U.S. patent application number 17/405943 was filed with the patent office on 2022-03-10 for metal-enzyme sandwich layers.
The applicant listed for this patent is CALIFORNIA INSTITUTE OF TECHNOLOGY. Invention is credited to Dvin ADALIAN, Samson CHEN, Muhammad M. JILANI, Xiomara Linnette MADERO, Axel SCHERER, Richard SMITH.
Application Number | 20220071529 17/405943 |
Document ID | / |
Family ID | |
Filed Date | 2022-03-10 |
United States Patent
Application |
20220071529 |
Kind Code |
A1 |
ADALIAN; Dvin ; et
al. |
March 10, 2022 |
METAL-ENZYME SANDWICH LAYERS
Abstract
Measurement of target analytes is carried out with an
enzyme-based sensor. The enzyme hydrogel is protected by a porous
layer of a metallic material. The size of the pores is small enough
to prevent degradation of the enzyme layer caused by the immune
system of an organism, but large enough to allow transfer of
molecules that participate in the electrochemical reaction allowing
the enzyme to detect the target analytes.
Inventors: |
ADALIAN; Dvin; (PASADENA,
CA) ; CHEN; Samson; (PASADENA, CA) ; JILANI;
Muhammad M.; (PASADENA, CA) ; SCHERER; Axel;
(PASADENA, CA) ; MADERO; Xiomara Linnette;
(PASADENA, CA) ; SMITH; Richard; (PASADENA,
CA) |
|
Applicant: |
Name |
City |
State |
Country |
Type |
CALIFORNIA INSTITUTE OF TECHNOLOGY |
Pasadena |
CA |
US |
|
|
Appl. No.: |
17/405943 |
Filed: |
August 18, 2021 |
Related U.S. Patent Documents
|
|
|
|
|
|
Application
Number |
Filing Date |
Patent Number |
|
|
16191357 |
Nov 14, 2018 |
11109784 |
|
|
17405943 |
|
|
|
|
62654701 |
Apr 9, 2018 |
|
|
|
International
Class: |
A61B 5/1486 20060101
A61B005/1486; A61B 5/145 20060101 A61B005/145; A61B 5/1473 20060101
A61B005/1473 |
Goverment Interests
STATEMENT OF INTEREST
[0002] This invention was made with government support under Grant
No. HR0011-15-2-0050 awarded by DARPA. The government has certain
rights in the invention.
Claims
1. (canceled)
2. A sensing device comprising: a substrate; and a working
electrode on the substrate; the working electrode comprising: a
metallic layer contacting the substrate; an enzyme layer contacting
the metallic layer, the enzyme layer having larger lateral
dimensions than the metallic layer; and a metallic porous layer
contacting the enzyme layer, the metallic porous layer being
separated from the metallic layer by the enzyme layer.
3. The sensing device of claim 2, wherein the enzyme layer
comprises an imperfect enzyme sensing layer and a porous electrical
insulator over the imperfect enzyme sensing layer.
4. The sensing device of claim 3, wherein the porous electrical
insulator comprises polyurethane.
5. The sensing device of claim 3, wherein the porous electrical
insulator comprises a sulfonated tetrafluoroethylene-based
fluoropolymer-copolymer.
6. The sensing device of claim 2, wherein the metallic layer
comprises an array of nanopillars.
7. The sensing device of claim 2, wherein the metallic porous layer
comprises an array of pores having a diameter between 2 nm and 2
micrometers.
8. The sensing device of claim 7, wherein the metallic porous layer
comprises an array of pores having a diameter smaller than 200
nm.
9. The sensing device of claim 8, wherein the metallic porous layer
comprises an array of pores having a diameter smaller than 20
nm.
10. The sensing device of claim 7, wherein the array of pores has a
total area between 0.001% and 50% of a total area of the metallic
porous layer.
11. The sensing device of claim 7, wherein the array of pores has a
total area between 0.1% and 10% of a total area of the metallic
porous layer.
12. The sensing device of claim 10, wherein the metallic porous
layer is made of a material selected from the group consisting of:
Pt, Au, W, Ti, Pd, Ir, Si, TiO2, WO3, SnO2, graphene, and InO2.
13. The sensing device of claim 2, wherein the metallic porous
layer comprises an array of pores having a diameter larger than
molecules taking part in the electrochemical reaction.
14. The sensing device of claim 2, wherein the enzyme layer
contacts the substrate and the metallic porous layer contacts the
substrate.
15. The sensing device of claim 2, wherein the metallic layer has a
thickness between 1 nm and 1 mm.
16. The sensing device of claim 15, wherein the metallic porous
layer has a thickness between 1 nm and 10 micrometers.
17. The sensing device of claim 16, wherein the enzyme layer has a
thickness less than 10 micrometers.
18. The sensing device of claim 17, wherein the enzyme layer has a
thickness less than 1 micrometer.
19. The sensing device of claim 16, wherein the enzyme layer has a
thickness less than 400 nm.
20. The sensing device of claim 2, wherein the enzyme layer
comprises an oxidase enzyme and a hydrogel.
21. The sensing device of claim 2, further comprising: a counter
electrode on the substrate; and a reference electrode on the
substrate, wherein: the counter electrode comprises: a metallic
layer contacting the substrate; an enzyme layer on the metallic
layer; and a metallic porous layer on the enzyme layer, and the
reference electrode comprises: a metallic layer contacting the
substrate; an enzyme layer on the metallic layer; and a metallic
porous layer on the enzyme layer.
22. The sensing device of claim 2, wherein the enzyme layer
comprises glutaraldehyde.
23. An array of sensing devices, each sensing device of the array
as described in claim 2, wherein a spacing between the sensing
devices of the array of sensing devices is configured to avoid
crosstalk due to leakage of intermediate reporting molecules.
24. The array of claim 23, wherein the spacing is between 10 and
250 nm.
25. The sensing device of claim 2, wherein the sensing device is
configured to be implanted internally to an organism, or attached
externally to skin of the organism.
Description
CROSS REFERENCE TO RELATED APPLICATIONS
[0001] The present application is a continuation of U.S. patent
application Ser. No. 16/191,357, filed on Nov. 14, 2018, which
claims priority to U.S. Provisional Patent Application No.
62/654,701, filed on Apr. 9, 2018, the disclosure of each of which
is incorporated herein by reference in its entirety.
TECHNICAL FIELD
[0003] The present disclosure relates to biosensing. More
particularly, it relates to metal-enzyme sandwich layers.
BRIEF DESCRIPTION OF DRAWINGS
[0004] The accompanying drawings, which are incorporated into and
constitute a part of this specification, illustrate one or more
embodiments of the present disclosure and, together with the
description of example embodiments, serve to explain the principles
and implementations of the disclosure.
[0005] FIGS. 1-2 illustrate exemplary steps in the fabrication
process of a sensor.
[0006] FIG. 3 illustrates an exemplary electrode configuration.
[0007] FIG. 4 illustrates exemplary images of irregular peaks.
[0008] FIG. 5 illustrates exemplary periodic nanopillars.
[0009] FIG. 6 illustrates an exemplary sensor with a protective
metallic top layer.
[0010] FIG. 7 illustrates exemplary data.
[0011] FIG. 8 illustrates a schematic approach of building an
electrochemical sensor on a flat substrate.
[0012] FIGS. 9-10 illustrate different embodiments of sensing
devices.
[0013] FIG. 11 illustrates an exemplary array of sensing
devices.
SUMMARY
[0014] In a first aspect of the disclosure, a sensing device is
described, the sensing device comprising: a working electrode on a
substrate; wherein: the working electrode comprises: a metallic
layer contacting the substrate; an enzyme layer on the metallic
layer; and a metallic porous layer on the enzyme layer, and the
working electrode, the metallic porous layer, and the enzyme layer
are configured to detect a target analyte through an
electrochemical reaction.
DETAILED DESCRIPTION
[0015] Layers containing enzyme molecules are important for the
electrochemical measurement of glucose, lactate, and many other
metabolites of medical importance in the body. As known to the
person of ordinary skill in the art, an enzymatic reaction can be
used to selectively convert analytes of interest into by-products,
releasing electroactive compounds during chemical reactions
involving the analyte molecules. Electrochemical measurement of the
rate of electroactive compound production can then provide
concentration data of the analyte molecules. These measurements are
not generally error-free, since the enzyme reaction may not be
fully selective to the material of interest. In other words, the
enzyme reaction may occur with false positives. Another factor is
that the reagents necessary for the reaction to occur should be
present in sufficient quantity, otherwise the electrochemical
measurement will be limited by the reagents' quantity and not be
entirely accurate. Further, the conversion rate from the enzyme
should be high enough to generate electroactive compounds
proportional to the analyte molecules, otherwise the signal can be
lost in the noise.
[0016] In some embodiments, the electroactive compounds may be
ions. However, in some embodiments, the electroactive compounds may
be molecules other than ions. For example, the product of interest
in the typical glucose oxidase reaction (hydrogen peroxide) is not
an ion. In the present disclosure, some embodiments may refer to
ions as the electroactive compounds, however these embodiments may
also be implemented more generally without ions but with
non-charged electroactive compounds. In the present disclosure
"electroactive compounds" refer to ions and non-charged compounds
that are capable of generating currents through an electrode
through chemical reactions on the surface of that electrode.
[0017] If a significant current from the electroactive product can
be collected, accurate measurement of metabolic biomarkers is
possible, as long as the enzyme molecules remain active. Therefore,
enzyme degradation has to be taken into account, and prevented or
at least reduced to prolong longevity of the sensor. The efficiency
of conversion for the analyte reaction into measured current is
characterized by the conversion efficiency, which can be optimized
if electroactive products are effectively collected onto the
measurement electrode and are not lost to the surrounding
environment through diffusion. The reaction rate should also be
limited by the presence of the analyte of interest, and the
concentration gradient of reagents to the enzyme molecules
maintained by an adequate turnover rate.
[0018] Enzyme reactions have been used since the 1960s to measure
glucose and other analytes. The first electrochemical glucose
detector measured the current on a working electrode, biased with a
negative 0.6 V with respect to a counter electrode, in order to
collect charge from hydrogen peroxide that is generated by an
enzymatic reaction between glucose and oxygen, catalyzed by glucose
oxidase (GOx). In this catalytic reaction, glucose is oxidized to
hydrogen peroxide and D-glucono-lactone. The electronic current
from the hydrogen peroxide conversion to water and oxygen at the
metallic working electrode was measured with a potentiostat. The
current is proportional to the glucose concentration over a wide
range of concentrations. Since the early 1960s, many other enzymes,
both oxidases and dehydrogenases, have been identified to measure
other analytes of interest. Such enzymes can selectively react with
glucose, lactate, alcohol, urea, cholesterol, xanthine, and several
dozens of other analytes of interest.
[0019] In electrochemical enzyme-based sensors, the conversion
efficiency of analyte molecules into measurable current generally
changes over time, as enzyme molecules are attacked, dissolved or
lose their co-factors. These processes lead to the deterioration of
the sensitivity and accuracy of the sensor over time. There exist
several failure modes for electrochemical sensors that rely on
enzymatic reactions. For example, enzymes can deteriorate by
becoming oxidized due to the molecular compounds in their
surrounding environment, or due to the reactive compounds that the
sensors themselves generate. Enzymes can also lose the co-factors
which enable the efficient reaction with analytes. Other necessary
reagents for the enzymatic reaction, such as oxygen for oxidase
enzymes, can be depleted, or the reactive enzyme film can simply
delaminate by degradation of its adhesion to the substrate. When
sensors are located within living organisms, these failure modes
are generally accelerated, as the organism's immune system will
attack the sensor through its foreign body response, and contribute
to enzyme deterioration by oxidation. In vivo implants can become
encapsulated, a process through which the immune response builds
scar tissue diffusion barriers around the implanted device,
preventing reagents and analytes from penetrating to the sensor
electrodes and isolating the implant from the rest of the organism.
All of these effects ultimately limit the sensor lifetime, or the
length of time during which the electrochemical sensor can provide
useful and accurate information. Most sensors are limited to
several weeks in the challenging environments encountered in living
organisms.
[0020] The present disclosure describes new methods for avoiding
the common degradation mechanisms and prolonging the lifetime of
enzyme-based electrochemical sensors, and other sensors that rely
on monitoring selective (binding) reactions on or close to
electrochemically-active contacts. Typically, electrodes are used
to form electrochemical sensors by either using needle or wire
geometries or by defining electrodes on flat surfaces. Ideally,
electrode materials should be inert to the surrounding and not
suffer from rapid corrosion. Platinum, gold, tungsten, iridium and
titanium are therefore preferred materials since these are
bio-compatible materials that do not deteriorate rapidly in the
body, and do not generate deleterious immune system responses.
[0021] The impedance of these electrodes ultimately determines the
current that can flow through the electrodes at a given voltage, as
it is desirable for electrodes to minimize that circuit resistance
in order to maximize the sensor sensitivity and the signal to noise
ratio. This impedance can be modified by changing the surface area
of the electrode. An increased surface area can be accomplished by
increasing the lateral dimensions of the sensor, or by fabricating
a three-dimensional electrode surface with an increased surface
area for the same lateral dimensions. For example, the surface area
can be increased through corrugation, by growing sponge-like
geometries, or by fabricating regular or irregular nanopillars or
jagged structures. Typically, large surface-to-volume ratio
electrodes can provide low impedance contacts to electrochemical
reactions.
[0022] In traditional electrochemical sensors that rely on
monitoring the ion generation of enzymatic reactions, the metal
contact surface is coated with a thin adhesive layer that contains
the enzyme. This layer is often called a "hydrogel", as it consists
of a relatively porous material that enables the smaller molecule
analytes (glucose, lactate, etc.) to diffuse through the layer to
react with enzyme molecules that are immobilized within the
molecular structure of this layer. A traditional chemistry used in
glucose monitors is the glutaraldehyde system, a polymer that
cross-links to form a chemically stable scaffold into which enzyme
molecules can be immobilized. The polymer is porous enough to
enable the enzyme molecules to remain active, and enable reagents
to enter and products to leave the polymer. In commercial sensors,
the glutaraldehyde-based layer containing the enzyme is typically
several micrometers thick. The enzyme molecules generate ions that
need to travel through this thickness of several micrometers, in
order to be read at the electrode's surface. Unfortunately, any
residual glutaraldehyde is immunogenic and can accelerate the
natural immune response. Therefore, the glutaraldehyde-based layer
is generally covered with a protection layer consisting of
polyurethane or PEGylated with polyethylene glycol to reduce the
immune response. These additional layers often reduce the
sensitivity of the enzyme sensor, and are slowly eroded as they are
attacked by natural oxidizing species, in turn leading to
relatively short sensor lifetimes.
[0023] The present disclosure describes a method for prolonging the
lifetime of enzyme-based electrochemical sensors. The approach uses
a thin metal layer deposited on the top surface of a thin layer
containing the active enzyme. For example, platinum can be
deposited in a very porous geometry often referred to as "platinum
black". Other materials such as titanium, gold or tungsten can also
be used for fabricating a metallic porous layer that allows target
analytes, reagents and by-products to move through the porous
electrode in both directions, thus allowing access to the enzyme
hydrogel underneath. The platinum (or other metal) layers can be
deposited through a vacuum deposition process in which platinum is
sputtered from a platinum target and deposited onto the sample. The
sputter deposition process enables the control of both the
thickness and the microstructure of the platinum layer by adjusting
the deposition parameters, which include the voltage, power, and
vacuum pressure during the deposition. Therefore, a platinum black
layer can be used as a low-impedance contact surface for
electrochemical electrodes. Sputter deposition is not the only way
to deposit metals, and alternative methods include evaporation from
a heated source. The source can be heated through electrical
resistance methods, or with an electron beam. In other embodiments,
other methods of deposition can be used, such as pulsed laser
deposition. The key requirement of the deposition technique is that
the local heating of the enzyme layer and the time required for the
deposition should be minimized. Local heating could, in fact,
degrade the enzyme layer. Although corrosive solutions may be
needed for deposition of materials, it is also possible to deposit
metal layers by electroplating or electroless plating from plating
baths.
[0024] Instead of depositing a single enzyme-containing active
layer on top of the platinum electrode layer, the present
disclosure describes the deposition of a platinum/enzyme/platinum
sandwich or multilayer. The structure comprises a top platinum
layer which is relatively porous to enable analytes to reach the
enzyme molecules in the middle layer, and a bottom platinum layer
which acts as an electrode. This approach brings with it several
important advantages. The platinum layer protects the enzyme layer
underneath from oxidizing species that would attack the polymer
through the immune response process. The pores in the top metallic
layer are nanoscale, therefore the porous platinum layer prevents
the immune system from probing and recognizing the underlying
chemistry, removing the immunogenic behavior. In fact, the active
components of the immune system are generally larger than a
nanoscale pore. The top platinum layer optionally acts as another
contact surface that captures escaping ions that would otherwise be
lost, and enables their contribution to the electrical measurement.
This is possible if the top platinum layer is electrically
connected to the bottom platinum layer, either through vias through
the enzyme layer or with electrical pathways around the enzyme
layer. The platinum layer also enables the retention of oxygen,
which is generated when peroxide ions react at the metal surface
converting into water and dissolved oxygen. The oxygen is needed
for subsequent reactions that require oxygen as a reagent, for the
enzyme assisted conversion to successfully take place. The distance
between the bottom platinum layer and the top platinum layer can be
reduced to control the loss of co-factor through diffusion to the
solution surrounding the sensor. All of these effects have been
shown to improve the linearity of the glucose response of the
sensor.
[0025] In some situations, it may be preferable to electrically
attach the top porous metal layer to the bottom electrode. This
process can effectively double the available electrode area, which
may be desirable in cases where sensitivity is important. The
attachment can be carried out in several ways. As shown in the
cross-sections of FIG. 9, the enzyme-based sensing layer (905) may
be fabricated to be slightly smaller than the bottom electrode
layer (910). In other words, the lateral dimensions of the enzyme
layer are slightly smaller, leaving some space at the sides. The
top conductive porous layer (900) may then be fabricated to be
larger than the enzyme-based sensing layer. In this way, the top
layer is electrically connected to the bottom electrode layer with
conductive pathways around the periphery of the enzyme layer.
Alternatively, the enzyme-based sensing layer (920) may be
fabricated with multiple holes or vias (902), through which the top
porous layer (915) is electrically connected to the bottom
electrode layer (925). These holes may be formed by lithographic
patterning techniques, and in some cases these holes may be formed
by adjusting the deposition parameters of the enzyme-based sensing
layer so that it natively contains holes. One advantage of this
fabrication technique is that the large number of conductive
attachment points improves both the reliability of the electrical
connection and its mechanical reliability. Another way of forming
these conductive attachments is to fabricate the bottom electrode
layer (945) so that it has pillars (904). The enzyme-based sensing
layer (935) may then be deposited onto the bottom electrode layer
(945), leaving the top surface of the pillars exposed. When the
porous layer (930) is deposited onto the structure, it contacts the
top surfaces of the pillars, forming a multitude of conductive
pathways between the top and bottom electrodes.
[0026] These conductive attachments can themselves increase the
effective electrode area, as each conductive attachment adds some
surface area. Thus, in scenarios where an increased electrode area
is beneficial, the width of each conductive attachment may be as
small as 5-40 nanometers, which is the limit of most practical
fabrication techniques for this type of structure. The spacing
between attachments may be as small as 100 nanometers. However,
such attachments may be fragile and are generally complex to
produce. If such high effective surface area is unnecessary or
potentially not beneficial due to higher background currents,
larger attachments, such as 1 to 25 micrometers, and a larger
spacing, such as 1 to 25 micrometers, are preferable. The shapes of
these attachments are typically dictated by the fabrication
technique, and could have, for example, circular or rectangular
cross sections. The height of these attachments is generally
selected based on the desired thickness of the enzyme sensing
layer.
[0027] In some embodiments, it may be preferable to ensure that the
top conductive porous layer is electrically disconnected from the
bottom electrode layer. In fact, the top conductive porous layer
itself can generate interfering currents from interfering compounds
generated outside of the sensor. Additionally, the increased
surface area can increase background currents. These interfering
and background currents may be large enough in some scenarios to
negate the benefit of increased surface area for collecting
molecules from the enzyme sensing layer. There are multiple ways of
maintaining electrical disconnection between the top porous layer
and the bottom electrode layer. In FIG. 10, one way of maintaining
the electrical separation is shown. The enzyme sensing layer (1005)
is fabricated to have larger lateral dimensions than the bottom
electrode layer (1000). The top porous layer (1010) is then
fabricated to have larger lateral dimensions compared to the enzyme
sensing layer, fully encapsulating the enzyme layer, while avoiding
any electrical connection between the top and bottom layers.
[0028] In some embodiments, the enzyme-based sensing layer natively
contains holes when deposited, causing the top and bottom layers to
be electrically connected. In these cases, it may be necessary to
deposit an additional insulator layer. For example, if an imperfect
enzyme sensing layer (1020), containing holes, is deposited on the
bottom electrode (1025), a thin, porous electrical insulator (1015)
may be deposited on top of the enzyme sensing layer (1020), to
separate the bottom electrode (1025) from the top, porous
conductive electrode (1030). A variety of porous insulating
materials may be used, including for example polyurethane and
Nafion.TM., which are commonly used for this purpose in
electrochemical sensors. In some cases, an inorganic insulator,
such as silicon nitride or silicon dioxide, may be fabricated with
pores to accomplish the same task. In general, the material for
layer (1015) must be an electrical insulator and contain
sufficiently large holes to readily allow the analyte of interest
to reach the enzyme sensing layer. The thickness of this layer is
determined by a number of factors, including the thickness required
to maintain sufficient electrical insulation, porosity of the
layer, nature of the deposition method of the top porous layer, and
roughness and thickness of the underlying enzyme sensing layer. In
general, this layer is made as thin as possible so that it does not
impede the transport of analytes or other molecules to and from the
sensing layer. For example, inorganic insulators which typically
have superior insulation performance. If using an inorganic
insulator, a layer between 10 nm to 500 nm may be sufficient to
perform the function. If an organic insulator is used, a layer
thickness of 100 nm to 5 micrometers may be used.
[0029] As known to the person of ordinary skill in the art,
alternative methods for obtaining linear sensor performance require
the use of organic control layers to control diffusion, for example
polyurethane. The precise thickness of these layers, as well as
their adhesion chemistry, are very dependent on the ambient
humidity and temperature at which they are deposited. The
variability, being hard to control, can introduce uncertainty in
the measurements. The porous metal diffusion control layers
described in the present disclosure offer a solution with improved
ion collection efficiency and oxygen recycling, as none of the
peroxide molecules are lost to the surrounding tissue, resulting in
almost no reduction in measured sensor sensitivity. Therefore,
using a porous metallic layer decreases the variability between
different sensors, in turn increasing the accuracy and reliability
of each sensor.
[0030] In addition to providing more linearity and oxygen
insensitivity, the metal surface protective layers also enable the
protection of enzyme molecules from attack, by hiding the sensing
chemistry from the immune system, thereby yielding efficient
electrochemical sensors with very thin enzyme layers that offer
long lifetimes. The thickness of the enzyme layer determines the
delay between the reaction of the sensor to the analyte molecules
of interest, and the observation of the electrical signal from the
released ions at the contact surfaces. This delay can be reduced,
in the structures of the present disclosure, to be limited by the
enzyme reaction speed determined by turn-over, rather than be
limited by the diffusion rate of the ions within the polymer
matrix.
[0031] Metal deposition through vacuum processing brings with it
many potential pitfalls when thin and delicate organic materials
are to be coated. However, by appropriate polymerization of the
enzyme layer with glutaraldehyde and subsequent metallization using
a platinum sputter deposition process, it is possible to deposit
metallic layers on enzymes without much deterioration in their
chemical performance. The metal/enzyme/metal sandwich layer can
even be deposited onto a photoresist lift-off mask, and patterned
by using a lithographic approach. This fabrication approach leaves
the active sandwich layer only on top of the electrodes of
interest, and thereby enables the functionalization of select areas
with micrometer accuracy. The lift-off chemistry, which consists of
using acetone to dissolve the photoresist stencil layer
selectively, to define the pattern of metal/enzyme/metal
functionalized areas, does not significantly reduce the enzyme
activity of glucose oxidase, and permits sandwich structures to be
defined on top of electrochemical potentiostat detectors. This may
also hold true for other enzymes, such as lactate oxidase or
urecase, and leads to the opportunity of defining several different
chemical sensors on the same substrate by a series of lithographic
processes depositing different enzyme layers onto electrochemical
working electrodes. Therefore, in some embodiments, a sensor may
comprise a plurality of areas, each area having a different enzyme
layer, in order to detect a plurality of target analytes with the
same sensor. Unlike sensor systems known in the art, in which
peroxide or ammonia molecules can escape, metal-based multilayers
as described in the present disclosure prevent escape of these
molecules, avoid cross-talk and enable precise measurements of
several analytes in close proximity.
[0032] In such traditional enzyme-based sensor systems, measuring
more than one analyte within a compact sensor is typically very
difficult, due to crosstalk. Specifically, many enzymes for
different analytes generate the same reporter molecule (the
reporter molecule being the molecule that generates sensing), and
reporter molecules generated due to one analyte may be sensed by
the electrode for another analyte. For example, an implantable
sensor designed to measure both lactate and glucose may have two
electrodes, one coated with lactate oxidase and one coated with
glucose oxidase. When glucose levels are high, hydrogen peroxide
generated by glucose oxidase may diffuse to the lactate oxidase
electrode, falsely elevating the lactate reading. These electrodes
may need to be separated, normally, by a distance of 1 mm or more
to reduce crosstalk to acceptable levels, depending on the
environment in which the sensor is implanted, and depending on the
required accuracy. Wireless implantable sensors can be smaller than
1 mm, with electrodes as small as 200.times.200 microns, therefore
this required separation can limit minimum device size when more
than one analyte must be sensed. Because the metal/enzyme/metal
sensor described in the present disclosure has a porous top layer
designed to react with or capture any escaping hydrogen peroxide,
or other reporter molecules, much less cross-talk occurs, and
multi-analyte sensors can have separations of as little as 10-250
micrometers between working electrodes, depending on the required
porosity of the top metal layer and depending on the required
accuracy of each sensor. This advantage permits the fabrication of
wireless implantable sensors with areas smaller than 1 mm.sup.2'
which can detect multiple analytes in one device.
[0033] In some embodiments, the sensor can also comprise wireless
electronics to transmit its measurements from within a human body
to external equipment. The sensor, for example, may have dimensions
of 1.2 mm by 1.2 mm, a fast response time of 0.1 s, use a low power
of 5 microWatts, and cost only 10 cents. Sensors with millimeter
scales dimensions do not move with respect to surround cells,
maintaining a constant impedance. The sensors can also be implanted
in many locations due to their small size, they can be injected,
and produce small tissue irritation. Possible applications comprise
monitoring of chronic diseases such as diabetes, kidney failure,
cardiovascular problems and cancer. The sensors can also be used in
skin patches to monitor cortisol, alcohol, glucose, and
applications such as sport medicine and post-operative monitoring.
Power can be transmitted to the sensor through wireless
transmission, for example using coil antennas. A typical
configuration for the electrode may be used, for example with
concentric, square, working, reference and counter electrodes. For
example, FIG. 3 illustrates a configuration with a counter
electrode (305), a reference electrode (310), and a working
electrode (315).
[0034] In some embodiments, platinum is used to fabricate all three
electrodes. Platinum allows the use of lower voltages and a more
stable operation compared to known Ag/AgCl electrodes. FIG. 4
illustrates atomic force microscope images of an exemplary
electrode structure with higher irregular peaks for use in high
current sensors, e.g. 500 nA: (405,410). FIG. 4 also illustrates
lower peaks for use in low current sensors, e.g. 5 nA: (415,420).
FIG. 5 illustrates exemplary nanopillars that can be fabricated on
the surface of an electrode to increase its surface area and
sensitivity. The nanopillars are shown in a close-up in the top
picture of FIG. 5 and in a top perspective view in the bottom
picture of FIG. 5. The nanopillars can be of the same metallic
material to form a uniform electrode. In some embodiments, the
sensors are fabricated by spin coating the enzyme, instead of drop
coating. Spin coating allows thinner electrode layers. As hydrogen
peroxide must diffuse to the metal electrode surface to be
measured, thinner enzyme layers increase the collection efficiency
of the sensor. In some embodiments, the enzyme hydrogel layers are
no thicker than 1 micrometer, and generally between 2 nm and 1
micrometer. Enzyme molecules are typically larger than 1-2
nanometers, therefore the minimum enzyme layer is typically the
size of the enzyme molecule. The top electrode controls the rate of
transfer for the target analyte to diffuse through the pores in the
metal, towards the enzyme layer, and the rate of transfer of the
peroxide. The top metallic layer can have pores (or pinholes)
between 1 nm and 1 micrometer. The metallic layer forms an
electrochemical capacitor, therefore its porous structure affects
the electrical properties of the sensor. The bottom metallic layer
can be as thick as necessary since it provides structural support
for the other layers.
[0035] In some embodiments, platinum provides the best choice for
the porous layer, as it matches the working electrode and reference
electrode materials, it is inert, and lasts for a long time. In
some embodiments, gold can be used as it not very reactive, and can
be generally considered biocompatible. Gold can complicate the
measurements within the electronic cell as there may be galvanic
voltages and currents between gold and the reference electrode.
Tungsten can also be used, as it is inert and generally
biocompatible, but can also complicate measurements with galvanic
interactions. Titanium is a biocompatible material, therefore it
can also be used, though it tends to oxidize, but the oxide is
conductive so oxidization is not generally a problem, though
electrochemical interaction can be problematic. Palladium is a
material used in many implants and can be considered reasonably
biocompatible. Iridium is biocompatible but very expensive. Silicon
has some conductivity and it is reasonably biocompatible, however
it can generate galvanic interactions. In some embodiments, the
above metals or other metals can be used, as well as their alloys.
Semiconductors may also be used, as well as conductive metallic
oxides such as, for example, TiO.sub.2, WO.sub.3, SnO.sub.2,
InO.sub.2, etc. Deposition techniques used may be vacuum
techniques, such as sputter deposition and vapor deposition, as
well as electroplating or electro-less plating.
[0036] The pores in the metallic, protective structures are large
enough for glucose, oxygen, and water to go through, but small
enough to avoid cells and large molecules from attacking the
enzyme. Therefore, the pore size or diameter can, in some
embodiments, range from 2 nanometers to 2 micrometers. In some
embodiments, the pore size is below 200 nanometers, or below 20 nm.
Controlling the porosity of the layer enables controlling the
amount of oxygen that can escape from the enzyme layer, and the
amount of glucose that can enter the enzyme layer. It is possible
to define a porosity ratio for the sensor by considering the area
of the pores as the open area of the sensor which allows passage of
chemical species, and comparing it to the closed area of metal
around the pores. The porosity ratio can be, in some embodiments,
between 0.001% and 50%. In some embodiments, the porosity ratio is
between 0.1% and 10%, and can be controlled by controlling the
deposition parameters. In other words, the open area of the pores
is between 0.001% and 50% of the total area of the porous layer, or
between 0.1% and 10% of the total area of the porous layer.
[0037] In some embodiments, the layer thicknesses of the three
individual layers of the sandwich structure are as in the
following. The bottom (contact) layer can be between 1 nm and 1 mm
thick, as it needs to provide a conducting substrate. Needles could
be used, or metallic layers could be deposited onto silicon chips
or polymer supports. The top (protective) layer can be between 1 nm
and 10 micrometer thick. In some embodiments, the top metallic
layer is thinner than 1 micrometer, or thinner than 100 nm. In some
embodiments, the enzyme layer can be between 10 nm and 50
micrometers. In some embodiments, the enzyme layer has a thickness
below 10 micrometers, or less than 1 micrometer, or less than 400
nm.
[0038] FIG. 1 illustrates a substrate (105) and a metallic layer
(110) for a working electrode, for example made of Pt by sputter
deposition. The metallic layer is patterned to produce a higher
surface area (115). The pattern may be irregular pillars such as in
FIG. 1, or regular, periodic pillars as in FIG. 5. As illustrated
in FIG. 2, an enzyme layer (205) is spin coated onto the patterned
metallic layer. For example, an enzyme hydrogel can be used. For
example, an enzyme and BSA coating can be used. A layer of
glutaraldehyde (210) can be deposited to crosslink the enzyme
hydrogel, for example by vacuum evaporation. For example,
glutaraldehyde in a water solution within an open container is
inserted in a vacuum chamber, leading to evaporation onto the
sample within the chamber. In some embodiments, this process is
applied on the working electrode only, while in other embodiments
it can also be applied to one or two of the other electrodes of
FIG. 3. In some embodiments, layer (210) in FIG. 2 may comprise a
bottom layer of glutaraldehyde and a porous metallic layer on top
of the glutaraldehyde. In some embodiments, layer (210) does not
comprise glutaraldehyde but only a porous metallic layer.
[0039] FIG. 6 illustrates an exemplary sensor with a protective
metallic top layer. The sensor comprises a substrate (625), a first
metallic, patterned layer (620), an enzyme hydrogel layer (615),
and a second, metallic patterned layer to protect the enzyme
hydrogel. The second layer comprises a porous layer (610), which
has openings allow passage of select chemical species while
blocking other chemical species. The second layer can also have a
patterned area on top (605). For example, the top porous Pt layer
may be 50 nm thick, the enzyme layer may be 300 nm thick, and the
Pt bottom layer may be 100 nm thick.
[0040] FIG. 7 illustrates exemplary data comparing a sensor with an
uncoated enzyme layer (710), with the superior performance of a
sensor having a protective porous layer as described in the present
disclosure (705). In some embodiments, the sensor can comprise
enzymes such as glucose oxidase, lactate oxidase, xanthine oxidase,
cholesterol oxidase, sarcosine oxidase, cortisole oxidase, urate
oxidase, alcohol oxidase, glutathione oxidase, and nicotine
oxidase. In some embodiments, the sensor can be attached to the
skin with an adhesive patch, for example to monitor sweat or
iontophoretically extracted materials. The external sensor may be
part of a fitness tracker with an external battery and
communication modules.
[0041] FIG. 8 illustrates an exemplary photolithography process: a
sensor comprises three electrodes (805); a photoresist is spin
coated on the sensor (810); after exposition to light, part of the
photoresist is lifted off (815), allowing the deposition of
multiple layers, such as the enzyme and crosslinker, and the porous
metallic protective layer; the remaining photoresist is completely
lifted off (820), and the three sensors now have an enzyme hydrogel
and porous metallic layer on top. For example, as visible in FIG. 8
(825), three layers can be deposited on the resist, comprising a
first metallic layer, an enzyme layer, and a porous metallic layer
on top. In some embodiments, where the sensing device is fabricated
by lithographic processes, it is possible to use the metal
deposited on top of the enzyme layer to protect the underlying
enzyme molecules during lithographic processing (i.e. exposure to
acetone, etc.).
[0042] The present disclosure describes a structure, consisting of
a layer of reactive material on a substrate, covered with a porous
inorganic layer capable of controlling the flow of reagents and/or
products into or out of the reactive material to control reaction
rates. In some embodiments, the inorganic porous layer is platinum,
gold, tungsten, iridium, titanium, carbon (including graphene or
carbon nanotubes) or an otherwise biocompatible material with very
low immunogenicity, and the reactive layer is enzymatic. In some
embodiments, the inorganic layer is platinum or another catalytic
metal capable of reducing reactive oxygen or oxidation species
before they can reach the sensitive reactive layer. In some
embodiments, the sensor is capable of quantifying a specific
analyte, the reactive layer is an enzyme in the oxidase family, and
the porosity of the inorganic layer is controlled such that oxygen
can freely enter the reaction layer, the analyte is restricted, and
other interfering or damaging species (including immune cells) are
rejected, improving the overall sensitivity, oxygen independence
and high analyte concentration performance of the system.
[0043] In some embodiments, the sensor is part of a power
generation cell, the reactive layer is an enzyme, and interfering
or damaging species (including immune cells) are rejected from the
compartment. In some embodiments, the sensor comprises a layer of
reactive material on a substrate, covered with a porous conductive
layer capable of controlling the flow of reactants and/or products
into or out of the reactive material and electrochemically
transducing chemical species within the reaction chamber into or
out of external circuitry, thereby improving the transduction
efficiency of that chemical species. In some embodiments, the
substrate is itself conductive, and periodic or occasional
electrically conductive attachments are made between the porous
conductive layer and the substrate, thereby greatly improving the
transduction efficiency of the chemical species of interest. In
some embodiments, the substrate and porous conductive layer
individually are composed of platinum, PEDOT:PSS, or other
biocompatible conductive material capable of electrochemical
transduction, the conductive attachments are composed of a
biocompatible conductive material, and the reactive layer is
enzymatic.
[0044] In some embodiments, the sensor is capable of quantifying a
specific analyte, and the porosity of the inorganic layer is
controlled so that the analyte and reactants can freely enter the
reaction layer and other interfering or damaging species (including
immune cells) are rejected, thereby permitting electrochemically
active products to be transduced with very low loss due to the near
complete encapsulation of the reactive layer. In some embodiments,
the sensor is part of a power generation cell, and the porosity of
the inorganic layer is controlled such that the reactants can
freely enter the reaction layer and other interfering or damaging
species (including immune cells) are rejected, thereby permitting
electrochemically active products to produce energy with very low
loss due to the near complete encapsulation of the reactive layer.
In some embodiments, the biosensor comprises a reactive layer based
on an oxidase enzyme. In some embodiments, the pores are sized to
promote the flow of oxygen into the reactive layer, reject the
entrance of interfering and damaging species, including immune
cells, into the reactive layer, reduce the loss of hydrogen
peroxide from the reactive layer, and optionally reduce the flow of
analyte into the reactive layer with respect to oxygen to improve
the linearity of the biosensor at high analyte concentrations.
[0045] In some embodiments, the porous layer and substrate are
composed of platinum, which has nearly no immunogenicity, and is
capable of catalytically reducing damaging reactive oxygen species.
In some embodiments, the nearly complete encapsulation of the
reactive layer permits nearly all hydrogen peroxide generated by
the reactive layer to be captured, and much of the oxygen generated
by the oxidation of hydrogen peroxide is reused by the reactive
layer, improving the sensitivity and linearity of the
biosensor.
[0046] In some embodiments, the porous, conductive electrode is
partially covered and filled with a reactive material, such that
the transduction of desirable reaction products is maximized, and
regeneration of reactants for the reactive material by the
conductive electrode is maximized. In some embodiments, the
conductive electrode is part of a sensor and composed of platinum,
the reactive material contains an oxidase enzyme, the desirable
reaction product is the analyte of interest, the regenerated
reactant is oxygen, thereby improving the linearity of the sensor
is improved at high analyte concentrations, and the overall
sensitivity is improved. In some embodiments, the electrode is part
of an energy harvesting system.
[0047] In some embodiments, the enzyme layer is deposited by
dipping, inkjet printing, spin coating and the top protecting metal
is deposited by a vacuum deposition technique to deliberately
contain pinholes or other microfabricated conduits for products and
reagents to travel through the metal membrane. In some embodiments
the metallic layers of the electrodes comprise irregular jagged
peaks, or an array of nanopillars. In some embodiments, the device
can be implanted internally. In some embodiments, the three
electrodes have concentric square shapes as in FIG. 3.
[0048] In some embodiments, the sensing device can comprise only
two electrodes instead of three electrodes. In these embodiments,
the two electrodes each comprise a multilayer as described in the
present disclosure, comprising a protective, porous metallic layer.
In other embodiments, the sensing device may comprise a single
electrode instead of three or two electrodes. For example, one
electrode systems use the body of the organism that is being
measured as ground, and the single working electrode to take
measurements.
[0049] In some embodiments, the top porous layer is not necessarily
metallic but may be conductive as to improve the collection of any
electroactive species, and may have catalytic activity towards any
damaging chemical species (such as reactive oxygen species
generated by the immune system). For example, this top porous
conductive layer may be composed of PEDOT:PSS or a polypyrrole. As
explained with reference to FIG. 10, in some embodiments the
metallic porous layer is electrically disconnected from the
metallic layer.
[0050] In some embodiments, the porosity of the top layer is
intrinsic to the deposition method. For example, adjusting
deposition parameters during e-beam evaporation or sputtering of
metal films can be used to control porosity. In some embodiments,
the porosity of the top layer may be more deliberately patterned.
For example, the exact positions and sizes of the pores may be
selected using e-beam lithography or photolithography. For example,
a mask containing a pattern of holes may be deposited on top of a
non-porous top layer, and subsequently used to etch controlled
holes into the top layer. In this way, the porosity of the top
layer may be better controlled than may be possible when the film
is natively deposited with pores.
[0051] In some embodiments, the bottom sensing layer is not
necessarily a metal, but a conductive electrode material. For
example, carbon, polypyrroles, and PEDOT have commonly been used as
non-metallic electrode materials. In some embodiments, the sensing
devices can be arranged in an array. A biosensor can therefore
comprise an array of sensing devices as described in the present
disclosure, wherein a spacing between the sensing devices of the
array of sensing devices is configured to avoid crosstalk due to
leakage of intermediate reporting molecules produced during the
sensing process. FIG. 11 illustrates an exemplary array of sensing
devices, where each device (1105) is as described above in the
present disclosure.
[0052] The examples set forth above are provided to those of
ordinary skill in the art as a complete disclosure and description
of how to make and use the embodiments of the disclosure, and are
not intended to limit the scope of what the inventor/inventors
regard as their disclosure.
[0053] Modifications of the above-described modes for carrying out
the methods and systems herein disclosed that are obvious to
persons of skill in the art are intended to be within the scope of
the following claims. All patents and publications mentioned in the
specification are indicative of the levels of skill of those
skilled in the art to which the disclosure pertains. All references
cited in this disclosure are incorporated by reference to the same
extent as if each reference had been incorporated by reference in
its entirety individually.
[0054] It is to be understood that the disclosure is not limited to
particular methods or systems, which can, of course, vary. It is
also to be understood that the terminology used herein is for the
purpose of describing particular embodiments only, and is not
intended to be limiting. As used in this specification and the
appended claims, the singular forms "a," "an," and "the" include
plural referents unless the content clearly dictates otherwise. The
term "plurality" includes two or more referents unless the content
clearly dictates otherwise. Unless defined otherwise, all technical
and scientific terms used herein have the same meaning as commonly
understood by one of ordinary skill in the art to which the
disclosure pertains.
* * * * *