array strip system
200. FIG. 2B shows a three dimensional view of the schematic of the
microneedle array strip system 200. FIG. 2C shows another three
dimensional view of the schematic of the microneedle array strip
system 200 inserted into a circuit board connector 219.
[0078] In another example, individually addressable electrodes
(microneedles) can be loaded with a carbon paste, carbon fiber, or
conducting polymer transducer and be employed for the detection of
patterns of biomarker changes that reflect optimal health and/or
performance. Using potentiometry, amperometry, or voltammetry,
various catalysts, biocatalysts, substrates, reagents, cofactors,
and/or coreagents can be immobilized within the transducers to
impart selectivity towards the analyte of interest. Likewise,
ion-selective membranes (or solid state ion selective components)
can be employed with electrochemical measurements to impart
selectivity towards the ions of interest. Significant predictive
and diagnostic information can be available in monitoring multiple
biomarkers and in measuring the dynamical pattern of those species
as a measure of the overall health/performance/fitness of the
subject. Patterns in multiple biomarkers can be integrated and
changes in those markers can be assessed over extended time periods
in order to provide a more detailed and accurate temporal
characterization of the negative effects of stress and overtraining
in addition to a plethora of diseases and illnesses.
[0079] For example, arraying the microneedles can allow for
measuring patterns in multiple bioanalytes. Moreover, an analyte or
multiple analytes, such as a catalyst/biocatalyst or other analyte
or biomarker substance, can be immobilized by robust means that can
include electropolymerization/polymer entrapment, electrostatic
interactions, covalent attachment, and direct adsorption. In one
example, a planar solid-state transducer can be an electrode, for
example, use of carbon fiber, carbon paste, and conducting polymers
to form the electrochemical transducer.
[0080] The exemplary device can be utilized in the following
manner. A transdermal microneedle array can be employed; each
microneedle constituent can contain a bored cylindrical vacancy
inside which a three-electrode electrochemical sensing element is
housed (e.g., such as potentiometric, voltammetric, amperometric,
conductometric, impedometric, etc. sensing elements). In one
example, an enzyme (with affinity to a particular biochemical
moiety) can be immobilized on the working electrode of the
three-electrode contingent and amperometry can be performed. In
another example, an ion-selective membrane (with suitable
ionophore) or solid state functionalization can be applied to the
working electrode and potentiometry or voltammetry is performed.
The presence of the analyte, metabolite, electrolyte, or ion of
interest can result in perturbations in the detected current
(enzyme electrode) or potential (ion-selective electrode),
respectively.
[0081] FIGS. 3A-3C show data plots of exemplary results of
multiplexed detection of various biomarkers on an electrochemical
array strip using individual microneedles. FIG. 3A shows a data
plot of multiplexed detection of pH; FIG. 3B shows a data plot of
multiplexed detection of lactate; and FIG. 3C shows a data plot of
multiplexed detection of glucose.
[0082] FIGS. 4A and 4B show exemplary individual microneedle
microsensors that can be addressable through a microelectrode array
mated with the reverse side of the device. For example, the
array(s) can be fabricated using a number of techniques including
photolithography, inkjet printing, and screen printing, among other
techniques. An example of a screen printed microneedle array is
shown in FIG. 4A, and a microneedle array fabricated by a
stenciling process used to define the pattern is shown in FIG.
4B.
[0083] In another embodiment of the disclosed technology, a
minimally-invasive multi-component microneedle device for
electrochemical monitoring and biosensing is described. This
embodiment can comprise the same embodiment(s) like those
previously described, and can therefore implement the entirety of
functionalities of the individual embodiments on a single
embodiment. For example, the disclosed technology can be
implemented for the electrochemical monitoring and biosensing of
the excitatory neurotransmitter glutamate and glucose. In this
exemplary embodiment, a device can include tight integration of
solid and hollow microneedles into a single biosensor array device
containing multiple microcavities. Such microcavities can
facilitate the electropolymeric entrapment of the recognition
enzyme within each microrecess. The resulting microneedle biosensor
array can be employed as an on-body minimally-invasive transdermal
patch, e.g., eliminating extraction/sampling of the biological
fluid, thereby simplifying device requirements.
[0084] Exemplary implementations were performed to demonstrate
various functionalities of the device, e.g., including the
electropolymeric entrapment of glutamate oxidase and glucose
oxidase within a poly(o-phenylenediamine) (PPD) thin film. For
example, the PPD-based enzyme entrapment methodology can enable the
effective rejection of coexisting electroactive interferents
without compromising the sensitivity or response time of the
device. The resulting microneedle-based glutamate and glucose
biosensor can exhibit high selectivity, sensitivity, speed, and
stability in both buffer and untreated human serum. For example,
high-fidelity glutamate measurements (e.g., down to the 10 .mu.M
level) were obtained in undiluted human serum. The exemplary recess
design can also protect the enzyme layer upon insertion into the
skin. The described robust microneedle design can be well-suited
for diverse biosensing applications in which real-time metabolite
monitoring is a core requirement.
[0085] The exemplary microneedle-based glutamate and glucose
biosensor were implemented in ways that demonstrate clinical
application of microdevices for the on-body monitoring of relevant
bioanalytes by minimally-invasive electrochemical biosensors. In
this regard, the microneedle arrays can be configured to provide
pain-free biosensing and, being highly integrated biocompatible
devices, these devices can be fabricated on an industrial scale and
at low cost. For example, the described microneedle arrays can
perform monitoring and biosensing applications without involving
fluid sampling/extraction. For example, a feature of the technology
is that the uptake of bodily fluids (e.g., such as transdermal
fluid) is not required, as is in conventional microfluidic sensing
systems. Through the execution of electrochemistry at the
microneedle-transdermal fluid interface, useful chemical
information can be extracted and directly transduced to the
electronic domain. In this manner, sophisticated and costly
mechanical devices that regulate flow to a detector array/separate
sensing unit can be eliminated from implementation.
[0086] The exemplary microneedle sensing array device can employ a
recess-based microcavity structure that can be designed to confine
the recognition enzyme and protect it upon penetration of the skin.
For example, a bicomponent microneedle biosensor can include an
array of platinum-coated solid microneedles, which can serve as the
working electrode, and a hollow microneedle cover, which can
provide a microcavity that surrounds each solid microneedle. This
exemplary bicomponent microneedle biosensor is shown in FIGS.
5A-5D. FIGS. 5A-5D also illustrate a process to fabricate the
exemplary bicomponent microneedle biosensor.
[0087] FIG. 5A shows a solid microneedle constituent 501 and a
hollow microneedle constituent 503 of an exemplary array of
bicomponent microneedle electrodes 500. The solid microneedle
constituent can be coated with a conductive material to form a
working electrode, e.g., such as a platinum working electrode 502.
FIG. 5A also shows a process to assemble the platinum-coated solid
microneedles 502 with the hollow microneedle cover 503. In some
examples, the assembly includes applying a sealing agent (e.g.,
epoxy) to the non-detecting, non-recessed regions between the solid
microneedle constituent 501 and the hollow microneedle constituent
503. FIG. 5B shows the bicomponent microneedle array electrode 500
fully-assembled. As shown in FIG. 5B, each bicomponent microneedle
electrode 500 in the array includes a recess region that exposes
the needles (e.g., the platinum-coated electrodes 502) of the solid
microneedle constituent 501, e.g., such that the recess can
facilitate enzyme immobilization. FIG. 5C shows the growth of the
glutamate oxidase (GluOx)-functionalized PPD film 504 at the solid
microneedle surface within the recess region of the microneedle
electrode 500 from the o-phenylenediamine (o-PD) monomer of a
o-PD-GluOx solution. For example, fabrication of the array of
bicomponent microneedle electrodes 500 can include applying a
GluOx-PPD thin film to the platinum working electrodes 502 by
immersing the microneedles in an o-PD-GluOx solution. FIG. 5D shows
the biocatalytic behavior of the electropolymerized glutamate
oxidase-poly(o-phenylenediamine) film 504 (illustrated in purple),
e.g., enabling the quantification of glutamate levels within the
transdermal fluid. In these exemplary figures, glucose oxidase
(GOx) becomes substituted in place of GluOx for the quantification
of glucose.
[0088] The recess-based microneedle electrodes 500 can enable
electropolymeric entrapment of the enzyme within the individual
microcavities. As a result, direct transdermal biosensing can be
accomplished without requiring the uptake of the transdermal fluid,
e.g., thereby simplifying device requirements and the sensing
process. The bicomponent recess geometry of the microneedle
biosensor can also provide a greater surface area for enzyme
immobilization with microneedles containing embedded planar
electrodes.
[0089] The recess-based microneedle electrodes 500 can include
electropolymerized PPD thin films that can be employed for the
confinement of enzymes into miniaturized electrode transducers,
e.g., while imparting remarkable permselective properties and a
stable response. In one example, PPD can be used for entrapping
different oxidase enzymes such as glucose oxidase, lactate oxidase,
and glutamate oxidase, along with permselective detection of the
liberated hydrogen peroxide product. As a consequence of their
remarkable permselective properties, PPD films can impart high
selectivity and stability through exclusion of co-existing
electroactive interferences and proteins normally present within
bodily fluids. The described biosensor devices, which can employ
PPD films, can thus facilitate the amperometric detection of
hydrogen peroxide with high substrate selectivity, excellent
sensitivity, operational stability, and rapid response time. In
this manner, the described biosensor devices that employ
enzyme-functionalized PPD films can exhibit considerable sensing
advantages when compared with those based on other immobilization
techniques, as described herein.
[0090] For example, to illustrate the versatility of the disclosed
bicomponent microneedle array platform, an exemplary biosensor
device for amperometric glutamate biosensing is demonstrated. This
exemplary platform can be subsequently extended to glucose
monitoring for the management of diabetes mellitus. For example, an
excitatory neurotransmitter, glutamate, can be implicated in a
number of pathologic medical conditions such as ischemic neuronal
injury, hypoglycemic injury, epilepsy, Alzheimer's disease, and
traumatic brain injury. In addition, elevated glutamate levels in
the circulatory system can be associated with excitotoxicity. Blood
glutamate levels have risen from an average value of 37.5 .mu.M
among healthy patients to 141.3 .mu.M among patients who have
sustained moderate to severe trauma related to intracranial injury.
As such, serum glutamate levels can provide useful insight into the
overall condition of the central nervous system following brain
trauma.
[0091] For example, the described biosensing platform can be
advantageous over biosensors that quantify glutamate levels with a
high degree of invasiveness, e.g., such as by uptaking the
cerebrospinal fluid (CSF) via a catheter or a microdialysis probe
for further analysis. Also for example, the described biosensing
platform can be advantageous over biosensors that typically are
clinically implemented in a hospital setting, e.g., as such
clinical analysis can be a painful, time-consuming, and costly
proposition. In addition, the described biosensing platform can be
amenable to on-body continuous monitoring, especially when access
to the CSF is not feasible. As blood glutamate levels correlate
well with the levels found in the CSF, its extraction from this
hard-to-access bodily fluid is unnecessary under the disclosed
embodiment.
[0092] An exemplary demonstration of the described biosensor device
involves the enzymes glutamate oxidase (GluOx) and glucose oxidase
(GOx) that can be entrapped within the microcavities of the
exemplary microneedle device using different PPD growth processes,
each with its own specific advantage that can be tailored to
specific applications. The PPD-based confinement of the enzymes
within the microneedle cavities can enable the efficient
quantification of glutamate and glucose at pathophysiological
levels within buffer solutions and undiluted human serum. For
example, the minimally-invasive nature of the exemplary device,
combined with its convenient means to achieve enzyme entrapment and
protection, as well as its attractive electroanalytical
performance, can demonstrate its applicability as a practical
patch-type on-body biosensor.
[0093] Exemplary materials and methods to implement the disclosed
embodiment of the technology are presented. The following chemicals
and reagents were used in the described implementations, which
included glutamate oxidase (GluOx, E.C. 1.4.3.11) from E. coli
(recombinant), glucose oxidase (GOx, E.C. 1.1.3.4) from Aspergillus
niger, 1,2-phenylenediamine (o-Pd), L-glutamatic acid (GLU),
D-(+)-glucose (GLC), L-ascorbic acid (AA), uric acid (UA),
L-cysteine (CYS), acetaminophen (ACT), sodium sulfate,
ethylenediaminetetraacetic acid (EDTA), potassium phosphate
monobasic, potassium phosphate dibasic, and serum from human male
(type AB). The exemplary implementations (with the exception of the
serum calibration) were performed in 0.1 M phosphate buffer (pH
7.40) with 0.5 mM EDTA. Ultrapure water (18.2 M.OMEGA.cm) was
employed in all exemplary implementations.
[0094] The instrumentation used in the described implementations
included the following, which was utilized in exemplary
demonstrations and implementations of the disclosed embodiment
under exemplary conditions disclosed herein. A CH Instruments
(Austin, Tex.) model 1232A electrochemical analyzer was employed
for electrochemical measurements. An external Ag/AgCl reference
electrode (CH Instruments CHI111) and a 0.5 mm diameter platinum
wire counter electrode (BASi, West Lafayette, Ind.) were used to
establish a three-electrode electrochemical system. Voltammetric
and chronoamperometric studies were used to evaluate the
electrochemical behavior of the microneedle array electrode at room
temperature (22.degree. C.). In these exemplary electrochemical
implementations, glutamate (or glucose) was added into 2 mL of
phosphate buffer solution or serum (stirred) in order to obtain the
desired concentration. Chronoamperometric currents were sampled for
15 s following the application of the potential step. The
morphology of the bicomponent microneedle array was examined using
a field emission scanning electron microscope (SEM) (Philips XL30,
Amsterdam, the Netherlands). Specimens were coated with chromium
prior to SEM analysis using a sputtering instrument (Energy Beam
Sciences Emitech K575X, E. Granby, Conn.). A deposition current of
130 mA was applied for 30 s to deposit .about.15 nm of chromium on
the sample surface.
[0095] The exemplary solid and hollow microneedle arrays used in
the exemplary implementations were developed in the following
manner. The microneedle designs were prepared using a CAD software,
e.g., Solidworks (Dassualt Systemes S.A., Velizy, France).
Substrate support structures were created with Magics RP 13
(Materialise NV, Leuven, Belgium). For example, the solid needles
were designed and fabricated with a conical in shape and possess
abase diameter of 390.+-.14 .mu.m and a height of 818.+-.35 .mu.m.
The hollow needles were pyramidal in shape with a triangular base.
The dimensions of each hollow microneedle were as follows: an edge
length of 1174.+-.13 .mu.m, a height of 1366.+-.15 .mu.m, and a
vertical cylindrical bore of 342.+-.5 .mu.m diameter on one of the
faces of the pyramid structure. Both the solid and hollow needles
were arranged into 3.times.3 square arrays with 2 mm periodicity.
Substrates for the microneedle arrays were 10 mm.times.10 mm in
extent and possessed thickness values of 2000 .mu.m and 500 .mu.m
for solid and hollow variants, respectively. The three-dimensional
computer models were transferred to a Perfactory.RTM. SXGA Standard
UV rapid prototyping system (EnvisionTEC GmbH, Gladbeck, Germany)
for fabrication. This system used computer models to precisely
guide light from a 150 W halogen bulb over a photocurable material,
e.g., resulting in the selective polymerization of the exposed
material. In some aspects, Eshell 200 acrylate-based polymer
(EnvisionTEC GmbH) can be utilized as the constituent material to
fabricate the microneedle arrays since the resin selectively
polymerizes under visible light and exhibits a Young's modulus of
elasticity of 3050.+-.90 MPa. Moreover, the polymer features
Class-IIa biocompatibility per ISO 10993. A 550 mW output power
beam (step size=50 .mu.m) with a zero-degree tilt was employed for
the polymerization of the resin. Following the fabrication routine,
the arrays were rinsed with isopropanol to remove the unpolymerized
material. The arrays were placed in an Otoflash post curing system
(EnvisionTEC); and post-build curing was performed for 50 s. A
Compex 201 krypton-fluoride (KrF) excimer laser (Coherent, Santa
Clara, Calif.), which can be operated with a 10 Hz repetition rate
and a wavelength of 248 nm, was used to ablate a
commercially-obtained high purity Pt target. This process resulted
in the deposition of thin films of Pt (.about.12 nm) on the surface
of the solid microneedle array. A background pressure of 5 .mu.Torr
was maintained during the 2 min pulsed laser deposition (PLD)
routine, performed at room temperature.
[0096] Adhesive non-conducting epoxy can be applied to the
periphery of the solid microneedle substrate. The hollow
microneedle cover can then be placed over the solid microneedle
substrate. This exemplary procedure is diagrammatically represented
in FIGS. 5A and 5B. For example, the two components (e.g., the
solid microneedle substrate and the hollow microneedle cover) can
be arranged under an optical microscope to align the solid
microneedles within the hollow microneedle aperture. This can form
the bicomponent microneedle array electrode (BMAE), as shown in
FIG. 5C, e.g., which can be mated with a 3 mL syringe (BD
Biosciences, Franklin Lakes, N.J.). For example, the nozzle portion
of the syringe can be removed to facilitate the attachment of the
BMAE, which can be affixed using adhesive epoxy to the syringe tip
for easier handling. A copper wire can be subsequently inserted
into the open end of the syringe in order to create an electrical
contact to the Pt working electrode. A poly(o-phenylenediamine)
(PPD) film can be electropolymerized from a solution of the
o-phenylenediamine (o-Pd) monomer, as shown in FIG. 5C, e.g., to
immobilize the GluOx and GOx enzymes on the electrode surface and
reject potential electroactive interferents. For example, a 0.1 M
phosphate buffer (pH 7.40) solution containing 10 mM o-Pd, 5 mM
sodium sulfate, and 100 U/mL GOx can be purged with nitrogen for 20
minutes at room temperature, which can be used to form the
GOx-functionalized electrode. For example, the BMAE, Ag/AgCl
reference, and platinum counter electrodes can then be immersed in
the solution; a potential of 0.75 V vs. Ag/AgCl can subsequently be
applied for 20 min in order to grow the GOx-entrapped PPD film, as
represented in FIG. 5C. This exemplary process represents a rapid
means to immobilize enzymes and is appropriate for applications in
which the enzyme is of sufficiently low cost such that the entire
extent of the electrode can be immersed in the enzyme-o-PD
solution.
[0097] In other examples, a slight variant of the aforementioned
process can be used, e.g, to conserve the costly GluOx enzyme
during the electropolymerization process. In this alternative
exemplary process, the BMAE, Ag/AgCl reference, and platinum
counter electrodes can be immersed in a solution of 0.1 M phosphate
buffer (pH 7.40) containing 10 mM o-Pd and 5 mM sodium sulfate; and
a potential of 0.75 V vs. Ag/AgCl can subsequently be applied for 5
min. The electrode can then be rinsed and dried at room
temperature. A 0.5 L aliquot solution of 7.5 U/mL GluOx in 0.1 M
phosphate buffer (pH 7.40) can then be dispensed in each recess of
the BMAE; this step can be repeated an additional six times on each
microneedle using a low-retention micropipette. Following this
process, the solution can be allowed to dry at room temperature.
The drop-casting procedure can be repeated five additional times.
Subsequently, a solution of 0.1 M phosphate buffer (pH 7.4)
containing 10 mM o-Pd, 5 mM sodium sulfate, and 1 U/mL GluOx can be
dispensed on each microcavity of the microneedle array. A potential
of 0.75 V vs. Ag/AgCl can subsequently be applied for 15 min in
order to electropolymerize the GluOx-entrapped-PPD film. For
example, whereas the previous described process can be adapted for
simplicity at the expense of increased enzyme usage, this
implementation can facilitate the electropolymerization of more
costly enzymes.
[0098] Following each electropolymerization process, the BMAE can
be rinsed and immersed in a 0.1 M phosphate buffer solution (pH
7.4) for 30 min to remove monomeric residue from the microneedle
structure as well as any non-bound enzyme. When not in use, the
BMAE can be stored in phosphate buffer at 4.degree. C. This
exemplary process, which is diagrammatically represented in FIG.
5D, can enable the quantification of glutamate and glucose, as well
as other analytes.
[0099] The exemplary implementations of the described microneedle
array biosensor device included an evaluation of the surface
morphology of the BMAE which was performed to ascertain the
electrode geometry and surface features. A close examination of the
BMAE surface morphology was executed using SEM. FIG. 6A shows
exemplary scanning electron micrographs of the solid microneedle
arrays. FIG. 6B shows exemplary scanning electron micrographs of
the hollow microneedle arrays. As shown in the SEM images, the
features of the exemplary microneedles closely corresponded with
those specified in the exemplary computer-aided design file. For
example, a uniform distribution of the solid and hollow
microneedles can be observed at the microneedle array. With respect
to the exemplary solid microneedle (FIG. 6A), a notable observation
is the ribbed structure, which can be attributed to the
layer-by-layer approach that was utilized to polymerize the E-shell
200 resin. A uniform pyramidal structure and a triangular base can
be observed at each component of the hollow microneedle array (FIG.
6B), along with apertures of uniform size distribution. Minimal to
no microneedle-to-microneedle geometric variation was observed.
[0100] The exemplary implementations of the described microneedle
array biosensor device included electrochemical characterizations
of the bicomponentmicroneedle electrode towards the amperometric
detection of glutamate. The initial electrochemical
characterization of the BMAE was aimed at constructing hydrodynamic
voltammograms (HDVs), e.g., to select an optimal detection
potential. For example, an HDV can be obtained using
chronoamperometry at varying potentials between 0.1 and 0.6 V vs.
Ag/AgCl (in 50 mV increments). These exemplary characterizations
were performed in the blank buffer solution containing 100 .mu.M of
glutamate. The redox currents was sampled at 15 s following the
potential step. An identical procedure can be followed using
GOx-BMAE for the detection of 10 mM glucose. FIG. 7 shows exemplary
hydrodynamic voltammograms of the glutamate bicomponent microneedle
array electrodes, e.g., in which redox currents were sampled at
t=15 s following the potential step. As is evident from the figure,
the presence of glutamate caused a concomitant rise in the anodic
current, corresponding to the oxidation of the H.sub.2O.sub.2
enzymatic product. The onset of the peroxide oxidation can occur at
.about.0.25 V vs. Ag/AgCl. To minimize the potential oxidation of
interferents in real samples, a potential of 0.40 V vs. Ag/AgCl can
be selected for further electrochemical implementations of the BMAE
biosensor.
[0101] Examples are described for biosensing of glutamate in a
buffer matrix and human serum at the bicomponent microneedle array
electrode. The sensitivity of the exemplary BMAE biosensor was
evaluated using chronoamperometric potential steps to the selected
potential of 0.40 V vs. Ag/AgCl. FIG. 8A shows a data plot of the
average of triplicate chronoamperometric experiments for increasing
levels of glutamate over the entire pathophysiological range (e.g.,
0-140 .mu.M in 20 .mu.M increments, sampled at t=15 s) in a buffer
matrix. A linear calibration plot (R.sup.2=0.995) can be observed
(as shown in the inset of FIG. 8A) over the entire range under
investigation. The exemplary calibration plot exhibits high
sensitivity (s.sub.x=7.129 nA/.mu.M) and a low deviation
(RSD=3.51%); a limit-of-detection (LOD) of 3 .mu.M can be estimated
based on the signal-to-noise characteristics of the experimental
data (S/N=3). The LOD lies well below normal physiological levels,
reflecting the ability of the microneedle sensor to detect
physiologic levels of glutamate.
[0102] Following the calibration experiments in the buffer
solution, exemplary implementations of the described BMAE biosensor
was evaluated by quantification of glutamate in undiluted human
serum samples. FIG. 8B shows a data plot of the average of
triplicate chronoamperometric calibration experiments in human
serum for increasing levels of glutamate over the 0-140 .mu.M range
(e.g., in 20 .mu.M increments). As with the exemplary buffer study,
a linear calibration plot can be observed (as shown in the inset of
FIG. 8B; R.sup.2=0.992) over the entire range. In addition, for
example, the calibration data exhibit high sensitivity
(s.sub.x=8.077 nA/.mu.M) and low deviation (RSD=6.53%). An LOD of
10 .mu.M can be estimated based on the signal-to-noise
characteristics of these data (S/N=3). As with the exemplary buffer
experiments, the LOD obtained from the experimental data resides
below the limit of normal physiological levels. The attractive
behavior of the BMAE in untreated serum samples reflects the
protective ability of the PPD film. Furthermore, the similar
sensitivity obtained for both the buffer- and serum-based trials
again can underscore the robustness of the PPD immobilization
scheme despite the prolonged exposure of the biosensor to
protein-rich serum medium.
[0103] Examples are described for interference investigation(s)
employing physiologically-relevant electroactive compounds. For
example, another advantageous feature of the PPD coating includes
its ability to reject coexisting electroactive interferents even at
moderate oxidation potentials. Accordingly, the selectivity of the
response was examined in the presence of physiological levels of
ascorbic acid (60 .mu.M), uric acid (500 .mu.M), cysteine (200
.mu.M), and acetaminophen (200 .mu.M). FIG. 9 shows a data plot of
exemplary chronoamperograms recorded in 0.1 M phosphate buffer (pH
7.40) for the blank buffer solution, 100 .mu.M glutamate, and 100
.mu.M glutamate in the presence of the electroactive physiological
interferents ascorbic acid (AA, 60 .mu.M), uric acid (UA, 500
.mu.M), cysteine (CYS, 200 .mu.M), and acetaminophen (ACT, 200
.mu.M). The exemplary implementations were carried out under the
same conditions as those represented in FIGS. 8A and 8B. FIG. 9
illustrates the negligible contribution imparted by the presence of
these electroactive compounds upon the current signal for 100 .mu.M
glutamate. Physiological levels of ascorbic acid, uric acid,
cysteine, and acetaminophen resulted in only 0.44%, 0.31%, 1.93%,
and 6.37% average deviations from the 100 .mu.M glutamate current
response, respectively. Hence, natural metabolic fluctuations in
the levels of these electroactive species are not expected to
interfere with the in vivo quantification of glutamate using the
exemplary BMAE device as an on-body biosensor.
[0104] Exemplary implementations were performed for stability
analysis of the exemplary bicomponent microneedle electrode array.
For example, the ability to operate over prolonged periods with
minimal deterioration in the current response can represent another
important feature of on-body biosensors. Accordingly, the stability
of the BMAE response was examined using a 100 .mu.M glutamate
solution over an eight-hour period. FIG. 10 shows a data plot
showing stability of the glutamate response over extended time
periods with each data item referenced to the original current
level at t=0 (100%). In this exemplary implementation, data was
generated from chronoamperograms recorded in 0.1 M phosphate buffer
(pH 7.40) with 140 .mu.M glutamate. The exemplary implementations
were carried out under the same conditions as those represented in
FIGS. 8A and 8B. The exemplary time-course profile of FIG. 10
indicates that the exemplary biosensor exhibits a highly stable
current response, retaining 105% of the original signal level after
eight hours of continuous sampling. In this example, the measured
current never exceeded 110% of the original level over the entire
time period. Consequently, the BMAE can be anticipated to perform
reliably over extended periods associated with body-worn
biosensors.
[0105] Examples are described for the biosensing of glucose with
the exemplary bicomponent microneedle electrode array. For example,
with the glutamate BMAE methodically characterized, the exemplary
platform can be subsequently migrated for use as a glucose
biosensor towards diabetic monitoring. GOx was confined within the
BMAE cavities using a slight variant of the described GluOx
immobilization process. This technique can be amenable to
cost-per-quantity considerations, as stated above. FIG. 11A shows a
data plot featuring chronoamperometric calibration data for
increasing levels of glucose over the entire pathophysiological
range (e.g., 0-14 mM in 1 mM increments) in a buffer matrix. A
well-defined response can be observed over the entire range,
leading to a linear calibration plot (as shown in the inset of FIG.
11A; R.sup.2=0.996). In addition, these exemplary calibration data
exhibit high sensitivity (s.sub.x=0.353 .mu.A/mM) and low deviation
(RSD=6.44%, n=3), along with a LOD of 0.2 mM (S/N=3). It can be
noted that the GOx-functionalized BMAE exhibited a lower
sensitivity towards its substrate when compared with the
GluOx-functionalized platform. For example, this can be attributed
to the different PPD growth process, which may affect the transport
properties of the substrate and product. Accordingly, the GluOx
immobilization process can be followed when high sensitivity is
desired.
[0106] FIG. 11B demonstrates the high selectivity of the exemplary
glucose microneedle biosensor. FIG. 11B shows a data plot of
exemplary chronoamperograms recorded for the blank buffer solution,
10 mM glucose, and 10 mM glucose in the presence of the
electroactive physiological interferents ascorbic acid (AA, 60
.mu.M), uric acid (UA, 500 .mu.M), cysteine (CYS, 200 .mu.M), and
acetaminophen (ACT, 200 .mu.M). FIG. 11B illustrates the
contribution imparted by the presence of potential electroactive
interferents upon the current signal for 10 mM glucose.
Physiological levels of ascorbic acid, uric acid, cysteine, and
acetaminophen resulted in negligible deviations of 1.07%, 0.88%,
1.65%, and 2.21%, respectively, from the current response for 10 mM
glucose. Consequently, as with the exemplary interference
implementation conducted with the glutamate BMAE, natural metabolic
fluctuations of these compounds is not be anticipated to interfere
with the monitoring of glucose.
[0107] An exemplary stability evaluation of the GOx-functionalized
BMAE was performed using a buffer solution containing 10 mM glucose
over an 8-hour period. The GOx BMAE yields a highly stable current
response, with 97% of the original signal level extracted at the
conclusion of the measurement period. In this example, throughout
the time period under investigation, the measured current response
never fell below 87% of the original level. The similar results
described earlier with the GluOx-functionalized BMAE indicate that
both PPD-based immobilization schemes can yield a stable response
over a prolonged period of continuous use.
[0108] The disclosed technology described for this embodiment
includes a bicomponent microneedle array biosensor platform that
can be used for minimally-invasive glutamate and glucose
quantification. The bicomponent microneedle design merges the
inherent advantages of solid and hollow microneedles in order to
form a microcavity that can allow the electropolymeric entrapment
of an enzyme, which can provide protection for the enzyme layer
upon skin penetration, and can eliminate the need for the
extraction of biological fluids. Using a poly(o-phenylenediamine)
thin-film for entrapping the enzymes glutamate and glucose oxidase
can enable the highly sensitive, selective, stable, and rapid
electrochemical detection of glutamate and glucose, respectively.
The high fidelity detection of glutamate in undiluted human serum
samples over the entire pathophysiological range can further
substantiate the utility of the platform as a practical on-body
biosensor. The exemplary patch-type on-body biosensor can enables
the transdermal monitoring of a number of relevant metabolites.
[0109] In another embodiment of the disclosed technology, a
minimally-invasive, multiplexed, multi-component microneedle
actuator device that enables the controlled delivery of multiple
therapeutic agents is described. This embodiment can comprise the
same embodiment(s) like those previously described, and can
therefore implement the entirety of functionalities of the
individual embodiments on a single embodiment. In this described
embodiment, a device can deliver a drug in response to
injury/trauma in an autonomous, minimally-invasive, and controlled
manner that leverages microneedle arrays as the delivery structure,
which can be referred to as a smart NanoPharmacy-on-A-Chip. For
example, microneedle array(s) can be integrated on an adhesive
patch that is placed on the skin in order to deliver on demand a
targeted therapeutic intervention transdermally. The exemplary
technology can integrate the microneedle platform with
stimuli-responsive conducting polymer nanoactuators (with tunable
permeability through an autonomous porosity change controlled by an
integrated sensing or enzyme-logic system). For example, the
disclosed biosensor-actuator device can be used to aid in the rapid
administration of multiple therapeutic agents and counteract
diverse biomedical conditions.
[0110] The described embodiment includes multiple
individually-addressable channels on a single microneedle array,
each paired with its own reservoir and conducting polymer
nanoactuator, which are used to deliver various permutations of the
multiple unique chemical species. For example, upon application of
suitable redox potentials to the selected actuator, the conducting
polymer is able to undergo reversible volume changes, thereby
serving to release a model chemical agent in a controlled fashion
through the corresponding microneedle channels. Exemplary
implementations of the drug delivery contingent of the disclosed
biosensor-actuator device were performed and are described herein.
For example, time-lapse videos were recorded and can offer direct
visualization and characterization of the membrane switching
capability, and, along with calibration investigations, confirmed
the ability of the device to alternate the delivery of multiple
reagents from individual microneedles of the array with high
precision and temporal resolution. Analytical modeling is described
herein, which can offers prediction of the volumetric flow rate
through a single microneedle, and accordingly, can be used to
assist in the design of the microneedle arrays.
[0111] In some examples, conducting polymers such as polypyrrole
(PPy), polyaniline (PANI), and poly(3,4-ethylenedioxythiophene)
(PEDOT) can be employed for utilization in the described controlled
release systems and drug delivery actuators. These exemplary
materials include unique properties (e.g., PPy in particular),
which include their reversible mechanical behavior as "artificial
muscles" and their ability to change porosity and undergo volume
changes in response to applied electrochemical stimuli. The
disclosed actuator technology engineers these exemplary materials
in devices and systems to provide a means to deliver medications in
an effective and minimally-invasive manner, e.g., which can be
implemented in practical body-worn devices for the amelioration of
disease and injury in the acute phase, amenable to extended
durations of pain-free wear.
[0112] FIG. 12A shows a schematic illustration of an exemplary
microneedle-based multi-channel, multiplexed drug delivery actuator
device 1200. The device 1200 includes a hollow microneedle array
1201. The device 1200 includes a gold-sputtered polycarbonate
membrane 1204, which can be functionalized with sodium
dodecylbenzenesulfonate-doped polypyrrole (PC/Au/PPy/DBS). The
device 1200 includes a polydimethylsiloxane (PDMS) reservoir 1207
that can include multiple reservoirs to store chemical
(therapeutic) agents.
[0113] FIG. 12B shows a schematic illustration of the assembled
multiple-channel drug delivery actuator device 1200, in which the
reservoir 1207 is configured to have two exemplary reservoirs
containing two different drugs, e.g., drug 1211 and drug 1212. The
schematic in FIG. 12B also shows the assembled multiple-channel
drug delivery actuator device 1200 including electrode connections
1216 and 1217 corresponding to the two reservoirs containing the
drug 1211 and 1212, respectively.
[0114] The disclosed biosensor-actuator device enables the
controlled and switchable delivery of multiple therapeutic agents.
Exemplary implementations of the described device utilized
still-frame imaging and real-time video capture to show the
alternating release of dye from different microneedles from the
same array platform. For example, image analysis software (e.g.,
such as ImageJ) and ultraviolet-visible (UV-Vis) spectrophotometry
were employed to demonstrate the switching accuracy and
repeatability of the microneedle volumetric flow rate. These
exemplary results were correlated with an analytical model that
assesses the fluid flow characteristics through a single
microneedle, which can subsequently be used to assist in the design
and development of other embodiments of the disclosed multi-section
microneedle arrays technology, e.g., which can be applied for
practical body-worn devices that can deliver on demand different
therapeutic agents.
[0115] By employing the disclosed biosensor-actuator technology, a
unique drug therapy can be released at each microneedle constituent
of the array, thereby enabling custom-tailored dosages of
medications. The described biosensor-actuator technology includes
an active solid-state device that requires no moving parts or
integrated microelectromechanical systems (MEMS). Thus, this
simplifies low-profile device design and eliminates the need for
sophisticated microfluidics-based components, which can complicate
system architecture and increase both size and cost. Additionally,
for example, the described microneedle multi-drug delivery
technology can be implemented with implantable devices, and thus is
well-positioned to serve as the core component in an autonomous
`wearable nanopharmacy` in connection to multiplexed microneedle
sensor arrays.
[0116] Exemplary materials and methods to implement the disclosed
embodiment of the technology are presented. The following chemicals
and reagents were used in the described implementations, which
included sodium dodecylbenzenesulfonate (NaDBS), methylene green
(MG), chresol red (CR), potassium phosphate monobasic
(KH.sub.2PO.sub.4), and potassium phosphate dibasic
(K.sub.2HPO.sub.4), e.g., which were obtained from Sigma Aldrich
(St. Louis, Mo.) and were used without further purification or
modification. Pyrrole was distilled daily under vacuum and stored
at 4.degree. C. prior to electropolymerization. All reagents were
prepared in a 0.1 M phosphate buffer solution (pH 7.00). Ultrapure
water (18.2 M.OMEGA.cm) was employed in all of the exemplary
implementations. Polydimethylsiloxane (PDMS) was obtained from Dow
Corning Corp. (Midland, Mich.) and mixed by hand in a 10:1 polymer:
fixing agent ratio. The suspension was then poured into a custom
mold and degassed in a vacuum desiccator. Subsequently, the PDMS
suspension was baked at 110.degree. C. for 15 min. The resultant
structures were exposed to UVO ozone (Jetline Co., Irvine, Calif.)
at a gas flow rate of 3 sccm for 5 minutes. 25 mm-diameter black
polycarbonate (PC) track etch membrane filters were procured from
SPI Supplies (West Chester, Pa.); these filters possessed a pore
diameter of 600 nm.
[0117] The instrumentation used in the described implementations
included the following, which was utilized in exemplary
demonstrations and implementations of the disclosed embodiment
under exemplary conditions disclosed herein. A CH Instruments
(Austin, Tex.) model 1232A electrochemical analyzer was employed
for all of the electrochemical measurements. An Ag/AgCl wire
reference electrode and a platinum wire counter electrode were used
to establish a three-electrode electrochemical system. A Shimadzu
(Kyoto, Japan) UV-2450 UV-VIS spectrophotometer was used for all of
the optical measurements. A consumer digital video camera/camcorder
was employed to capture the still-frame images and videos. A
Philips XL30 field emission scanning electron microscope
(Amsterdam, the Netherlands) was employed to investigate the
surface morphology of the microneedle array. The arrays were coated
with a gold film (e.g., .about.15 nm) using an Emitech (East
Sussex, UK) K575X sputtering instrument prior to SEM imaging. The
resultant electron micrographs are shown in FIGS. 13A and 13B. FIG.
13A shows an SEM image detailing the surface morphology of an
exemplary hollow microneedle array. FIG. 13B shows an enhanced view
of the scanning electron micrograph of a single needle of the
exemplary hollow microneedle array featuring a well-defined
cylindrical lumen. The PC membranes were sputtered with a gold thin
film (.about.75 nm) using an Emitech K575X sputtering instrument
prior to the deposition of the NaDBS-doped PPy conducting
polymer.
[0118] The fabrication of the exemplary hollow microneedle arrays
used in the described implementations was performed in the
following manner. The hollow microneedle arrays were fabricated, in
which, the microneedle designs were originally prepared using
Solidworks (Dassualt Systemes S.A., Velizy, France). Substrate
support structures were subsequently created with Magics RP 13
(Materialise NV, Leuven, Belgium). For example, the hollow needles
were pyramidal in shape with a triangular base. For example, the
dimensions of each hollow microneedle were as follows: an edge
length of 1174.+-.13 .mu.m, a height of 1366.+-.15 .mu.m, and a
vertical cylindrical bore of 342.+-.5 .mu.m diameter on one of the
faces of the pyramid structure. The hollow needles were arranged
into 3.times.3 square arrays with 2 mm periodicity. For example,
substrates for the microneedle arrays were 10 mm.times.10 mm in
extent and possessed thickness values of 500 .mu.m.
[0119] The preparation of the exemplary electrically-actuatable
nanoporous membranes (e.g., PC/Au/PPy/DBS membranes) used in the
described implementations was performed in the following manner.
For example, gold-sputtered PC membranes (PC/Au) (e.g., pore
diameter.about.600 nm, porosity.about.0.2) were attached at the
periphery to a copper wire using silver conductive epoxy. A
solution of 0.1 M NaDBS was purged with nitrogen for 40 min after
which the pyrrole monomer was added to achieve a final
concentration of 0.25 M. Subsequently, the PC/Au membrane was
immersed in the solution and served as the working electrode in an
electrochemical cell while 0.6 V vs. Ag/AgCl was applied for 10
min. The application of this exemplary potential for the given
amount of time resulted in optimal deposition of the polypyrrole
polymer on the PC/Au membrane, thereby minimizing the leaching of
the solution through the membrane under the `closed` state while
enabling the solution to flow at appreciable rates under the `open`
state. Following electropolymerization of polypyrrole/DBS
(PPy/DBS), the PC/Au/PPy/DBS membranes were rinsed with deionized
water and stabilized by cycling between -1.1 V and 0.5 V vs.
Ag/AgCl for ten iterations in the buffer solution. This process
enabled the membrane to swell in the reduced state (-1.1 V) and
contract in the oxidized state (0.5 V) in a reversible manner. When
not in use, the membranes can be stored in the buffer solution at
room temperature.
[0120] The fabrication of the exemplary drug delivery actuator
contingent of the disclosed biosensor-actuator device used in the
described implementations was performed in the following manner.
The PC/Au/PPy/DBS membranes were cut into slivers possessing
dimensions of approximately 12 mm.times.4 mm. These slivers were
subsequently affixed to the reverse side of the exemplary 3.times.3
microneedle array using adhesive epoxy such that one sliver
completely covered a column of three microneedles. The center
column of the array was obstructed using modeling clay, enabling
formation of two individually-addressable electrically-actuatable
channels, exemplified in the component-level schematic illustrated
in FIG. 12A. Electrical leads were attached using silver epoxy to
each of the two PC/Au/PPy/DBS membranes to facilitate ohmic contact
with each actuator. The PDMS dual-channel reservoir was
subsequently aligned over the membranes and affixed using adhesive
epoxy. As shown in FIG. 12B, the reservoirs were finally loaded
with .about.20 .mu.L of the model chemical agent(s).
[0121] Initial implementations of the exemplary microneedle array
actuator device were aimed atvalidating and visualizing the
switching capability of the PC/Au/PPy/DBS membrane and the
dual-channel operation. For example, both reservoirs in the
assembled multiplexed drug delivery actuator, e.g., reservoir 1
(R1) and reservoir 2 (R2), were initially loaded with 12 mM of
methylene green (MG) dye and immersed in a buffer solution along
with the counter and reference electrodes. Continuous agitation at
a constant speed (e.g., 140 rpm) was applied with a magnetic
stirring bar. The DBS-doped PPy membrane entered the reduced state
and engorged upon biasing with -1.1 V vs. Ag/AgCl, thereby
obstructing the flow of the solution through the porous material.
Ejection of the MG dye at either channel was not observed at this
potential (represented as being in the `OFF` state), as shown in
image (A) of FIG. 14. Subsequently, the R2 membrane nanoactuator
was maintained at the reduced state (-1.1 V vs. Ag/AgCl, `OFF`) and
the membrane at R1 was switched to the oxidized state (`ON`) by
applying a potential of 0.5 V vs. Ag/AgCl. This "ON" state caused
the DBS-doped PPy membrane to become oxidized and contract, thereby
facilitating the flow of the solution through the nanoporous
membrane and subsequently through the microneedles. As can be
observed from the image (B) of FIG. 14, the emission of MG from R1
is visible whereas R2 remained closed and did not permit the
release of the dye. Following this operation, R1 was kept at the
oxidized state (0.5 V vs. Ag/AgCl, `ON`) while R2 was switched to
the oxidized state (0.5 V vs. Ag/AgCl, `ON`), thus releasing MG
from both reservoirs (as shown in image (C) of FIG. 14).
Subsequently, R1 was switched to the reduced state "OFF" and R2 was
kept at the oxidized state "ON", as shown in image (D) of FIG. 14.
This controlled and alternating release of MG from the individual
reservoirs by switching potentials on the nanoporous membranes was
illustrated in a real-time manner. The execution of repeated
`ON-OFF` cycles demonstrates that the drug delivery array maintains
its ability to open and close in a cyclic fashion, e.g., even
following 10 iterations or more. Furthermore, the temporal duration
(.about.30 s) required to observe the release of MG at the tenth
cycle was identical to that of the first cycle. In the exemplary
implementations, the time duration for complete flow shutoff was
approximately 35 s following the application of the "OFF"
potential. Based on the above results, R1 was loaded with CR dye
and R2 was loaded with MG dye. All four "ON/OFF" permutations were
applied. The controlled ejection of dye from alternating
microneedle array reservoirs was demonstrated based on the
potential applied to each nanoporous membrane.
[0122] The exemplary implementations included image analysis and
UV-Vis spectrophotometry techniques to analyze the drug delivery
capability of the microneedle array by experimentally quantifying
the flow rate of the MG dye from a single microneedle channel. FIG.
15 illustrates the release of MG from a single microneedle into a
quiescent buffer solution at fixed time intervals of 30 s. FIG. 15
shows exemplary time-lapse still frame images of the release of
methylene green (MG) from a single microneedle at distinct time
intervals of 30 s (shown in image (A)), 60 s (shown in image (B)),
90 s (shown in image (C)), and 120 s (shown in image (D)). For
example, a potential of 0.5 V (vs. Ag/AgCl) was applied to open the
nanoporous membrane and release the dye during the implementations.
The exemplary flow rate of released dye was determined to be
6.3.+-.0.4 .mu.L/hour (n=10) through analysis of multiple
time-lapse video still-frames. After 30 s of applying this
potential, the dye began to emerge from the microneedle aperture. A
small column of dye was clearly observed at 60 s. A well-defined
column of dye possessing a height of approximately 0.5 cm was
observed after 120 s. Afterwards, the estimated experimental flow
rate of the released dye was calculated by measuring its column
height (h) with image processing software (e.g., ImageJ) in
conjunction with the flow rate equation (Eq. 1):
Q = .pi. .times. .times. d 2 .times. h 4 .times. ( t - t 0 ) ( 1 )
##EQU00001##
where d is the microneedle channel diameter and h is the column
height associated with a particular point at time t.
[0123] UV-Vis spectrophotometry was employed to quantify the amount
of released dye and subsequently assess the microneedle flow rate
as well as the repeatability of the release. FIG. 16 shows an
exemplary UV-Vis spectrum data plot illustrating the absorbance for
the release of methylene green (MG) dye from a single microneedle
at a 2 minutes release interval over a 20 minute period. The upper
(top left) inset data plot displays the UV-Vis spectra, and the
lower (bottom right) inset data plot displays absorbance of 10
distinct experimental implementations over a constant time release.
The lower (bottom right) inset data plot in FIG. 16 substantiates
the reproducibility of the MG release from the drug-delivery
nanoactuator over the same release time. The maximum deviation
among these ten repetitions was 5.5% from the original absorbance,
which was measured at the maximum wavelength. Linear regression
analysis was performed on the absorbance vs. time plot, yielding a
slope of 3.5 mOD min.sup.-1 with a high coefficient of
determination (R.sup.2=0.993) and low relative standard deviation
(RSD=2.74%, n=3); this result indicated a constant release of dye
from the microneedle. From these implementations, the fluid flow
rate was calculated to be 5.5.+-.0.2 .mu.L/hour, which is in good
agreement with the image analysis data collected from the
time-lapse video still-frames. The fabricated membranes exhibited
excellent reproducibility, e.g., calculated flow rates deviated by
less than 10% under identical electropolymerization conditions.
[0124] The understanding of the fluid flow characteristics of the
microneedle array is important for delivering the precise amount of
drug to subcutaneous tissue during transdermal drug delivery. In
some examples, to augment this understanding and to analytically
estimate the drug delivery capability, the fluid flow
characteristics of a single microneedle can be modeled via the
Modified Bernoulli Equation (Eq. 2):
P 1 .rho. .times. .times. g + V 1 2 2 .times. g + z 1 = ( P 2 .rho.
.times. .times. g + V 2 2 2 .times. g + z 2 ) + .times. f .times. L
D .times. V 2 2 2 .times. g + .times. KV 2 2 2 .times. g ( 2 )
##EQU00002##
where P.sub.1 and P.sub.2 are the atmospheric and microneedle
outlet pressure, V.sub.1 and V.sub.2 are the average fluid
velocities, z.sub.1 and z.sub.2 are the heights at the top of the
reservoir and microneedle outlet respectively, f is the friction
factor, .rho. is the fluid density, L is the channel or pore
length, and D is the hydraulic diameter. FIG. 17 shows an exemplary
schematic of a single microneedle 1700 during drug delivery. The
schematic shows the following exemplary microneedle components
including: a reservoir 1701 (e.g., which can store a chemical
agent, such as a drug), a lumen structure 1702 (e.g., which can be
a duct or cavity of a tubular structure, sized to a 342 .mu.m
diameter), a hollow microneedle structure 1703, an
electrically-actuatable nanoporous membranes 1704 (e.g., Au/PPY/DBS
nanoporous membrane), a PC membrane 1705, and the released chemical
agent exiting the lumen 1706.
[0125] The second term in Eq. 2 refers to the friction losses
through the actuating nanopores, polycarbonate membrane, and
microneedle channel, as shown in expanded form (Eq. 3):
.times. f .times. .times. L D = f pores .times. L pores D .times.
.tau. pores pores + f membrane .times. L membrane D .times. .tau.
membrane membrane + f microneedle .times. L microneedle D ( 3 )
##EQU00003##
where .tau. and represent the tortuosity and porosity of the
nanopores and polycarbonate membranes, respectively and D is the
diameter of a single microchannel (e.g., 342 .mu.m). The porosity
of the PC membrane 1705 can be configured to 0.1 and the porosity
of the actuating nanopores can be configured to 0.4, e.g., due to
the pore narrowing created by the Au/PPy/DBS functionalization. For
example, an approximate tortuosity value of 1.5 was assigned to the
PC membrane and the actuating nanopores to take into account the
increased channel curvature created by the nanopores. The
respective friction factors were calculated according to Stokes
flow theory for water flow in microchannels, where the product of
the friction factor and Reynolds number (fRe=64) utilized for
macroscale laminar flow in circular channels is employed. The
friction factors for each flow section can be obtained by the
Reynolds numbers obtained for fluid flow in each of the three flow
sections of the microneedle channel, as shown in Table 1. Table 1
shows dimensions and flow characteristics of a single microneedle
channel.
TABLE-US-00001 TABLE 1 Length Total Cross Flow Section (.mu.m)
Sectional Area Re Microneedle Channel 1366 A.sub.c 6 .times.
10.sup.-3 Polycarbonate Membrane 7 (0.2) A.sub.c 5 .times.
10.sup.-5 Nanoporous Membrane 0.75 (0.4) A.sub.c 9 .times.
10.sup.-5
[0126] The values presented in Table 1 are calculated according to
the following exemplary considerations. For example, A, is the
cross sectional area of the microneedle
[.pi.(D.sub.microneedle).sup.4/2]. The Reynolds number
(Re=.rho.VD/.mu.) was calculated using the density (.rho.=1000
kg/m.sup.3) and viscosity (.mu.=1.000 N s/m.sup.2) of water at room
temperature. The velocity used to estimate the Reynolds number for
fluid exiting the microneedle channel was determined apriori by
averaging the experimental velocities obtained by image analysis
and UV-Vis spectrophotometry. Furthermore, the apriori velocities
for the polycarbonate membrane and the nanoporous membrane were
obtained by utilizing Conservation of Mass for incompressible
fluids (V.sub.1A.sub.1=V.sub.2A.sub.2) in conjunction with the
average experimental velocity to calculate the corresponding
Reynolds numbers.
[0127] The last term (.SIGMA.K) in Equation (2) represents the sum
of minor losses due to the inlet, exit, and hydrodynamic
development length, which is shown in expanded form below:
.SIGMA.K=K.sub.inlet+K.sub.outlet (4)
where K.sub.inlet and K.sub.outlet are loss coefficient factors for
a square edge inlet (0.5) and for an exit (1) typically associated
with hollow microneedles.
[0128] An expression (Eq. 5) for the theoretical flow rate of the
fluid exiting the microneedle channel can be formed by assuming
quiescent flow at the top of the reservoirs (V.sub.1=0), and
negligible pressure gradients throughout the flow network
(.DELTA.P=P.sub.1-P.sub.2=0):
Q 2 = A c .times. 2 .times. g .function. ( z 1 - z 2 ) .times. f
.times. .times. L D + K ( 5 ) ##EQU00004##
[0129] The theoretical flow rate calculated by Eq. 4 and the
experimental flow rates obtained through image analysis and UV-Vis
spectrophotometry were in good agreement, e.g., validating the
veracity of the microneedle fluid flow model presented herein, as
shown in Table 2. Table 2 exhibits a comparison of calculated
theoretical and experimental microneedle flow rates.
TABLE-US-00002 TABLE 2 Flow Rate (Q.sub.2) (.mu.l/hr) Theoretical
Model 6.4 Image Analysis 6.3 .+-. 0.4 UV-Vis Spectrophotometry 5.5
.+-. 0.2
[0130] The ability to transdermally release multiple drugs may be
important for the autonomous treatment of metabolic syndromes
(e.g., a combination of hypertriglyceridemia, hypertension, and
insulin dependent diabetes mellitus), human immunodeficiency virus,
and other chronic medical conditions. The disclosed embodiment
presents a self-contained multiplexed drug delivery system that
utilizes arrays of microneedles coupled with conducting polymer
nanoactuators for the controlled release of fluidic agents. The
ability of the exemplary PPy/DBS conducting polymer to undergo
volumetric changes with applied electrical potentials permits the
release of fluid in a controlled and switchable fashion, e.g.,
without the need for moving parts or integrated
microelectromechanical systems. These nanopore-actuated microneedle
arrays are well suited for integration into wearable drug delivery
devices, in which cost and power constraints must be minimized.
[0131] For example, a method to sense an analyte and deliver a
therapeutic agent is described, e.g., which can be implemented
using the described devices and systems of the disclosed
embodiment. The exemplary method can include a process to detect a
signal produced by an analyte at an interface with a functionalized
probe configured to electrochemically interact with the analyte
within a biological fluid, in which the signal is transduced to an
electrical signal by the functionalized probe. For example, the
functionalized probe can be one of an array of multiple
functionalized probes, and the functionalized probe can be
chemically functionalized to interact with one or more target
analytes in the fluid. For example, the biological fluid can
include at least one of transdermal fluid, intraocular fluid,
vitreous humor, cerebrospinal fluid, extracellular fluid,
interstitial fluid, plasma, serum, lacrimal fluid, saliva,
perspiration, mucus, or blood, among other biological fluids in a
living organism. The exemplary method can include a process to
process (e.g., implementing signal processing techniques) the
electrical signal to determine a parameter of the analyte (e.g.,
such as the concentration of the analyte). The exemplary method can
include a process to, e.g., based on the determined parameter,
apply an electrical stimulus to a valve (e.g., in which the valve
can be a porous polymer film having pores of a reversibly tunable
porosity, as described herein), in which the valve attached to a
container containing a therapeutic agent. The exemplary method
includes the electrical stimulus altering the permeability of the
pores, e.g., from a closed state to an open state, thereby
releasing the therapeutic agent into the biological fluid. For
example, the therapeutic agent can include, but is not limited to,
a drug, vaccine, hormone, vitamin, anti-oxidant, or pharmacological
agent.
[0132] The disclosed multiple-drug delivery microsystem can be
integrated with the described microneedle sensor array, e.g.,
coupling multiplexed analyte detection with the corresponding
therapeutic intervention. This can enable a closed-loop
sensing/drug delivery microneedle paradigm that is well-positioned
to precisely deliver multiple therapeutic agents in an on-demand
basis. This type of autonomous "Sense-Act-Treat" system, devices,
and methods can provide an avenue for responding to biomarker
fluctuations with a targeted therapy, as well as provide
self-regulating drug delivery that can adjust patient dosage based
on the severity of the injury or the disease process. The
development of such responsive multiplexed drug-delivering systems
can be implemented to transform outpatient, home-based civilian
medical treatments as well as military medical care.
[0133] In another embodiment of the disclosed technology, a
minimally-invasive multi-component microneedle device with carbon
paste electrodes within a hollow microneedle array for
electrochemical monitoring and biosensing, which can be fabricated
using a digital micromirror device-based stereolithography
techniques, is described. This embodiment can comprise the same
embodiment(s) like those previously described, and can therefore
implement the entirety of functionalities of the individual
embodiments on a single embodiment. In this embodiment, a rapid
prototyping method to fabricate exemplary microneedle
biosensor-actuator devices is described that uses a dynamic mask
(e.g., such as a Digital Micromirror Device (DMD)). In some
examples, the exemplary method can employ the dynamic mask for
selective polymerization of a photosensitive acrylate-based polymer
resin into an exemplary microneedle sensor-actuator device.
Exemplary implementations were performed that demonstrated that the
hollow microneedles remained intact after puncturing the outermost
layer of skin in a living organism. For example, in these exemplary
implementations, the carbon fibers underwent chemical modification
in order to enable detection of hydrogen peroxide and ascorbic
acid; electrochemical measurements were demonstrated using
integrated electrode-hollow microneedle devices. The disclosed
technology includes an approach for implementing real-time,
minimally invasive point-of-care sensing using an exemplary device
capable of obtaining biological samples (e.g., interstitial fluid)
through the skin while protecting the sensing transducer from
biofouling elements. Such devices can be used as in vivo sensors to
provide real-time detection of physiological processes, such as
monitoring of a neurotransmitters, medically-relevant molecules,
cancer biomarkers, and pathogenic microorganisms.
[0134] Exemplary materials to fabricate and implement the disclosed
embodiment of the technology are presented. For example, a variety
of materials can be used for microneedle fabrication, including
silicon, glass, metal (e.g., stainless steel and nickel), and
resorbable polymers (e.g., polyglycolic acid and polylactic acid).
In one example, an acrylate-based polymer, e.g., e-Shell 200, was
utilized for microneedle fabrication. The material is a Class-IIa
biocompatible, water-resistant material; it has been used in
thin-walled hearing aid shells, solid microneedle arrays, as well
as nonmedical applications. e-Shell 200 contains 0.5-1.5% wt
phenylbis(2,4,6 trimethylbenzoyl)-phosphine oxide photoinitiator,
15-30% wt propylated (2) neopentyl glycoldiacrylate, and 60-80% wt
urethane dimethacrylate. Energy-dispersive X-ray spectroscopy
indicated that e-Shell 200 contains carbon, oxygen, and titanium;
these elements are known to possess excellentbiocompatibility.
e-Shell 200 exhibits a water absorption value of 0.12% (D570-98
test method) and a glass transition temperature of 109.degree. C.
(E1545-00 test method). It exhibits tensile strength of 57.8 MPa
(D638M test method), flexural strength of 103 MPa (D790M test
method), and elongation at yield of 3.2% (D638M test method).
[0135] For example, a method of fabrication included use of a
Digital Micromirror Device-stereolithography instrument to
fabricate hollow microneedles and the integration of carbon fiber
electrodes within the bores of these hollow microneedles. The
carbon fibers can be chemically modified to enable detection of two
medically significant molecules, hydrogen peroxide and ascorbic
acid. Electrochemical characterization was performed on the
chemically modified electrode-hollow microneedle devices. For
example, hydrogen peroxide (H.sub.2O.sub.2) is a reactive oxygen
species that is monitored in many common enzyme-based
electrochemical sensors. For example, H.sub.2O.sub.2 and
gluconolacone are produced in reactions between glucose and glucose
oxidase; monitoring of released hydrogen peroxide is used for
quantification of glucose. For example, monitoring of released
hydrogen peroxide may also be used for quantification of glutamate
in brain dialysate; hydrogen peroxide is produced in reactions
between glutamate and glutamate oxidase. Glutamate is an excitatory
neurotransmitter, e.g., which has been linked with aggressive
activity. Ascorbic acid can be an indicator of oxidative stress
that is experienced by cells.
[0136] The following processes were implemented to demonstrate the
disclosed embodiment of a microneedle array device and fabrication
methods thereof. Exemplary implementations included proliferation
of human dermal fibroblasts and neonatal human epidermal
keratinocytes on e-Shell 200 surfaces, which was evaluated using
the MTT (3-(4,5-dimethylthiazol-2-yl)2,5-diphenyl tetrazolium
bromide) assay, e.g., which involves reduction of a yellow
tetrazolium salt to a purple formazan dye by mitochondrial succinic
dehydrogenase. In the described exemplary implementations, e-Shell
200 wafers (diameter=15 mm, thickness=2 mm) were compared against
glass cover slips (diameter=15 mm). The cover slips and e-Shell 200
wafers were rinsed and sterilized in two 30 minute rinses of 70%
ethanol; the materials were subsequently rinsed in sterile
deionized water. The e-Shell 200 wafers were placed in sterile
Petri dishes in a laminar flow cabinet and sterilized with
ultraviolet B light, e.g., both surfaces were exposed to
ultraviolet B light. The materials were rotated 90 degrees after a
minimum of two hours light exposure. Polymers were transferred to
sterile 24-well culture plates, rinsed twice in sterile Hank's
Balanced Salt solution, and once in the appropriate cell culture
medium. The e-Shell 200 wafers were placed in 2 mL of the
appropriate cell culture medium and held in the incubator until
seeded.
[0137] Cryopreserved neonatal human epidermal keratinocytes (HEK)
and human dermal fibroblasts (HDF) were obtained, and fibroblast
growth media (FGM-2) and keratinocyte growth media (KGM-2) were
also obtained (e.g., Lonza, Walkersville, Md.). The human dermal
fibroblasts and neonatal human epidermal keratinocytes were
propagated in 75 cm.sup.2 flasks, e.g., grown to 75% confluency,
and subsequently harvested. The cells were seeded (e.g.,
concentration=40,000 cells per well) in a 24-well plate on e-Shell
wafers 200 (n=4), glass cover slips (n=4), and polystyrene well
plates (n=4). Material rinsing and all media changes were performed
by moving the test materials from one solution to the other using a
forceps. The materials were placed in fresh medium after 48 hours;
this time point corresponded with 80% confluency for human dermal
fibroblasts and neonatal human epidermal keratinocytes. MTT
viability was assessed 24 hours later. The materials with cells
were rinsed using Hank's Balanced Salt solution; desorption using
isopropyl alcohol and agitation were subsequently performed.
Isopropyl alcohol (e.g., quantity=100 .mu.L) was transferred to a
new 96-well plate. Absorbance was determined (e.g., X=550 nm) with
a Multiskan RC plate reader (Labsystems Inc., Franklin, Mass.). The
mean values for percent viability were calculated. Significant
differences (p<0.05) were determined using the PROC GLM
Procedure (SAS 9.1 for Windows) (SAS Institute, Cary, N.C.). When
significant differences were found, then multiple comparisons were
performed using Tukey's Studentized Range HSD (Honestly Significant
Difference) test at p<0.05 level of significance.
[0138] Arrays of hollow microneedles were fabricated from
three-dimensional drawings that were created using Solidworks
(Dassualt Systemes S.A., Velizy, France). Support structures were
fabricated from three-dimensional drawings that were created using
Magics RP 13 (Materialise NV, Leuven, Belgium). In the
tetrahedron-shaped microneedle design, two faces of the microneedle
exhibit a vertical orientation with respect to the substrate. The
microneedle input dimensions included a triangular base with 1.2 mm
sides, a height of 1.5 mm, and a vertical cylindrical channel
(diameter=400 .mu.m). The needles were arranged into a 2.times.2
square array with 2 mm inter-microneedle spacing. The substrate
input dimensions included lateral dimensions of 1 cm.times.1 cm and
a thickness of 500 .mu.m. Rapid prototyping of the microneedle
array was performed using a Perfactory III SXGA+instrument
(EnvisionTEC GmbH, Gladbeck, Germany). A 150-W halogen bulb was
used as the light source for polymerization of liquid e-Shell 200
resin. Selective polymerization of the resin in the X-Y plane was
achieved using Digital MicrornirrorDevice (DMD) optics (Texas
Instruments, Dallas, Tex.), specifically a DMD SXGA+guidance chip
with 1280.times.1024-pixel resolution. This instrument contains a
build envelope of 90 mm.times.67.5 mm. After fabrication, the
microneedle array was washed in isopropanol in order to remove
unpolymerized material. Post-building curing was accomplished using
an Otoflash Post Curing System instrument (EnvisionTEC GmbH,
Gladbeck, Germany), which contains two photo-flash lamps and
provides light exposure over a wavelength range of 300-700 nm.
[0139] A Hitachi 5-3200 (Hitachi, Tokyo, Japan) variable pressure
scanning electron microscope with a Robinson backscattered electron
detector was used for imaging the exemplary microneedle arrays. The
exemplary microneedle arrays were coated with 60% gold-40%
palladium using a Technics Hummer II instrument (Anatech, Battle
Creek, Mich.) prior to imaging. Skin penetration testing was
performed with full-thickness cadaveric porcine skin since human
skin and porcine skin exhibit similar structures. Trypan blue
(Mediatech, Inc., Manassas, Va.), a toluidine-based dye, was used
to assess the transdermal drug delivery functionality of the hollow
microneedle arrays. Cadaveric full-thickness weanling
Yorkshire/Landrace skin was stored at 3.degree. C. until testing
was performed. Hollow microneedle arrays were inserted into
full-thickness porcine skin. After removal of the arrays, Trypan
blue was applied to the insertion site; the site was subsequently
washed with isopropanol swabs. The Trypan blue-treated skin was
subsequently imaged using optical microscopy. Images of a
microneedle device before insertion into porcine skin and after
insertion into porcine skin were obtained using optical
microscopy.
[0140] FIGS. 18A-18D show illustrative schematics showing
processing steps for assembly of an exemplary microneedle array
device of the disclosed embodiment. In this embodiment, the
disclosed microneedle array device includes two layers. An upper
layer includes the microneedle array, and the lower layer provides
support for the carbon fibers and facilitates alignment of the
carbon fibers to the microneedle array. For example, the support
component can be fabricated from a 1.6 mm thick
poly-(methylmethacrylate) (PMMA) piece. An array of holes can be
laser drilled through the PMMA piece using a Model PLS instrument,
including a 6.75 60-watt CO.sub.2 laser and a computer-controlled
XY stage (Universal Laser Systems, Scottsdale, Ariz.), as shown in
FIG. 18A. The holes can be placed in a square pattern with 2 mm
spacing. For example, using a Model HPDFO high power density
focusing optics lens (Universal Laser Systems, Scottsdale, Ariz.),
the diameter of the exemplary hole at the exit surface was measured
at -45 .mu.m. For example, to control the carbon fiber length
beyond the support surface, the support component was placed on top
(exit-side down) of a well with a depth of 100 .mu.m. The carbon
fiber can be inserted into each of the holes (entrance-side) and
allowed to rest at the bottom of the well, as shown in FIG. 18B.
The fibers can be secured in place with acrylic adhesive on the
entrance side after a desired well depth has been achieved, as
shown in FIG. 18C. FIGS. 19A and 19B show optical images of an
array of carbon fiber electrodes and a single carbon fiber
electrode in focus, respectively. The support layer and microneedle
layer can be brought together in such a manner that the carbon
fibers are positioned within the hollow shafts of the microneedles.
The layers can be subsequently adhered to each other. For example,
metallic epoxy can be applied to the back of the fibers in order to
create the connection for the working electrode, as shown in FIG.
18D.
[0141] The exemplary implementations included 7 .mu.m carbon fibers
(Alfa Aesar, St. Louis, Mo.) that were activated in a KOH solution
(concentration=0.1 M) at a pH of 13 and at a potential of 1.3 V for
five minutes. In situ diazotation of 2-amino-4-nitrophenol was
performed by mixing a solution of 8 mM sodium nitrite and 6 mM
2-amino-4-nitrophenol on ice for 5 minutes to create the
corresponding diazonium salt. After five minutes, the activated
carbon fibers were inserted. Two cyclic voltammetry (CV) scans were
run from 0.4 V to -0.8 Vat 0.1 V/s to enable electrochemical
grafting of the 2-nitrophenol and subsequent reduction to the
aminophenol. The carbon fibers were modified with palladium to
enable detection of hydrogen peroxide. Activated carbon fiber
bundles were placed in a solution of 1 mM palladium (II) chloride;
Pd was deposited by applying a potential of -0.8 V for 120 s. The
electrochemical measurements were obtained using a PGSTAT12 Autolab
electrochemical instrument (EcoChemie, Utrecht, the Netherlands).
Data was acquired versus an Ag/AgCl reference and a Pt counter
electrode.
[0142] The Digital Micromirror Device-based stereolithography
instrument was employed for the fabrication of approximately 200
arrays over a three-hour period. FIGS. 20A and 20B show SEM images
of an exemplary hollow microneedle array and an exemplary single
hollow microneedle of the disclosed embodiment prior to
incorporation of carbon fiber electrodes, respectively.
Measurements obtained from the SEM images showed that the exemplary
microneedles exhibited heights of .about.1030 .mu.m, triangular
bases with side lengths of .about.1120 .mu.m, and bore diameters of
.about.375 .mu.m. Good microneedle-to-microneedle uniformity was
noted among the microneedles in the microneedle array. In this
exemplary implementation, it is noted that differences between
input and measured dimensions may be attributed to translation of
the computer-aided design drawing by the software. For example,
Digital Micromirror Device-based stereolithography and other rapid
prototyping techniques involve tessellation, conversion of the
surface of the computer-aided design drawing into a series of
polygons. This polygon series is converted into a series of
cross-sectional layers, which is subsequently used for
layer-by-layer fabrication of the microneedle device. It is not
possible to predict how the computer-aided design drawing is
manipulated by the software.
[0143] For example, microneedles undergo bending forces,
compressive forces, shear forces, and skin resistance during skin
insertion; the pressure necessary for human skin penetration can
exceed 3.0.times.10.sup.6 Pa. The skin penetration properties of
the microneedle devices were evaluated using cadaveric porcine
skin, which has been previously used as a model for assessing
microneedle functionality. FIG. 21A shows an image of porcine skin
after application of the microneedle array, removal of the
microneedle array, and application of Trypan blue. The Trypan blue
spots indicate penetration through the stratum corneum layer
(outermost layer) of the epidermis by the microneedle array and
localization of Trypan Blue within microneedle-generated pores.
FIG. 21B and FIG. 21C show optical micrographs of hollow
microneedles before insertion into porcine skin and after insertion
into porcine skin, respectively. These exemplary images indicate
that the microneedles remain intact after skin insertion.
[0144] For example, the positioning of the exemplary carbon fiber
electrodes within the microneedle device can facilitate
interactions with the biological sample and minimize carbon fiber
exposure to stresses associated with microneedle insertion into
skin and movement at the microneedle device-skin interface. To
facilitate interactions between the biological sample and the
carbon fiber electrodes, the carbon fiber electrodes can be
positioned at the centers of the microneedle bores. In addition,
dead space between the carbon fiber electrodes and the microneedle
sidewalls may allow for infiltration of the biological sample. FIG.
22A shows an SEM image of a hollow microneedle array, and FIG. 22B
shows an SEM image of a single hollow microneedle after
incorporation of carbon fiber electrodes. The exemplary SEM image
of FIG. 22B reveals that the carbon fiber electrodes do not extend
beyond the tip of the microneedle bore. Placement of carbon fibers
within the microneedle bores included precise alignment of the
microneedle bores and the carbon fibers, e.g., the positions of the
laser-ablated holes in the lower layer of the microneedle device
were coordinated with the positions of the microneedle bores in the
upper layer of the microneedle device.
[0145] The exemplary implementation included the evaluation of the
electrochemical response of the exemplary carbon fibers within the
electrode-hollow microneedle device towards 5 mM
Fe(CN).sub.6.sup.3-/4-/1 M KCl. FIG. 23 shows a data plot of a
cyclic voltammetric scan of 5 mM ferricyanide in 1 MKCl versus
Ag/AgCl and Pt reference counter electrodes, respectively, e.g., at
a scan rate of 100 mV/s. For example, well defined
oxidation/reduction waves were observed, indicating interaction
between the carbon fiber electrodes and the test solution as a
result of permeation of the microneedle bore by the test solution.
The average formal potential (E.sup.0) for Fe(CN).sub.6.sup.3-/4-
was measured at 220 mV vs. Ag/AgCl reference and platinum counter
electrodes, respectively. The average peak separation was
.DELTA.E.sub.p=125 mV. These exemplary results indicate that the
carbon fibers within the electrode-hollow microneedle device were
capable of providing electrochemical measurements.
[0146] The exemplary implementation included the evaluation of
palladium-catalyzed oxidation of hydrogen peroxide on the carbon
fibers within the exemplary electrode-hollow microneedle devices.
For example, palladium was deposited onto the carbon fibers by
applying a potential of -0.8V for 120 sec in 1 mM Pd/0.5M HCl prior
to insertion into the microneedle device. FIG. 24 shows a data plot
of cyclic voltammetric scans of 0, 50, 100, 300, and 500 .mu.M
hydrogen peroxide, e.g., as shown represented by pink, black,
green, blue, and red curves, respectively, versus Ag/AgCl and Pt
reference counter electrodes, respectively, at a scan rate of 100
mV/s. This exemplary data shown in FIG. 24 shows an increase in
reductive currents after additions of 0, 50, 100, 300, and 500
.mu.M hydrogen peroxide, exhibiting a linear range of 100-500 .mu.M
and a detection limit of .about.15 .mu.M (based on the response of
50 .mu.M hydrogen peroxide; S/N=3).
[0147] For example, the carbon fibers were modified with
aminophenol (o-AP) groups following in-situ diazotination and
electrografting of the corresponding diazonium salt. Modification
of the carbon fiber electrode with o-AP can result in
electrocatalytic oxidation of ascorbic acid and selective oxidation
of ascorbic acid in the presence of common interferents, e.g., such
as uric and citric acid. Uric acid is a well-known interferent in
electrochemical analysis of ascorbic acid, e.g., which can be
attributed to the fact that uric acid and ascorbic acid possess
similar oxidation potential values. Linear sweep voltammograms of
100 mM phosphate buffer (blank solution) and 1 mM ascorbic acid in
100 mM phosphate buffer (pH=7) versus Ag/AgCl and Pt reference
counter electrodes, respectively, at a scan rate of 100 mV/s are
shown in FIG. 25. This exemplary result indicates that the carbon
fibers within the electrode-hollow microneedle device are able to
detect the ascorbate analyte with the low potential oxidation of
ascorbic acid at 195 mV. Electrochemical measurements by the carbon
fibers within the electrode-hollow microneedle device of the
disclosed embodiment were demonstrated. In addition, chemical
modification of the exemplary carbon fibers for selective analytes
was shown, and detection of hydrogen peroxide and ascorbic acid
using these modified carbon fibers was demonstrated.
[0148] In another embodiment of the disclosed technology, a
minimally-invasive multi-component microneedle device with carbon
paste electrodes (CPEs) for electrochemical monitoring and
biosensing is described. This embodiment can comprise the same
embodiment(s) like those previously described, and can therefore
implement the entirety of functionalities of the individual
embodiments on a single embodiment. The exemplary carbon paste
electrodes can exhibit a renewable nature or functionality that
enables the packing of the exemplary hollow non-planar microneedles
with pastes that contain assorted catalysts and biocatalysts. For
example, smoothing the surface can result in
microelectrode-to-microelectrode uniformity. Optical and scanning
electron micrographs show the surface morphology at the microneedle
apertures. Exemplary implementations of the disclosed microneedle
electrode arrays included low-potential detection of hydrogen
peroxide at rhodium-dispersed carbon paste microneedles in vitro
and lactate biosensing by the inclusion of lactate oxidase in the
metallized carbon paste matrix. The exemplary implementations
demonstrated highly repeatable sensing, e.g., for following
consecutive cycles of packing/unpacking the carbon paste. For
example, the operational stability of the exemplary array was
demonstrated, as well as the interference-free detection of lactate
in the presence of physiologically relevant levels of ascorbic
acid, uric acid, and acetaminophen. The described microneedle
design can be well-suited for diverse biosensing applications,
e.g., including subcutaneous electrochemical monitoring of a number
of physiologically-relevant analytes.
[0149] For example, carbon paste can be characterized by a high
degree of moldability that is essential for optimal packing and can
be employed in electroanalysis. CPEs can include the advantages of
low background current, low cost, as well as convenient surface
renewal and modification (e.g., via the inclusion of the modifiers
within the paste). Exemplary microneedle arrays of the disclosed
embodiment can include a nine-element arrays of pyramidal-shaped
hollow microneedles, which possess a 425 .mu.m-diameter aperture
through which the modified carbon paste is extruded and can act as
a transducer. For example, rhodium-dispersed carbon paste, which
can be used for low-potential detection of hydrogen peroxide, can
be packed within the microneedles to minimize the contribution of
co-existing electroactive interferents. The described microneedle
array CPE sensor device obviates the need for integrated
microchannels and extraction of the interstitial fluid.
[0150] Exemplary materials and methods to implement the disclosed
embodiment of the technology are presented. The following chemicals
and reagents were used in the described implementations, which
included lactate oxidase from Pediococcus sp. (LOx, E.C.
1.13.12.4), rhodium on carbon (5% Rh weight), polyethyleneimine
(PEI), mineral oil (e.g., d=0.838 g/mL), L-lactic acid, hydrogen
peroxide (H.sub.2O.sub.2), L-ascorbic acid (AA), uric acid (UA),
acetaminophen (AC), ethyl alcohol, potassium phosphate monobasic,
and potassium phosphate dibasic were obtained from Sigma Aldrich
(St. Louis, Mo.) and were used without further purification or
modification. All experiments were performed with 0.1 M phosphate
buffer (pH 7.0). Ultrapure water (e.g., 18.2 M.OMEGA.cm) was
employed in the exemplary implementations.
[0151] The exemplary solid and hollow microneedle arrays used in
the exemplary implementations were developed in the following
manner. The hollow microneedle arrays were fabricated with the aid
of Solidworks (Dassualt Systemes S.A., Velizy, France) computer
models. Substrate structures were designed with Magics RP 13
(Materialise NV, Leuven, Belgium). For example, the needles were
pyramidal in shape with a triangular base. For example, the
dimensions of each microneedle were as follows: an edge length of
1250 .mu.m, a height of 1500 .mu.m, and a vertical cylindrical bore
of 425 .mu.m in diameter on one of the faces of the pyramid
structure. The exemplary needles were arranged into 3.times.3
square arrays with 2 mm periodicity. Substrates for the microneedle
arrays were 10 mm.times.10 mm in extent and possessed a thickness
of 500 .mu.m. The three-dimensional computer models were
transferred to a Perfactory.RTM. SXGA Standard UV rapid prototyping
system (EnvisionTEC GmbH, Gladbeck, Germany) for production. This
system uses these computer models to precisely guide light from a
150 W halogen bulb over a photocurable material, resulting in the
selective polymerization of the exposed material. Eshell 200
acrylate-based polymer (EnvisionTEC GmbH, Gladbeck, Germany) was
utilized as the constituent material to fabricate the microneedle
arrays since the resin selectively polymerizes under visible light
and exhibits a Young's modulus of elasticity of 3050.+-.90 MPa. The
polymer also offers Class-IIa biocompatibility per ISO 10993. A 550
mW output power beam (e.g., step size=50 .mu.m) with a zero-degree
tilt was employed for the polymerization of the resin. Following
fabrication, the arrays were rinsed with isopropanol for removal of
the unpolymerized material and subsequently placed in an Otoflash
post curing system (EnvisionTEC GmbH, Gladbeck, Germany) for
post-build curing.
[0152] The exemplary enzyme-functionalized rhodium-dispersed carbon
paste microelectrode array was prepared in the following manner.
For example, 100 mg of Rh-on-carbon and 10 mg of LOx were
thoroughly homogenized via 10 alternating 5-min cycles of vortexing
and ultrasonication. The mixture was then vortexed for an
additional 1 hr. Following the homogenization process, 125 mg of
the mineral oil pasting liquid and 15 mg of the PEI enzyme
stabilizer were added to the solid mixture. Homogenization of the
resulting paste mixture was accomplished by grinding the mixture
with a mortar and pestle for an additional 1 hr.
[0153] For example, a 3 mL syringe (BD Biosciences, Franklin Lakes,
N.J.) can be utilized as the support to extrude the metallized
carbon paste through the microneedle array. The nozzle portion of
the syringe was removed to facilitate the attachment of the
microneedle array, which was affixed (e.g., using adhesive epoxy)
to this cleaved end for durability. A copper wire was subsequently
inserted into the back end of the syringe barrel in order to create
an electrical contact to the microneedle transducer. Following this
exemplary procedure, the carbon paste mixture was loaded into the
syringe from the back end and then extruded with a plunger until
the paste began to expel through the microneedle microholes. Excess
paste was removed from the openings; the surface was later smoothed
using wax paper. In order to investigate the repeatability of the
response after repacking the microneedles with new paste, the array
was carefully removed from the syringe and subsequently immersed in
ethanol under ultrasonication in order to remove the extraneous
carbon paste residue. A 0.15 mm diameter iridium wire was used to
facilitate removal of the paste from the microhole. The
aforementioned assembly and packing protocols were then followed in
order to generate a new electrode from the cleaned microneedle
array.
[0154] The instrumentation used in the described implementations
included the following, which was utilized in exemplary
demonstrations and implementations of the disclosed embodiment
under exemplary conditions disclosed herein. A CH Instruments
(Austin, Tex.) model 1232A electrochemical analyzer was employed
for electrochemical measurements. An external Ag/AgCl reference
electrode (CH Instruments CHI111) and a 0.5 mm diameter platinum
wire counter electrode were used to establish a three-electrode
electrochemical system. The electrochemical experiments were
performed in a 7 mL cell at room temperature (22.degree. C.).
Voltammetric and chronoamperometric studies were used to evaluate
the electrochemical behavior of the exemplary carbon paste
microneedle array electrode. In these electrochemical
implementations, either H.sub.2O.sub.2 or lactate was added into 5
mL of potassium phosphate buffer solution in order to obtain the
desired concentration. Chronoamperometric currents were sampled at
15 s following the potential step. In order to obtain hydrodynamic
voltammograms, fixed potential amperograms were recorded in a
stirred phosphate buffer solution containing the desired
H.sub.2O.sub.2 concentration by varying the potential between -0.20
and +0.60V vs. Ag/AgCl (e.g., in 0.05V increments). The solution
was continuously stirred using a magnetic stirrer at a rate of 100
rpm. The morphology of the carbon paste microneedle array was
examined using a field emission scanning electron microscope
(Philips XL30, Amsterdam, The Netherlands). All of the specimens
were coated with chromium prior to analysis using a sputtering
instrument (Energy Beam Sciences Emitech K575X, East Granby,
Conn.). A deposition current of 130 mA was applied for 30 s to
deposit-15 nm of chromium on the sample surface.
[0155] Exemplary implementations of the disclosed embodiment
included characterization of the surface morphology of the carbon
paste microelectrode array. For example, unmodified and modified
carbon pastes can readily conform with the non-planar features of
microneedle array devices. Initial implementations were aimed at
characterizing the morphology of the carbon paste-loaded
microneedle array and were initiated with a close examination of
the microelectrode surface. FIGS. 26A and 26B show optical
micrographs of the unpacked and Rh-carbon paste packed microneedle
array, respectively. An optical micrograph of the exemplary
unpacked microneedle array is shown in FIG. 26A. This image shows
uniform pyramidal microneedle structures (with triangular bases)
possessing a height of 1500 .mu.m as well as the cylindrical
openings (425 .mu.m diameter). FIG. 26B depicts an exemplary
microneedle array that has been packed with carbon paste and
subsequently polished. It indicates that the surface has been
smoothly polished to obtain a highly reproducible exposed area,
thereby facilitating reliable electrochemical sensing. As shown in
the figures, microelectrode-to-microelectrode uniformity can be
observed.
[0156] Pursuant to the characterization of the surface morphology,
SEM imaging of the microneedle was performed. FIGS. 27A and 27B
show SEM images of the unpacked and Rh-carbon paste packed
microneedle constituent of the array. FIG. 27A shows an electron
micrograph of a single microneedle. The structure of the
microneedle can be observed, e.g., the bored cylindrical vacancy
and the ribbed structure created by rastering of the light source
over the polymer resin. FIG. 27B shows the surface details of a
single microneedle packed with the carbon paste. For example, as
shown in the figure, a well-formed surface, a relatively smooth
morphology, and defined edges can be observed, e.g., reflecting the
effective filling of the cylindrical microhole. Such surface
quality can be achieved by extruding excess paste and later
polishing the surface. It should be noted that the microneedle and
the opening appear to be elongated due to the oblique angle at
which the SEM image was acquired.
[0157] Exemplary implementations of the disclosed embodiment
included electrochemical characterization of the carbon paste
microelectrode array towards peroxide-based amperometric sensing.
For example, initial electrochemical implementations were carried
out to characterize the response of the carbon paste microneedle
array to H.sub.2O.sub.2. A hydrodynamic voltammogram (HDV) was
recorded over the -0.20 to +0.60 V range in order to deduce a
suitable operating potential and to demonstrate the strong
catalytic ability of the Rh-CPE towards the redox processes of
H.sub.2O.sub.2. The exemplary results, shown in FIG. 28A, elucidate
that the Rh-CPE offers convenient detection of H.sub.2O.sub.2 over
the entire range tested, with a crossover point occurring around
0.22V (vs. Ag/AgCl). FIG. 28A shows plots of hydrodynamic
voltammograms of 0.1M potassium phosphate buffer (data plot (a))
and 10 mM H.sub.2O.sub.2 (data plot (b)) at the rhodium-dispersed
carbon paste microneedle electrode. For example, such lowering of
the overvoltage enables the selection of a low operating potential
of -0.15V vs. Ag/AgCl for subsequent sensor implementations. At
this exemplary potential, a reduction current of 5.95 .mu.A can be
achieved for 10 mM H.sub.2O.sub.2. Contributions imparted by common
electroactive interferences were negligible.
[0158] The exemplary microneedle CPEs display a wide dynamic range
for H.sub.2O.sub.2 detection. FIG. 28B shows plots of
chronoamperograms obtained using the rhodium-dispersed carbon paste
microneedle electrode (e.g., 0-500 .mu.M H.sub.2O.sub.2 in 50 .mu.M
increments, a.fwdarw.k; E.sub.APP=-0.15 V vs. Ag/AgCl). An
exemplary calibration curve is shown in the inset of FIG. 28B. For
example, as shown in FIG. 28, well-defined currents, proportional
to the H.sub.2O.sub.2 concentration, were observed. The exemplary
resulting calibration curve, based on sampling the current at 15 s
following the potential step, displays high linearity
(R.sup.2=0.999; as shown in the inset). The response for 50 .mu.M
H.sub.2O.sub.2 (curve b) indicated a limit of detection (LOD) of
.about.20 .mu.M (S/N=3). The ability to detect H.sub.2O.sub.2 at
low potentials is an attractive feature of the disclosed Rh-CPE
microneedle array when positioned for use in minimally-invasive
oxidase-based biosensors.
[0159] Exemplary implementations of the disclosed embodiment
included evaluations on the effect of reconstitution of the carbon
paste matrix within the microelectrode array. For example, a key
advantage of carbon paste-based electrodes is their renewable
surface, which can be readily regenerated. Such regeneration should
facilitate the re-use of the microneedle array. Accordingly, the
effect of repetitive packing of the array upon the resulting
response was evaluated. As such, five calibration experiments were
executed for H.sub.2O.sub.2 over the 50 to 500 .mu.M H.sub.2O.sub.2
range, which involved successively reconstituted carbon paste
surfaces. Between each experimental implementation, the electrode
was thoroughly disassembled, cleaned, reassembled, and repacked;
its electrochemical response was then characterized. FIG. 29 shows
a plot of a calibration curve obtained for H.sub.2O.sub.2
concentrations from 0 to 500 .mu.M in 50 .mu.M increments (e.g.,
E.sub.APP=-0.15 V vs. Ag/AgCl, t=15 s). The effect of
reconstitution of the Rh-dispersed carbon paste microneedle array
is illustrated for five subsequent reconstitution operations. The
results, illustrated in FIG. 29, are indicative of a
highly-repeatable calibration. For example, the response of
successive packings deviated by no more than 5.4% from the average
current at each level over the examined concentration range. Highly
linear results were observed over the concentration range
(R.sup.2=0.997), along with a very low standard deviation (e.g.,
.sigma.<10 nA). These exemplary data demonstrate that repeated
packing/unpacking of the carbon paste constituent in the
microneedle array resulted in a reproducible electrochemical
response.
[0160] Exemplary implementations of the disclosed embodiment
included the biosensing of lactate at the microneedle CPE arrays.
For example, an exemplary microneedle array CPE biosensor for
lactate was developed. Lactate oxidase (LOx)-dispersed metallized
carbon paste was prepared using PEI for the electrostatic
entrapment of the enzyme within the matrix. Chronoamperometric
calibration experiments were performed using the LOx-Rh-carbon
paste microneedle array at -0.15V vs. Ag/AgCl for increasing levels
of lactate (e.g., 0 to 8 mM in 1 mM increments). Typical
chronoamperograms are displayed in FIG. 30A, which shows exemplary
plots of chronoamperograms obtained for lactate concentrations from
0 to 8 mM in 1 mM increments (e.g., E.sub.APP=-0.15 V vs. Ag/AgCl).
FIG. 30B shows an exemplary calibration curve corresponding to the
chronoamperometric current at t=15 s. For example, high linearity
(R.sup.2=0.990) and low deviation (e.g., .sigma.<10 nA) were
observed. A detection limit was estimated to be 0.42 mM lactate
(S/N=3), which is well below normal physiological levels and is
therefore more than sufficient for relevant applications. It should
be noted that the exemplary linear concentration range encompasses
the entire physiological and pathological range of lactate in
transdermal fluids, e.g., indicating the diagnostic value of the
microneedle-based lactate biosensor.
[0161] Exemplary implementations of the disclosed embodiment
included evaluating the microneedle CPE arrays for interference by
common electroactive compounds. For example, in order to ascertain
that the exemplary biosensor could function as intended in the
presence of common electroactive substances found in transdermal
fluids, an interference investigation was conducted using
physiological levels of these compounds. FIG. 31 shows plots of
chronoamperograms showing the effect of physiologically-relevant
electroactive interferents upon the detection of 1 mM lactate in
the presence of 60 .mu.M ascorbic acid (AA), 500 .mu.M uric acid
(UA), and 200 .mu.M acetaminophen (AC) (e.g., E.sub.APP=-0.15 V vs.
Ag/AgCl). As shown in the figure, the addition of any of these
common electroactive interferents resulted in a negligible effect
on the lactate response of the exemplary biosensor device. For
example, a maximum current deviation of only 1.5% from the 1 mM
lactate level was observed for the addition of AC. For example,
such interference-free lactate detection reflects the strong and
preferential electrocatalytic activity of the Rh-CPE towards
H.sub.2O.sub.2, which can be detected by the described microneedle
paste biosensor, e.g., for lactate monitoring in transdermal
fluids.
[0162] Exemplary implementations of the disclosed embodiment
included evaluating the stability of the lactate response of the
microneedle CPE arrays. For example, the stability of the
microneedle array-based biosensor was examined from repetitive
chronoamperograms for 2 mM lactate over a 2 hour period. In some
examples, an initial short preconditioning step was implemented.
This process involved the immersion of the exemplary CPE
microneedle array in a 0.1 M potassium phosphate buffer (pH 7.0)
and the concomitant recording of six chronoameprograms, followed by
the immersion of the array in a 2 mM lactate solution for 10 min
while recording two chronoamperograms. After the exemplary
preconditioning, the current was sampled every 10 min over the
entire 2 hour stability test period. FIG. 32 illustrates the
time-course profile of the resulting current response, e.g., with
the initial reading at t=0 min normalized to 100%. FIG. 32 shows a
data plot of the stability of the electrochemical response of the
microneedle array for 2 mM lactate (E.sub.APP=-0.15V vs. Ag/AgCl)
over a 2 hour duration. As shown in the figure, a stable current
was achieved almost immediately following the initialization of the
experiment, with only a slight increase (of 9.7%) over the entire 2
hour time course. The stable response reflects the integrity of the
exemplary CPE microneedle array biosensor. For example, tight
packing of the CPE, which can prevent the potential accumulation of
the enzymatic product within the microneedle openings, can
influence the stable response.
[0163] For example, in the disclosed embodiment, the coupling of
CPE transducers with microneedle hosts was shown to provide
low-potential detection of H.sub.2O.sub.2. For example, exemplary
implementations of the disclosed embodiment demonstrated that a
reproducible amperometric response can be achieved following
successive reconstitution of the carbon paste matrix. For example,
exemplary implementations of the disclosed embodiment demonstrated
that highly-linear lactate detection can be achieved over the
entire physiological range, along with the high selectivity
imparted by the very low cathodic detection potential. The high
selectivity, sensitivity, and stability of the described CPE
microneedle array demonstrates the ability of the array to be
implemented in diverse on-body sensing applications.
[0164] While this patent document contains many specifics, these
should not be construed as limitations on the scope of any
invention or of what may be claimed, but rather as descriptions of
features that may be specific to particular embodiments of
particular inventions. Certain features that are described in this
patent document in the context of separate embodiments can also be
implemented in combination in a single embodiment. Conversely,
various features that are described in the context of a single
embodiment can also be implemented in multiple embodiments
separately or in any suitable subcombination. Moreover, although
features may be described above as acting in certain combinations
and even initially claimed as such, one or more features from a
claimed combination can in some cases be excised from the
combination, and the claimed combination may be directed to a
subcombination or variation of a subcombination.
[0165] Similarly, while operations are depicted in the drawings in
a particular order, this should not be understood as requiring that
such operations be performed in the particular order shown or in
sequential order, or that all illustrated operations be performed,
to achieve desirable results. Moreover, the separation of various
system components in the embodiments described in this patent
document should not be understood as requiring such separation in
all embodiments.
[0166] Only a few implementations and examples are described and
other implementations, enhancements and variations can be made
based on what is described and illustrated in this patent
document.
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