GB2609633A - A wearable chemical sensor device and method of forming the same - Google Patents

Abstract

Electrochemical biosensor 10 detects target analyte 22, e.g. a metabolite or hormone such as cortisol, in biological fluid 30 e.g. sweat. The sensor includes a graphene layer 12 comprising catalytic particles 14 e.g. metal, metal oxide, promoting a redox reaction of the target. At least one semi-permeable membrane 16, e.g. ionomers or polymeric structures, allows selective movement of the analyte from the fluid to the graphene layer and blocks interferent species 24a e.g. ascorbic acid, preventing irritant functional catalysts from the graphene channel back to solution, maintaining biocompatibility. Flexible substrate 40 e.g. polyimide supports the graphene layer and source 11, drain 13 and gate 15 electrodes, the latter inducing current flow in the graphene channel depending on excess charges, e.g. electrons generated by redox reactions between target analyte and catalytic agent. The sensor may be thin and flexible and suitable for long-life operation with minimal performance drift.

Description

A WEARABLE CHEMICAL SENSOR DEVICE AND

METHOD OF FORMING THE SAME

The present disclosure generally relates to a method for non-invasive detection of molecular biomarkers from a skin-compatible wearable device, such biomarkers including but not limited to biomolecules such as glucose, lactate (also called metabolites), or cortisol (a hormone related to stress), relevant in health diagnostic applications, and to a wearable chemical sensor device for implementing said method.

Background of the Invention

Over the last few decades electrochemical, enzymatic-based, sensors have been used for detecting different chemical substances, or analytes, such as for instance glucose from blood (US Pat. No.5165407, Nov. 24, 1992), or from tears (US Pat. Appl. No. 2007/0043283A1, Feb 22, 2007, US Pat. Appl. No. 2013/0008803A1, Jan 10, 2013).

Other bodily fluids may be used as a source of biomarker analytes, such as urine, saliva and sweat. The latter would appear to be an ideal fluid for continuous analysis by a wearable device but tends to be underutilized as a fluid source due to a number of drawbacks associated with this type of fluid, such as very low sample volumes, contamination from the skin surface and external influences on the skin, such as skin temperature.

An example of existing work relating to this class of application has been reported by W. Gao et al. "Fully integrated wearable sensor arrays for multiplexed in situ perspiration analysis" (2016) Nature 529, 509-514, am:10.1038in attar 1 6521. In this paper a panel of target analytes (glucose, lactate, potassium and sodium) and skin temperature was selected to facilitate an understanding of an individual's physiological state, where skin temperature measurements were used to compensate for and eliminate the influence of temperature variation in the readings of the chemical sensors through a built-in signal processor.

The main limitation of this and other similar prior-art approaches relates to the use of conventional enzymatic sensing mechanisms Enzymatic sensors adopted for monitoring metabolites such as glucose or lactate are affected by a short lifetime, limited by decreasing enzymatic activity with time. This lifespan is typically one to two weeks. This relatively short lifetime increases cost and reduces the scope of enzymatic biosensors, and makes the approach unsuitable for long-term monitoring of metabolites from a wearable or long-term implanted device.

Another specific limitation of the approach outlined in the above paper is that the device is made to operate in a range of metabolite and metal ions concentrations in sweat that applies to healthy individuals, exercising physical activity in aerobic conditions. However, this approach does not necessarily cover analytically relevant values related to specific disease conditions. For instance, the sensor devices presented in the paper operate with concentration values up to 200 micromolar (0.2 millimolar) for glucose, whereas a person with diabetes has blood sugar level in the 220 millimolar range, that correspond to 0.2 to 0.6 millimolar in sweat (E. Witkowska New et al "Electrochemical Glucose Sensing: Is There Still Room for Improvement?", Anal Chem 88 (2016), 11271-11282, imps://doi.e,rg/10.' em.6b03151).

Therefore, a different mechanism to the conventional enzymatic-based sensing mechanism is desirable, that allows continuous monitoring of metabolites from sweat with a long lifetime, spanning several months/Years with no significant degradation of the sensor performance, and having the capability to compensate for skin and environmental-related parameters.

An alternative type of chemical sensor device are electrochemical sensor devices based on graphene and graphene-related active materials. Known methods in the art relating to this class of device and their use in point-of-care testing are reported in the review publication by I. Prattis, E. Hui, P. Gubeljak, G. S. Kaminski-Shierle, A. Lombardo, L.G. Occhipinti -Graphene For Biosensing Applications In Point-Of-Care Testing", li.encis In Biotech. (2021) u ci10.1016/ 2021.0.005).

An example of a device for glucose sensing in sweat outlined in this review is the publication by Xuan, X. et al -A wearable electrochemical glucose sensor based on simple and low-cost fabrication supported micro-patterned reduced graphene oxide nanocomposite electrode on flexible substrate" (2018) Biosens. Bioelectron. 109, 75-82 (intps::ic 0.10 I 6/j.bios.2( 8.02.054). Reduced Grapheme Oxide and metal nanoparticles were applied to a working electrode of an enzymatic electrochemical sensor device and covered with glucose oxidase enzyme. The developed sensor had an analytical performance with a detection range between 0.1 -2.3 inM within a 12-seconds time frame.

While this device may be suitable for a disposable, cost-effective, and responsive analytically performing enzyme-based glucose sensing patch from sweat, the detection method still requires the use of the enzymatic glucose-oxidase membrane electrodes limiting the operational lifetime of the device as with a conventional enzymatic sensor.

Furthermore, the device is affected by other limitations, such as the performance of selective measurement of targeted biomarkers by the presence of other substances in sweat that act as interferents; in fact, tests done by adding small concentrations of interferents to the reported device in targeted experiments induce variations to the readout signal that affect the accuracy of the glucose concentration value derived from the measurements. Additionally, the method and device do not allow for continuous monitoring of body temperature, and other environmental parameters, such as ambient temperature and humidity, that affect the enzyme-driven reduction-oxidation process used in correlating the concentration of the metabolite to the signal read from the working electrode of the device, resulting in: i) inability of the device to compensate for variations of the signal detected by the sensor associated to the redox reaction kinetics of the targeted biomarker, and inability of the device to perform reliably in the presence of large variation in the environmental conditions. The device is also affected by environmental sensitivity and rapid degradation of the enzyme molecules contained in enzymatic sensors, such as glucose oxidase, lactate oxidase, and others generally adopted in enzymatic-based electrochemical sensors and suffers from drift of the sensor response and performance characteristics over time due to the consumption and/or degradation of the enzyme used for the redox reaction.

Non-enzymatic electrochemical sensors have been developed where the working electrode surface is treated to include catalysts of one or more of the following categories: I) metal and metal compound catalysts, 2) metal oxides 3) composites such as alloys, bimetallic or polymer composites, 4) carbon-based micro-and nano-structured materials, including carbon nanotubes and graphene electrodes.

Generally-speaking, non-enzymatic sensors, such as those based on metal-oxides and microporous carbon electrodes do have longer lifetime than enzymatic sensors. Metal oxide sensors detect glucose via the direct oxidation reaction of glucose with an activated metal oxide contact. A review of non-enzymatic sensors and their respective mechanism of sensing is presented by K. Tian et al. "A review of recent advances in nonenzymatic glucose sensors" (2014) Mater Sol Eng C Mater App! 41, 100-118. traps://doi.org/10.1016d.msee.2014.04. 013. However, a limitation of metal oxide based non-enzymatic sensors is that they require an acidic medium to operate (pH >10), which is not the case of sweat.

An approach to develop such non-enzymatic sensors for glucose monitoring in sweat, has been reported in X. Strakosas et at "A non-enzymatic glucose sensor enabled by bioelectronic pH control" (2019) Sci. Reports 9, 10884 (hltps.:Aioi.orail0 1038/s1 1598-019-46302-9). The method is based on properties of a tri-cobalt tetra-oxide electrode contact which is converted to cobalt dioxide and acts as catalytic element for glucose oxidation in a reversible electron-producing reduction-oxidation process driven via local control of the solution pH around the electrode active area from neutral to alkaline. Unfortunately the cobalt oxide is known to produce skin irritation (MSDS=materials safety datasheet, e.g. littps://uwaterloo.cailine-ans./fsiWsica.firte-loadsiTilesicobal oxide pd1) and therefore is not suitable for application in wearable devices especially in regions of the device made to be in contact with skin, which is generally a preferred surface with which to contact a wearable device, for example a wristband.

Given the large surface area of graphene (-2630 m2/g), graphene electrodes continue to be investigated as ideal coating of electrodes in enzymatic-based sensors, in order to increase the sensitivity and lower the limit of detection. These types of devices are generally affected by poor selectivity to glucose against other interferents that are present in complex biological media such as sweat and also by a limited lifetime.

A device made of a solution-gated field effect transistor with graphene channel (SG-GFET) fabricated on a plastic flexible substrate, has been reported to react to changes in the pH of the solution, as reported in the publication by B. Mailly-Giacchetti, A. Hsu, V. Vinciguerra, F. Pappalardo, L. Occhipinti, E. Guidetti, S. Coffa, J. Kong, 1'. Palacios "pH sensing properties of graphene solution-gated field-effect transistors" (2013) Journal of Applied Physics 114, 84505 (iattp:;:th or.:11 O. 1063/I.4g 1921 9). This SG-GFET device operates in an ambipolar regime as a pH sensor through a shift of the Dirac point in the device trans-characteristics as a function of the solution pH (Figure I, prior art). indeed, the adsorption of hydroxyl and hydronium ions on the graphene surface due to the charging of the electrical double layer capacitance is responsible For the pH sensing mechanism, while the surface transfer doping mechanism does not have significant influence on the pH sensing response.

In this type of devices, other charged molecules in the solution can interact with the graphene channel, acting as interferents to the pH sensing mechanism.

The present invention aims to overcome or at least alleviate the abovementioned drawbacks by providing an improved method and device for the detection of target analytes in bodily fluid, in particular sweat.

Summary of the Invention

A first aspect of the present invention provides an electrochemical sensor device for detecting at least one target analyte in a biological fluid, the sensor device comprising: at least two electrodes, each comprised of an electrically conductive material; at least one layer of graphene, graphene oxide, reduced graphene oxide or other graphene derivative, the at least one graphene layer having incorporated therein or deposited thereon at least one catalyst to provide a catalytic graphene layer for promoting an oxidation-reduction reaction of the target analyte; and a substrate for supporting at least one of the electrodes and at least part of the at least one catalytic graphene layer; the device further comprising: a semi-permeable membrane formed over at least a part of catalytic graphene layer, the membrane being configured to allow selective movement of the target analyte from the biological fluid to the catalytic graphene layer.

More preferably, the electrochemical cell has three electrodes, a source, gate and drain, to provide a field-effect transistor (FET) that uses an electric field to control the flow of current in the device. Alternatively, the three electrodes may comprise a counter electrode, a reference electrode and a working electrode. The electrodes are preferably planar. Preferably, a voltage differential of 0.10.7V is provided between the source and drain or the reference and the working electrodes.

The substrate is preferably flexible, such as being comprised of one or more layers of poly-ethylene naphthalate (PEN), polyester terephthalate (PET), polyimmide (PI), Kapton, FR4, paper, polydymetil-syloxane (PDMS), polyurethane (PU), thermoplastic polyurethane (TPU), or similar or combinations thereof The graphene layer is preferably 0.335nm to 12 pm in thickmess.

The semi-permeable membrane may comprise a biological or artificial membrane that allows selective movement of a target analyte from the fluid to the graphene layer but blocks movement of interferent species above a predetermined size and/or charge. Selective movement may relate to selective passive movement, for example, diffusion, across the membrane or active transport, For example via a carrier molecule, such as a protein. The membrane is preferably a lipid bilayer, more preferably a modified lipid bilayer as defined below. However, it is to be appreciated that the membrane may comprise other types of membrane materials with ionomers or other polymeric structures suitable for facilitating diffusion of water with a selected analyte or analytes.

The lipid bilayer is preferably modified to include at least one channel or transport molecule to allow selective passive or active transport of the target analyte across the membrane to the catalytic graphene layer while preventing the movement of other molecules such as interferent species. In the context of this disclosure, an interferent is any molecule or species which can affect the true measurement of the actual target analyte concentration in the biological fluid. It is to be appreciated that the channel or transport molecule incorporated into the membrane to allow selective transport across the membrane will depend upon the analyte to be monitored. Preferably, the channel or transport molecule is a naturally occurring biological protein consisting of channel proteins or carrier proteins for the target analyte. For example, being selected from the group consisting of the channel protein Acquaporin and the carrier proteins glucose transporters (GLUTs), sodium-glucose cotransporters (SGLTs) and/or Lactose permease symporters.

The semi-permeable membrane formed over at least a part of the catalytic graphene layer to allow selective transport of the analyte From the fluid to the graphene layer preferably forms an inner membrane and at least one outer membrane is also provided in the device. The outer membrane is preferably less selective than the inner membrane, having a greater porosity than the inner membrane, for example having a porosity of 0.2 microns or less. Any suitable porous material may be used, such as water-absorbent cellulose acetate.

The device of the present invention enables an electrochemical process to occur at the catalytic graphene layer with electrons flowing from one chemical substance (the target analyte) to another, driven by an oxidation-reduction (redox) reaction. A redox reaction occurs when electrons are transferred from a substance that is oxidized to one that is being reduced, with the electrons produced being transferred to the graphene layer producing current which may be read by front-end electronics interfacing with the device of the invention.

Thus, the device provides a non-enzymatic redox reaction occurring in the graphene layer with a particular analyte to be monitored being selectively delivered to the graphene layer. For example, the analyte may be a metabolite such as glucose or lactate or a hormone such as cortisol.

The at least one graphene layer of the device of the present invention is provided with at least one catalyst for promoting the oxidation-reduction reaction of the target analyte. It is to be appreciated that the catalyst provided in or on the graphene layer will depend on the target analyte being monitored. Preferably, the catalyst is incorporated in or deposited on the graphene layer by any suitable chemical or physical deposition technique, such as chemical functionalization, chemical reduction, electrochemical reduction, or mixing with exfoliated graphene flakes and compounding with binder species in solution to formulate a suitable catalytic graphene ink or paste.

For example, the catalyst may comprise a metal or metal oxide nanopartide, nanocomposites, molecularly imprinted polymers or an immobilized hormone-specific monoclonal antibody. For detection of the metabolite glucose, a copper or copper oxide catalyst may be used. Metal oxides, such as Co304 (cobalt oxide), NiO (nickel oxide), Ni(OH)2 (nickel hydroxide), nanocomposites such as PdCu (palladium-copper bimetal), CuO-MWNTs (copper oxide with multi-walled carbon nanotubes), Cus(btc)2 (i.e., copper-benzene tricarboxylic acid), as well as metallic or polymeric compounds such as AgNW (silver nanowires), or molecularly imprinted poly (3-aminophynilboronic acid) may be used for detection of lactate.

Hormones, such as cortisol, may use other types of catalyst nanoparticles deposited on the graphene layer, including metal (e.g. Au nanowires), metal oxides (e.g. Fe203), molecularly imprinted polymers (e.g. poly(glycidylmethacrylate-co ethylene glycol dimethacrylate) (poly(GMA-co-EGDMA)), nanocomposites (e.g. Au-PAN1, Ag0Ag-PANI), or other reagents such as e.g. Dithiobis(succinimidyl propionate), hosting an immobilized cortisol-specific monoclonal antibody (Anti-Gith) The biological fluid for detecting at least one target analyte is preferably sweat.

The electrochemical sensor device according to the first aspect of the present invention preferably includes at least one additional sensor, preferably a temperature or humidity sensor, more preferably both, thereby providing a multifunctional sensor device that is able to take account of changes in the environmental conditions, for example ambient temperature, ambient humidity, body temperature, and/or perspiration-rate through skin.

The device is preferably incorporated into a wearable device. The flexible substrate and catalytic graphene layer enables the device to be formed as a thin and flexible smart wearable device, such as but not limited to an adhesive patch-like, a wristband-like, a chest-belt, an arm-belt or other elastic or adjustable belt-like devices, suitable for maintaining deep contact of the active sensing part of the device with the water-containing epidermis's stratum comeum of skin in the different parts of a human body corresponding to one or more of such device form-factors.

The device may include appropriate interfaces to suitable electronic circuitry to allow a readout of signal produced by the oxidation-reduction (redox) reaction, for instance being connected to a I0 microcontroller by appropriate operational amplifiers, transconductance amplifiers, potent ostat circuits, and analogue to digital converters.

A second aspect of the present invention provides a non-invasive method for detecting target analytes in biological fluid, the method comprising the steps of: absorbing biological fluid onto the substrate of an electrochemical sensor according to the first aspect of the present invention: applying a voltage to the two or more electrodes to provide a potential difference between the at least two electrodes; and measuring electric current flowing between the electrodes to provide an amperometnc reading generated by excess electrons produced by the oxidation-reduction reaction of the target analyte to determine at least a presence of the target analyte in the biological fluid.

Preferably, the method according to the second aspect quantitatively measures the target analyte in the biological fluid, more preferably on a continuous or intermittent basis.

Preferably, the method includes attaching the electrochemical sensor to the skin. The target analyte is preferably a metabolite or hormone, such as glucose, lactate or cortisol.

The non-invasive method preferably includes selective passive or active movement of the analyte to the catalytic graphene layer. The method includes promoting the oxidation-reduction reaction of the target analyte by the one or more catalytic agents incorporated in or deposited on the graphene laver.

A third aspect of the present invention provides a method of fabricating an electrochemical sensor according to the first aspect of the present invention comprising: depositing or incorporating at least one catalytic agent on or in at least one layer of graphene, graphene oxide, reduced graphene oxide or other graphene derivative layer to provide at least one catalytic graphene layer,the catalytic agent promoting an oxidation-reduction reaction of a target analyte for the sensor; assembling at least two electrodes and the at least one catalytic graphene layer onto a substrate; and depositing a semi-permeable membrane over at least a part of the catalytic graphene layer, the membrane being configured to allow selective movement of the target analyte from the biological fluid to the catalytic graphene layer.

The method of the third aspect of the present invention preferably includes selecting suitable lipid vesicles, preferably being unilarnellar lipid vesicles, and induing vesicle fusion to create a lipid bilayer on the catalytic graphene layer to form the semi-permeable membrane. More preferably, at least one type of transport or channel molecule is incorporated into a lipid vesicle, the at least one type of transport or channel molecule being selected dependent on the analyte to be sensed by the device, with vesicle fusion enabling deposition of a lipid bilayer incorporating a selected channel or transport molecule onto the catalytic graphene layer for selective movement of that analyte across the lipid bilayer.

Brief Description

For a better understanding of the present invention and to show more clearly how it may be carried into effect, reference will now be made by way of example only, to the accompanying drawings and following examples, in which: Figure 1 a is a schematic diagram of an example of a prior art device for pH sensing based on a solution-gated graphene field effect transistor on a flexible substrate and Figure lb illustrates the t)Tical device characteristics as a function of the solution pH; Figure 2 is a schematic diagram of the architecture of a solution-gated graphene field effect transistor forming an electrochemical sensor device according to one embodiment of the present invention; Figure 3 is a schematic diagram of the architecture of a solution-gated graphene field effect transistor forming an electrochemical sensor device according to an alternative embodiment of the present invention; Figures 4a-c illustrate a typical prior art lipid bilayer membrane deposition method from a vesicle to a support layer forming a supported lipid membrane; Figures Sit-c illustrates a modified process for the formation of a supported lipid bilayer on the surface of a graphene coated support layer according to an embodiment of the present invention; Figure 6 shows one example of detection of glucose concentration in sweat based on a method and device according to a preferred embodiment of the present invention; Figure 7 shows an alternative example of detection of glucose concentration in sweat based on a method and device according to a preferred embodiment of the present invention; Figure 8 is an embodiment of a wearable biosensor device according to an embodiment of the present invention combining the sensor architecture shown in Figure 3 with a temperature and humidity sensor; Figure 9 is an embodiment of a simplified data acquisition and processing system for interfacing with wearable biosensor device according to the invention, such as the device illustrated in Figure 8; and Figure 10 is shows examples of characterization data obtained with different concentrations of glucose in PBS on repeated set of experiments using a device according to the present invention.

Detailed Description

An example of a prior art sensor device 1 is illustrated in Figure la of the accompanying drawings, the device sensing pH based on a solution-gated graphene field effector transistor on a flexible substrate. The device comprises a reference electrode 2, a silver electrode 4 and a solution-gated field effect transistor with graphene channel (SG-GFET) 6 fabricated on a poly-ethylene naphthalate (PEN) substrate 8 (B. Mailly-Giacchern et al., supra). As discussed in the Background of the Invention above, this device operates in an ambipolar regime as a p1-1 sensor through a shift of the Dirac point in the device trans-characteristics as a function of the solution I3 pH. Figure lb illustrates typical device characteristics as a function of the solution pH, where Vgs is the gate source voltage, ID is the drain current and Vgsp is the value of Vgs where each characteristic corresponding to a specific pH value has the minimum value of the drain current. The adsorption of hydroxyl and hydroni um ions on the graphene surface due to the charging of the electrical double layer capacitance is responsible for the pH sensing mechanism, while the surface transfer doping mechanism does not have significant influence on the pH sensing response. While this device is satisfactory, other charged molecules in the solution can interact with the graphene channel, acting as interferents to the pH sensing mechanism, reducing the accuracy of the sensing.

The present invention aims to address this problem by the provision of a novel detection method and corresponding device architecture which rely on a modified graphene-based electrochemical sensor 10, as illustrated in further detail below in relation to Figures 2 to 9 of the accompanying drawings. The device is made of: a) a modified graphene-based working electrode 12 or graphenebased channel where the graphene is decorated with catalyst particles 14 designed to act as a non-enzymatic sensor of the selected biomarker 22 according to known redox reaction methods, such as for glucose or lactate sensing; b) a semipermeable membrane 16, with selected ion or molecular channels, such as a supported lipid bilayer with glucose transporting channels, cellulose acetate, Nafion, or another equivalent membrane material of either natural or synthetic origin that will allow the biomarker to permeate through the membrane and reach the active sensing area through diffusion or transport mechanisms, while limiting the diffusion of other interferent species 24a, the membrane being deposited on top of the decorated graphene layer 12 via a bespoke process as described hereinbelow; and c) an electronic front-end circuitry (see Figures 8 and 9) that allows a readout of the signals as current flowing thorough the sensing electrode or electrodes in the modified graphene-based sensor device, including conversion of such current into voltage and an associated readout, for instance by a microcontroller 100 via an analogue to digital converter 102.

The device of the invention is advantageous as it allows for long-life operations with no or minimal drift of the performance over time, or where drift of performance should exist, the same can be compensated for from other measurements taking place in proximity of the device, for instance by embedding a temperature and humidity sensor 80 (see Figure 8) in the same substrate 40 as the sensor device(s) described in this invention, to allow measuring of the temperature and/or the humidity of the skin surface and apply suitable compensation algorithms to the measured sensor response designed to remove dependency of the same response in relation to temperature and humidity variations in the medium where the metabolite concentration is to be measured.

Another advantage of the device of the invention is that materials are layered in the device architecture in such a way that no harmful or potentially harmful material gets in contact with the skin surface and therefore no skin irritation is produced by wearing the device in deep contact with skin, even for an extended period of time.

Figure 2 illustrates the components making up the device architecture according to one preferred embodiment of the present invention, corresponding to a solution-gated graphene field effect transistor forming an electrochemical sensor 10. In the illustrated embodiment, the electrochemical sensor has a three-coplanar electrode configuration, composed of: i) a source electrode 11, ii) a drain electrode 13 and (iii) a gate electrode 15 co-planar with the source and drain electrodes. A graphene layer 12 decorated with catalytic particles 14 is provided on a flexible substrate 40, such as polyimmide A semipermeable inner membrane 16 of a biological or artificial nature that is suitable for diffusion of species of interest 22 and blocking the diffusion of interferent species 24a, is provided over the graphene layer, followed by an outer membrane 18 for blocking larger interferent species 24b.

The source electrode 11 is made of an electrically conductive material patterned via subtractive (e.g. lithography) or additive (e.g. printing) methods, including for instance copper metal finished with nickel and gold by electroplating or electroless plating, or yet another material or method known to a person skilled in the art. The drain electrode 13 is also made of an electrically conductive material obtained with a similar process as for the source electrode I I and is maintained at a constant difference voltage potential with respect to the source electrode. For example, a

IS

voltage difference selected in the range of 0.1V-0.7V may be maintained between the drain and the source electrode.

The first layer made of grapheme 12 (which may comprise graphene oxide, reduced graphene oxide, single layer graphene or multilayered graphene) is decorated with catalytic metal or metal oxide particles 14 selected among those that are blown to catalyze the oxidation or reduction of the targeted metabolite 22 (further explanation of which is provided with reference to Figures 6 and 7 below). The layer 12 preferably has a thickness between 0.335 nanometers (one single atom layer of carbon) and 10 micrometers (a few layers of graphene) and acts as a channel in the solution-gated field effect transistor. The patterned metal gate electrode 15, coplanar with the source and drain electrodes, is driven by the front-end electronics to produce an electric field inducing a current flow in the graphene channel 12 that depends on the amount of excess charges (e.g. electrons) generated in the channel by means of red-ox reactions, occurring in the same channel when the target metabolite or hormone reaches its surface and reacts with the catalyst particles 14/22 decorating the graphene layer, and therefore correlates to the concentration of the targeted metabolite or hormone in sweat 30.

The semipermeable inner membrane 16 is suitable for diffusion of species of interest 22 but blocks the diffusion of interferent species 24a, this membrane being of biological or artificial nature. For example, the inner membrane may be a supported lipid bilayer membrane and/or other membrane materials with ionomers or other polymeric structures suitable for facilitating diffusion of water with selected metabolites, hormones or ions. The outer membrane 18 is less selective, for example having a porosity of 0.2 microns or less, suitable to act as a blocking layer for larger interferent species 24b present in sweat 30. One such material suitable for this purpose is water-absorbent cellulose acetate.

The inner membrane 16 being a supported lipid bilayer acts to protect the graphene channel from direct interaction with the solution 30 and the interferent species 24a contained in it, and also allows specific ions or molecules 22 to permeate the membrane and reach the graphene channel layer 12, resulting in a change of electrons flow that allow an amperometric readout of the sensor, as a result of the changes in the channel, due to a non-enzymatic redox reaction occurring in the surface of a catalyst-decorated graphene channel 12 (see Figures 6 and 7 below).

The substrate 40 and an insulating layer (polymer) 45 may partially surround the sensor to expose it to skin, with appropriate electronics being provided on another part of the substrate. An appropriate case (not shown) may also be provided for the device to enable its attachment to the wearer, ensuring continuous contact with the skin so that metabolites in sweat 30 of the wearer may be monitored by the sensor.

The method and device of the invention is advantageous versus other methods and devices known in the art as the inner membrane 16 acts as a barrier in two directions: a) preventing unwanted interferent species 24a moving from the solution 30 (water-containing outer skin layer) to the graphene channel 12 as the sensing layer, hence providing great selectivity to the device, b) preventing potential movement of possible toxic or skin-irritating functional catalysts 14 from the graphene channel back to the solution (water-containing outer skin layer), hence maintaining a great biocompatibility and avoiding skin irritation, and c) acting as an encapsulation layer for the graphene channel 12 and the catalytic species 14 contained in it, hence enabling long operational fetime.

Figure 3 of the accompanying drawings illustrates another embodiment of the device architecture of the present invention. Identical features already described in relation to Figure 2 are provided with the same reference numerals for the sake of simplicity and only the differences will be discussed in detail.

The device again comprises a three-coplanar electrode configuration, composed of i) a working electrode 32 coated with a first layer 12 made of graphene, graphene oxide, reduced graphene oxide, single layer graphene or multilayered graphene, decorated with catalytic metal or metal oxide particles 14 selected from those that are known to catalyze the oxidation or reduction of the targeted metabolite or hormone, preferably with a thickness between 0.335 nanometers (one single atom laver of carbon) and 10 micrometers (few layers graphene); ii) a reference electrode 34 with coating 34c, sometimes also known as pseudo-reference electrode, which establishes a stable reference electrochemical potential in the electrochemical cell against the working electrode, preferably between 0.1 and 0.6V; and iii) a counter electrode 36, made of a metal, preferably coated with a precious metal such as gold or platinum. The device is again provided with a semipermeable inner membrane 16 over the working electrode, the membrane being suitable for diffusion of species of interest 22 and blocking the diffusion of interferent species 24a and an outer membrane 18, preferably with a given porosity of 0.2 microns or less, such as water-absorbent cellulose acetate, suitable to ack as a blocking layer for larger interferent species 24b present in sweat 30 Similar to the embodiment shown in Figure 2, when the membrane is a supported lipid bilayer, the supported lipid bilayer acts by: i) protecting the graphene-modified working electrode from direct interaction with the interferent species 24a contained in sweat 30, and it) allowing specific ions or molecules 22 to permeate the membrane 16 and reach the graphene-modified electrode 32, resulting in a change of electrons flow that allow an amperometric readout of the sensor, as a result of the changes in the same graphene-modi fled electrode, due to a non-enzymatic redox reaction occurring in the surface of the catalyst-decorated graphene electrode layer. The inner membrane 16 again acts as a barrier in two directions: a) preventing potential movement of unwanted interferent species 24a from the solution 30 (water-containing outer skin layer) to the graphene sensing layer 12, hence providing great selectivity to the device, b) preventing movement of potentially toxic or skin-irritating functional catalysts 14 from the graphene layer 12 back to the solution 30 (water-containing outer skin layer), hence maintaining a great biocompatibility and avoiding skin irritation, and c) acting as an encapsulation layer for the graphene-modi fled electrode and the catalytic species contained in it, hence enabling long operational lifetime. Is

Figures 4a-c and 5a-5c of the accompanying drawings illustrate examples of lipid bilayer membranes that may be used for the inner membrane 16 shown in the embodiments of Figures 2 and 3. Figure 4 shows a typical, conventional lipid bilayer membrane deposition method from a vesicle 160 to a support layer 8 wherein the vesicles are caused to break 1606 forming a planar supported lipid membrane 16, as described by Kobayashi, T., Kono, A., Futagawa, M., Sawada, K., Tero, R. "Formation and Fluidity measurement of supported lipid bilayer on polyvinyl chloride membrane in ATP. Conference Proceedings 1585 (2014), 145-152.

However, in a more preferred embodiment, particularly for when the support layer is the graphene channel of the sensor device, a modification to this known process is required in order to achieve the required adhesion and maintain a desired selective permeation characteristic of the membrane 16 to the species of interest 22.

An embodiment of this modified process is illustrated with reference to Figures 5a-c. Unilamellar lipid vesicles 200 comprising spheres of liquid 202 surrounded by a lipid bilayer 216 are incubated in a solution containing transporter species 206 for a targeted biomarker (Figure 5a) until bespoke species transport channels 206 are formed in the modified lipid vesicles 216/206 (see Figure 5b). Vesicle fusion is then induced using osmotic shock or temperature incubation resulting in multiple neighboring modified vesicles 216/206 fusing together and eventually breaking 216/206b and precipitating on the substrate, namely on the catalyst containing graphene channel 12 forming a supported lipid bil ay er membrane 16.

It is to be appreciated that this modified process for the formation of a supported lipid bilayer on the surface of a graphene coated support layer may be adapted to different characteristics of the membrane in terms of permeability to the targeted biomarkers (e.g. glucose) and rejection of interferent species (e.g. ascorbic acid and other species present in sweat), as a person skilled in the art can easily derive. An example of such a process for producing a lipid bilayer incorporating selected target transport proteins for passage of specific biomarkers is provided below, together with examples of the detection of the metabolites, glucose and lactose, using the method and device of the present invention.

Example 1: A process for forming a supported lipid bilayer on a graphene channel according to an embodiment of one aspect of the invention.

A suitable lipid, such as 1.2-di oleoyl-sn-glycero-3-ph osphocholine (DOPC) and/or I -palmitoy1-2-oleoyl-glycero-3-phosphocholine (POPC) was suspended in Chloroform at 25mg/m1 * 1,2-di ol eoyl3-trimethylammonium-propane (DOTAP) was also suspended in Chloroform at 25mg/ml. To form the lipid vesicles, 4000 of the desired lipid in Chloroform, either DOPC or POPC, was mixed with 100R1 DOTAP in a small glass vial, forming a 4:1 ratio of lipid to DOTAP.

The bulk chloroform in the vial was evaporated using a nitrogen gas stream, with any remaining chloroform removed under vacuum using a vacuum desiccator for 1 hour. The remaining dried lipid was resuspended in 1:5ml Phosphate Buffered Saline (PBS) and stored at -20°C for 24 hours. After 24h, lml of lipid suspension was extruded using 20 passes through a 50nm pore, polycarbonate membrane to form unilamellar lipid vesicles (200 in Figure 5b). The filtered solution was then stored at 4°C for up to two weeks.

100R1 of the DOPC/DOTAP lipid vesicles in PBS filtered solution was added on top of the graphene-containing substrate using a cured polydimethhylsiloxane (PDMS) well temporarily attached to the substrate around the active graphene-containing sensing area. This facilitates confinement of the solution onto the same active sensing area of the device.

If required, 100R1 of Poly-Ethylene-Glycol (PEG) solution was then added to the substrate in the well and left for 10 minutes to induce vesicle fusion in any unruptured vesicles, after which the 100p.1 of PEG solution was removed, before washing the substrate gently twice with 100p.1 of PBS before adding 100111 of PBS.

In one embodiment, lipid vesicle fusion to create a lipid bilayer on a graphene substrate was induced using a temperature-based protocol. This involved the incubation of the DOPC/DOTAP lipid vesicles on the substrate at 60°C for 1 hour.

In another embodiment, prior to the addition of the lipid vesicles to the graphene substrate, the graphene-containing substrate was soaked in 100 nil of De-ionised (DI) water for 24 hours, as this results in improving lateral mobility in the supported lipid bilayer formed as described above.

In yet another embodiment, 4:1 lipid suspension of DOPC/DOTAP is replaced with a 4:1 lipid suspension of POPC/DOTAP.

Example 2: Detection of Glucose concentration in sweat using a method and device according to one embodiment of the invention.

Aspects of the invention provide a device and method for the detection of glucose G concentration in sweat 30. As illustrated in Figure 6 of the accompanying drawings, the catalyst particle 14 is a catalytic metal or metal oxide nanoparticle that promotes the oxidation of glucose G into gluconolactone GL, with the production of excess electron charges e-that are transferred and read amperometrically via the graphene-modified electrode or graphene channel 12 of the device/method of the invention.

For instance, when the metal catalyst 14 is based on copper (Cu), the following reactions are likely to occur in presence of OH-ions, that fomi the basis of non-enzymatic sensing of glucose G: (a) (b) Cu + 2 OH- CuO + H20 + 2 e- Cu + 2 OH--> Cu(OH)2 + 1-120 + 2 e- CuO + OH--> Cu0OH + e- Cu(OH)2 + OH--> Cu0OH + H20 + e-Either way (route (a) or (b)), the resulting Cu0OH reacts as oxidizing agent for glucose G and converts it into gluconolactone GL according to the following reaction.

Cu0OH + glucose Cu(OH)2+ gluconolactone + e-The electrons e-produced as result of the reaction are then transferred into the graphene-modified electrode or graphene channel 12 producing a current that is read by front-end electronics, while the oxidited gluconolactone molecules GL are dissolved in the solution and eventually converted into gluconic acid by hydrolyzation.

Example 3: Detection of Lactate concentration in sweat using the method and device according to one embodiment of the invention.

Figure 7 illustrates another embodiment of a device/method of the present invention adapted for the detection of lactate L concentration in sweat 30, where the catalyst particle 14 promotes the oxidation of L-lactone L into pyruvate P. with the production of excess electron charges that are transferred and read amperometrically via the graphene-modified electrode or graphene channel 12. Examples of catalyst particles 14 that can be used in this embodiment are metal oxides, such as Co304 (cobalt oxide), NiO (nickel oxide), Ni(OH)2 (nickel hydroxide), nanocomposites such as PdCu (palladium-copper bimetal), CuO-MWNTs (copper oxide with multi-walled carbon nanotubes), Cu3(btc)2 (i.e., copper-benzene tricarboxylic acid), as well as metallic or polymeric compounds such as AgNW (silver nanowires), or molecularly imprinted poly (3-aminophvnilboronic acid).

It is to be appreciated that the catalyst incorporated into the device/method of the invention may be selected based on the active molecule to be monitored by the invention. Similar approaches to those shown in figures 6 and 7 can be adopted for other metabolites or hormones, such as Cortisol, with the adoption of similar or different classes of catalyst nanoparticles, including metal (e.g. Au nanowires), metal oxides (e.g. Fe203), molecularly imprinted polymers (e.g. poly (glyci dylmethacrylate-co ethylene glycol dimethacry I ate) (poly (GMA-co-EGDMA)), nanocomposites (e.g. Au-PANT, Ag0Ag-P AND, or other reagents such as e.g. Dithiobis(succinimidyl propionate), hosting an immobilized cortisol-specific monoclonal antibody (Anti-Camb). These molecules will serve as anchoring site for the cortisol promoting its electrochemical reduction with liberation of 2 hydrogen ions and 2 electrons for each cortisol molecule that is reduced upon binding to the Anti-Cmab antibody: cortisol (ox) cortisol(red)+ 211+ + 2e -Example 4: A Graphene-based Sensor incorporating a temperature and humidity sensor for interfacing with a microcontroller based electronic system according to an embodiment of the present invention.

The graphene-based sensor 10 according to the invention may be combined with a temperature and humidity sensor 80. Figure 8 illustrates one embodiment of such a device, with the graphene-based sensor 10 as hereinbefore described in relation to Figure 3 combined with a temperature and humidity sensor 80. Features already described in relation to Figure 3 are identified by the same reference numerals.

The temperature and humidity sensor may be used to calibrate the device and compensate for variations in the reaction kinetics, as well as to calibrate the sensor response with respect to humidity values that correspond to the presence of sweat in the same region under the outer membrane 16 where the electrochemical sensor is located, according to the proposed data acquisition and processing method.

In this respect, it is known that non-enzymatic electrochemical reactions are affected by changes in the temperature of the solution where the reactions take place. Also, the accuracy of the reading is dependent on the amount of biological solution present in the active region of the graphene-based sensor. Therefore, in this embodiment of the invention, a temperature and humidity sensor 80 is integrated that will provide the following advantages: a) data extracted from the humidity sensor allows the establishment of a baseline where the measurement can be taken reliably from the graphene-based sensor related to the species of interest, and b) data extracted from the temperature sensor may be used to compensate the drift of the sensor response due to the reaction kinetics in order to improve the accuracy of the sensor output in relation to the actual concentration of the species of interest in sweat.

Figure 9 shows a non-limiting example of a simplified data acquisition and processing system and data processing where the sensor data collected from the graphene-based sensor device 10 is combined together with data collected from the temperature and humidity sensor device 80 for practical application in microcontroller-based electronic system, such as a wearable multifunctional patch-like device. The graphene-based sensor 10 and the temperature and humidity sensor 80 are each provided with an interface 90, 92 for connecting to a microcontroller unit 100 via analogue-to-digital convertors 102. The system is also provided with an appropriate communication unit and antenna 96 and energy and power management components 98.

Figure 10 shows examples of characterization data obtained with different concentrations of glucose in PBS. These were obtained using a lab-prototype of the sensor of the invention, for repeated experiments determining the presence of glucose in PBS solution at different concentration levels.

The device and method of the invention enable measurement of the concentration of targeted biomarkers in complex biological fluids, such as sweat, with high sensitivity and minimal or no interference effects to the active detection area of the device from other substances that are present in the complex biological fluid along with the targeted metabolites and/or metal ions.

The same device may be produced in a thin and flexible form suitable for integration in a smart wearable device, such as but not limited to an adhesive patch-like, a wristband-like, a chest-belt, an arm-belt or other elastic or adjustable belt-like devices, suitable for maintaining deep contact of the active sensing part of the device with the water-containing epidermis's stratum comeum of skin in the different parts of a human body corresponding to one or more of such device form-factors.

The device may be used for continuous monitoring of biomarkers produced by human subjects affected by chronic conditions for the purpose of informing their wellness or health-related status.

The device and method of the invention overcome a problem not addressed in the prior art of reliably measuring with a wearable device small concentration of biomarkers that are present in a complex biological fluid, along with other substances, such as the water-containing outer part of skin in a human body, in a way that the measurement is done continuously for a long period of time (limited for instance by energy stored in a rechargeable thin-film battery providing the power needed for the wearable device to operate) with no significant degradation of the sensor performance, and is resilient to multiple artefacts such as those introduced by movements of the person wearing the device, changes in the environmental conditions, for example ambient temperature, ambient humidity, body temperature, and perspiration-rate through skin, as well as to the concentration of the other substances than the targeted metabolites in the same biological fluid.

While several embodiments of the disclosure have been shown in the drawings and accompanying examples, it is not intended that the disclosure be limited thereto, as it is intended that the disclosure be as broad in scope as the art will allow and that the specification be read likewise. Therefore, the above description should not be construed as limiting, but merely as exemplifications of preferred embodiments. Thus the scope of the embodiments should be determined by the appended claims and their legal equivalents, rather than by the examples given.

Claims (20)

1. CLAIMS: I An electrochemical biosensor device for detecting at least one target analyte in a biological fluid, the sensor device comprising: at least two electrodes, each comprised of an electrically conductive material; at least one layer of graphene, graphene oxide, reduced grapheme oxide or other graphene derivative, the at least one graphene layer having incorporated therein or deposited thereon at least one catalytic agent to provide a catalytic graphene layer for promoting an oxidation-reduction reaction of the target analyte; and a substrate for supporting at least one of the electrodes and at least part of the least one catalytic graphene layer; the device further comprising: at least one semi-permeable membrane formed over at least a part of the catalytic graphene layer, the membrane being configured to allow selective movement of the target analyte from the biological fluid to the catalytic graphene layer.
2. 2 The electrochemical biosensor device as claimed in claim 1 wherein the electrochemical cell has three electrodes selected from a source, gate and drain and a counter, reference and working electrode.
3. 3 The electrochemical biosensor device as claimed in claim I or claim 2 wherein the substrate is flexible, being comprised of at least one layer of poly-ethylene naphthalate (PEN), polyester terephthalate (PET), polyimmide (PI), Kapton, FR4, paper, polydymetilsyloxane (PDMS), polyurethane (PU), thermoplastic polyurethane (TPU) or combinations thereof
4. 4. The electrochemical biosensor device as claimed in claim 1, 2 or 3, wherein the semipermeable membrane comprises a biological or artificial membrane that allows selective movement of a target analyte from the fluid to the catalytic graphene layer but blocks transport of interferent species above a predetermined size.
5. 5. The electrochemical biosensor device as claimed in claim 4 wherein the membrane is a lipid bilayer, more preferably a modified lipid bilayer.
6. 6 The electrochemical biosensor device as claimed in claim 5 wherein the lipid bilayer is modified to include at least one channel or transport molecule to allow selective movement of the target analyte across the membrane to the catalytic graphene layer while preventing the movement of interferent species.
7. 7 The electrochemical biosensor device as claimed in claim in claim 6 wherein the channel or transport molecule is selected from the group of naturally occurring biological proteins consisting of channel proteins or carrier proteins for the target analyte, preferably being selected from the group consisting of the channel protein Acquaporin and the carrier proteins glucose transporters (GLUTs), sodium-glucose cotransporters (SGLTs) and/or Lactose permease symporters
8. 8 The electrochemical biosensor device as claimed in any one of the preceding claims wherein the semi-permeable membrane formed over at least a part of the catalytic graphene layer to allow selective movement of the analyte from the fluid to the graphene layer forms an inner membrane and at least one outer membrane is provided in the device, the outer membrane being less selective than the inner membrane, having a greater porosity than the inner membrane.
9. 9 The electrochemical biosensor device as claimed in any one of the preceding claims, wherein the catalytic agent is selected from at least one of the following catalysts; a metal or metal oxide nanoparticle, nanocomposites, molecularly imprinted polymers and immobilized hormone-specific monoclonal antibodies.
10. The electrochemical biosensor as claimed in claim 9 wherein the catalyst is selected from the group consisting of: copper or copper oxide, Co304 (cobalt oxide), NiO (nickel oxide), Ni(OH)2 (nickel hydroxide), PdCu (palladium-copper bimetal), CuO-MWNTs (copper oxide with multi-walled carbon nanotubes), Cu3(btc)2 (i.e., copper-benzene tricarboxylic acid), metallic or polymeric compounds including AgNW (silver nanowires), molecularly imprinted poly (3-aminophynilboronic acid), Fe203, molecularly imprinted polymers poly(glycidylmethacrylate-co ethylene glycol dimethacrylate) (poly(GMA-co-EGDMA)), Au-PANT, Ag0Ag-PANT, and Dithiobis(succinimidyl propionate), hosting an immobilized protein-specific monoclonal antibody (Anti-Cniab).
11. 11. The electrochemical sensor device as claimed in any one of the preceding claims further comprising at least one additional sensor, preferably a temperature or humidity sensor.
12. 12. The electrochemical biosensor device as claimed in any one of the preceding claims incorporated into a wearable device selected from an adhesive patch, a wristband, a chest-belt, an arm-belt or other elastic or adjustable belt-like device.
13. 13 A non-invasive method for detecting target analytes in biological fluid, the method comprising the steps of: absorbing biological fluid onto the substrate of an electrochemical sensor according to any one of claims 1 to 12; applying a voltage to the two or more electrodes to provide a potential difference between the at least two electrodes; and measuring electric current flowing between the electrodes to provide an amperometric reading generated by excess electrons produced by the oxidation-reduction reaction of the target analyte to determine at least a presence of the target analyte in the biological fluid.
14. 14. The method according to claim 13 further comprising quantitatively measuring the target analyte in the biological fluid.
15. 15. The method according to claim 13 or claim 14 wherein the target analyte is a metabolite or a hormone, preferably being selected from glucose, lactate and cortisol.
16. 16 The method according to any one of claims 13 to IS further comprising selective passive or active transport of the analyte across the semi-permeable membrane to the graphene layer.
17. 17. The method according to any one of claims 13 to 16 further comprising sensing one or more additional parameters associated with the biological fluid.
18. 18 A method of assembling an electrochemical sensor according to any one of claims Ito 12, the method comprising: depositing or incorporating at least one catalytic agent on or in at least one layer of graphene, graphene oxide, reduced graphene oxide or other graphene derivative layer to provide at least one catalytic graphene layer, the catalytic agent promoting an oxidation-reduction reaction of a target analyte for the sensor; assembling at least two electrodes and the at least one catalytic graphene layer onto a substrate; and depositing a semi-permeable membrane over at least a part of the catalytic graphene layer, the membrane being configured to allow selective movement of a target analyte from a biological fluid to the catalytic graphene layer.
19. 19. The method according to claim 18 further comprising selecting lipid vesicles and inducing vesicle fusion to create a lipid bilayer on the catalytic graphene layer.
20. 20. The method according to claim 19 further comprising incorporating at least one type of channel or transport molecule into the lipid vesicle, the channel or transport molecule being selected dependent upon the analyte to be sensed by the sensor, whereby vesicle fusion enables deposition of a lipid bilayer incorporating a selected channel or transport molecule into or onto the catalytic graphene layer.