Bioanalyte detection and monitoring Technical field of the invention

The invention relates to a sensor system for bioanalyte detection in a medium of interest, particularly a body fluid, such as blood or interstitial fluid.

Background of invention

Transdermal wearable biosensors capable of interacting with the skin which contains biofluids like sweat, interstitial fluid (ISF), and peripheral blood rich in valuable biomarkers are desirable. Wearable sensor device technologies, which enable continuous monitoring of biological information from the human body, are promising in the fields of sports, healthcare, and medical applications. Further thinness, light weight, flexibility and low-cost are significant requirements for making the devices attachable onto human tissues or clothes like a patch.

Skin is composed of three layers starting with the epidermis that is the outermost layer with an average thickness of 150 pm, continuing with the dermis with a thickness ranging between 500 to 2000 pm and finally the hypodermis. While the epidermis does not contain any blood vessels and nerves, the lower part of the dermis and hypodermis contain nerve bundles and blood vessels. The interstitial space is the area surrounding parenchymal cells, vascular and lymphatic capillaries and this space is filled with interstitial fluid (ISF). ISF carries nutrients form blood vessels to the cells and also waste components from cells to the lymph vessels and therefore is a composition of serum and cellular materials including ions, proteins, and small molecules such as glucose and lactate. Due to the high similarity (93%) of ISF constituents to the ones found in blood serum and plasma, ISF can be used to monitor different biomarkers for disease diagnostics.

Transdermal monitoring via microneedles (MNs) or microneedle arrays (MNAs) which commonly target interstitial fluid (ISF) or peripheral blood are good alternatives. MNs enable painless needle insertion into the skin, are less invasive compared to the conventional needles, and provide higher interacting surface area. Moreover, tunable physical properties of MNs, like length, sharpness, and density, enables MNs to reach different depths of epidermal and dermal layers without stimulation of dermal nerves thus not causing pain or discomfort. MNs fabricated from silicon are particularly desirable due to silicon’s biocompatibility, high rigidity and mechanical strength that increase skin penetration, prevents MN breakage and accumulation inside the skin which can produced undesirable immune responses or other related health problems. Silicon has an elastic modulus of about 50 GPa to about 180 GPa, which is more than enough to withstand the force needed to pierce the skin. This is particularly important for the application of transdermal biosensors since one of the main challenges, next to biocompatibility, in designing MN is mechanical strength in order to prevent MN breakage inside the skin.

One aspect of the invention is to provide patients with painless bioanalyte or pH monitoring methods involving ISF instead of blood as a medium. Bioanalytes include glucose, lactate, ascorbic acid and biomarkers such as antigens, preferably those which are diagnostic or prognostic of disease, for example, cancer. Electrochemical based sensing, such as in vivo pH sensing or in vivo detection of bioanalytes in ISF requires a sensitive, selective and robust system of suitably functionalised electrodes which are mechanically strong enough to pierce the skin to reach various dermal layers of interest and also which can withstand friction associated with transdermal use. This is particularly the case where a complex sensing chemistry or system is applied to the electrode surfaces in the case of many working electrodes.

A reference herein to a patent document or other matter which is given as prior art is not to be taken as an admission that the document or matter was known or that the information it contains was part of the common general knowledge as at the priority date of any of the embodiments. Where any or all of the terms "comprise", "comprises", "comprised" or "comprising" are used in this specification (including the embodiments) they are to be interpreted as specifying the presence of the stated features, integers, steps or components, but not precluding the presence of one or more other features, integers, steps or components.

Statement of the invention

In a first aspect, the invention provides a working electrode comprising an array of solid, non hollow skin piercing microprojections at an array density of at least 2,000 microprojections/cm ² disposed on an electrode substrate, wherein the substrate, and each of the microprojections are coated with a conformal nanofilm of at least one biocompatible electrically conducting, chemically inert material which supports a redox reaction.

In a second aspect, the invention provides a working electrode comprising an array of solid, non-hollow skin piercing microprojections at an array density of at least 2,000 microprojections/cm ² disposed on an electrode substrate which is provided onto an electrode and connector support having a proximal portion comprising the sensing electrode substrate of the electrode and a non sensing distal portion for connection to one or more connectors, contacts or leads for transfer of electrical signal, wherein the substrate, each of the microprojections, and the support are coated with a conformal nanofilm of at least one biocompatible electrically conducting, chemically inert material which supports a redox reaction having a thickness of a thickness of from 5 to about 500 nm, and wherein non-sensing parts of the working electrode are coated with an electronically and ionically insulating material.

Preferably, one or more of the microprojections have an electrically conductive surface which is formed from nanolayer layer or nanofilm of a chemically inert but electrically conductive material. Preferably, the nanolayer layer or nanofilm is a conformal nanolayer or nanofilm of the chemically inert but electrically conductive material. The conformal film is a contiguous/continuous or integral nanofilm across the surface of all components of the electrodes as described herein. Suitably, the electrode substrate, each of the microprojections and the support are coated with the conformal nanofilm. The support has a proximal portion comprising the sensing electrode substrate of the electrode and a non sensing distal portion for connection to one or more connectors, contacts or leads for transfer of electrical signal.

Suitably, the support further comprises a bridging portion disposed between the distal portion and the proximal portion which adapted to transfer an electrical current from the microprojections to the distal portion. Suitably, the bridging portion is adapted with the same nanofilm of conducting material as is provided on the sensing portions of the working electrode.

Preferably, the non-sensing parts of the working electrode are coated with an electrically insulating material. The non-sensing parts are the non-sensing distal portion for connection to one or more connectors, contacts or leads for transfer of electrical signal and any intermediate or bridging portions between the sensing substrate/microprojections and the connectors, contacts or leads.

Desirably, the bridging portion is an elongate bridging portion in the form of an elongate body with the electrode substrate at a head of the elongate body and the connectors, contacts or leads are at a tail of the elongate body.

Preferably, the bridging portion is also coated with the conformal nanofilm of at least one biocompatible electrically conducting, chemically inert material and further coated with the electrically insulating material. In some embodiments the electrically insulating material is a non electronically conducting polymer, an electronic conduction insulator and/or an ionic conduction insulator material, such as a dielectric ink.

Suitably, the support is a flexible support that bends and adopts to skin contours. The flexibility is also useful for aiding finger pressure skin piercing of the high density microprojection array of the invention.

Preferably, the support is mounted on a base in the form of a fixing frame or a fixing board for mounting, positioning and fixing one or of supports (and associated substrates/arrays) in a desired configuration. Desirably the fixing frame or a fixing board is flexible and bends and adopts to skin contours. The flexibility is also useful for aiding finger pressure skin piercing of the high density microprojection array of the invention.

Desirably, the electrode and connector support is formed from a metal such as gold or silicon or in some embodiments may be a 3D printed support, for example, formed from a photocurable polymer such as OrmoComp®. In embodiments utilising a sputtered gold conducting nanofilm, a 3D printed support is preferred for reduced cost reasons as 3D printed materials are inexpensive and support bulk manufacturing. In embodiments utilising a thermally carbonised conducting nanofilm, it is preferred that the support is a metal support, particularly a silicon support, which can be readily carbonised, particularly thermally carbonised under high temperature conditions. Likewise, in embodiments utilising a conducting polymer nanofilm such as a polyaniline conducting nanofilm, it is preferred to use a support material that is compatible with electropolymerisation or electrodeposition of the desired conducting polymer nanofilm , such as a conducting metal or polymer material that supports the electropolymerisation reaction. Desirably, the microprojections, the substrate and/or the support are formed from a metal such as silicon, or a polymer such as a curable or crosslinkable polymer. In the case of the polymer microneedles, preferably, the polymer is one with mechanical properties suitable for skin piercing. Suitably, such polymers will have an elastic modulus of at least 1 GPa.

Suitably, the redox reaction involves (i) electrons from direct oxidation or reduction of a target analyte at the nanofilm surface, (ii) electrons from a cascade reaction involving the target analyte, or (iii) electrons from a redox active species present in the medium of interest.

Preferably, the nanofilm has a thickness of from 5 to about 500 nm, preferably about 150 nm.

Suitably, the conformal nanofilm is in the form of a metal film. Desirably, the conformal nanofilm is in the form of a metal film comprising a surface metal and an adhesive undercoat of a metal. In some embodiments, the adhesive undercoat prevents delamination of the surface metal which is important for a skin piercing sensor. Desirably, in one embodiment the metal film is a gold film, preferably a sputtered metal film, particularly a sputtered gold film. Desirably, the adhesive undercoat is a sputtered metal or an electrodeposited metal. Preferably, the adhesive undercoat is chromium.

In embodiments where the electrode surface is a conformal gold or platinum nanofilm, it is preferred that biorecognition elements, when present, are bound to functional groups of a self- assembled monolayer of organic molecules on surfaces of the nanofilm of the working electrode. Desirably, the self-assembled monolayer on surfaces of the conformal nanofilm of the working electrode is derived from a thiol containing organic molecule such polyaniline, or as an alkane thiol or a dialkene thiol, most preferably 3-mercaptopropionic acid multiwalled nanotubes (MWCNT) or a mixture of polyaniline and graphene or MWCNT. The self-assembled monolayer presents at least carboxylic acid groups on the conducting surface which can be activated to enable immobilisation of biorecognition elements thereon. Carboxylate activating agents such as an EDC/NHS couple can be used for such immobilisation. Preferably, after biorecognition element immobilisation onto the SAM, the surfaces of the conformal nanofilm of the working electrode are modified with 2- mercaptoethanol.

In one embodiment, the conformal nanofilm is in the form of a porous carbonized surface region. Suitably, the porous carbonized surface region is a porous thermally carbonized layer. Such layers have been found to have very high surface areas. Desirably, the pores of the carbonized porous surface layer have pore diameters of <100 nm, preferably 50 to 60 nm. Suitably, the carbonized porous surface layer is obtained by thermal treatment of a porous metal microarray needles/projections at high temperature of from 400-800 °C in the presence of acetylene. Pores in the needles/projections can be formed by anodization or electrochemical etching using known methods. Metals such as silicon etches with nanopores are preferred for thermal carbonisation. Where a carbonised nanofilm is used, it is preferred that biorecognition elements are immobilised within at least the nanopores via for example electrografting of a benzoic acid, preferably amino benzoic acid (ABA) in the presence of a diazonium salt grafted on the carbon nanofilm. In another embodiment, the conformal nanofilm on at the array substrate and microprojections is in the form of a conducting polymer layer, preferably a plurality of conducting polymer layer, such as electropolymerized conducting polymer layers, most preferably polyaniline. Such layers are preferred for potentiometric sensing, for example, where pH sensing is desired based on changes in the cell potential due to ions (protons or hydroxyl ions) in the medium of interest. Suitably, the conducting polymer layer(s) are electropolymerized or electrodeposited conducting polymer such as polyaniline, which react with hydrogen ions for pH sensing.

Desirably, surfaces of nanofilm of the working electrode are modified with biorecognition elements which are selective for a bioanalyte of interest. Preferred biorecognition elements include one or more of enzymes, antibodies, antigens, and aptamers.

In particular, aptamers can be used for insulin sensing where insulin selective aptamers are used.

Preferably, the enzymes are redox mediator modified enzymes. Desirably, enzymes are redox mediator dendrimer modified enzymes or redox mediator conductive nanoparticle modified enzymes. Suitably, the redox mediator is retained between the conducting material and the enzyme. In some embodiments, the redox mediator is provided at the focal point of the dendrimer. This arrangement is believed to reduce the working potential of the working electrode and in many cases also increases selectivity. Desirably, the dendrimer is immobilised onto the conformal nanofilm of biocompatible electrically conducting, chemically inert material. Preferably, the enzyme molecules are bound to the dendrimer in a single layer. Desirably, the dendrimer is selected from hyperbranched polyglycerol dendrimers, polyethylene glycol dendrimers, glutamic acid dendrimers and PAMAM dendrimers. Preferably, the dendrimer is a GO generation dendrimer or upwards, preferably, a PAMAM dendrimer. Suitably, the dendrimer is a G2 PAMAM dendrimer. Preferably, the redox mediator is provided underneath, or embedded in, a polymer matrix comprising the conductive nanoparticles such as gold nanoparticles; carbon nanotube nanoparticles; reduced graphene oxide nanoparticles or platinum nanoparticles.

In some embodiments, the redox mediator is ferrocene, methylene blue or Prussian Blue.

Suitably, the enzyme is glucose oxidase for glucose sensing, lactate oxidase (LOx) for lactate sensing, urease for urea sensing and alcohol oxidase (AOx) for ethanol sensing. In some cases cholesterol oxidase can be used for cholesterol sensing.

In one embodiment, the redox mediator is a ferrocene-dendrimer modified glucose oxidase (GOx) and the bioanalyte is glucose. Suitably, the sensitivity of the sensor system allows detection of mM variation in bioanalyte levels, particularly glucose levels. Desirably, the redox mediated dendrimer enzyme is selective for its target in a media comprising the target and one or more of from uric acid (UA), ascorbic acid (AA), dopamine (DA), lactic acid (LA) and glycine (gly).

In one embodiment, the redox mediator is a Prussian Blue-AU nanoparticle modified lactase (LOx) and the bioanalyte is lactate.

Suitably, the antibodies are anti-HER2 antibodies, preferably against ErbB2 antigen. Preferably, the microneedles/microprojections are formed from metal, such as silicon or a polymer, preferably a cured polymer having a modulus of at least 1 GPa.

Desirably, the cured polymer is a cured hybrid polymer, such as a cured silica-polymer hybrid or cured polylactide.

In one embodiment, the cured polymer is formed by curing in a mould generated from a microprojection array template, preferably a PDMS mould, for example, formed from a metal microneedle array or microprojection array template as described herein. Use of a template to form a polymeric array provides a more convenient manufacture in terms of speed and reduced cost of reproduction fabrication. It will be understood that such polymeric arrays need to be provided with a nanofilm of conductive material for embodiments involving redox reactions.

In a third aspect, the invention provides an electrochemical sensor system for in vivo pH sensing or for sensing of a bioanalyte in interstitial fluid, the sensor system comprising an electrochemical cell comprising a working electrode as defined in the first or second aspects.

In a fourth aspect of the invention, there is provided a wearable device for pH sensing or for bioanalyte monitoring comprising the sensor system of the third aspect. Desirably, the device is adapted for continuous monitoring. In the continuous monitoring embodiment, the sensor system is retained in place on the skin for a period of up to about 14 days. This embodiment enables real time monitoring of pH or of the bioanalyte. In the case of at least glucose this is desirable because while the concentration of glucose in ISF is identical to that in blood in steady-state conditions, there may be a lag time immediately after glucose intake due to the kinetics of diffusion. This lag time is estimated to be about 5 to about 15 minutes.

In a fifth aspect of the invention, there is provided for a use of an electrochemical sensor system as defined in the third aspect in the transdermal monitoring of pH or of a bioanalyte, preferably glucose, lactate, ethanol, urea, ascorbic acid or a biomarker such as an antigen or antibody.

Desirably, in some embodiments, the transdermal monitoring is continuous inline, in vivo monitoring, which does not require sampling or withdrawal of any bodily fluid during testing. In other embodiments, the sensor system can be used as bioanalyte capture device, such as an immunocapture device, for capturing biomarkers (such as antigens or antibodies) from ISF in the skin, withdrawing or extracting such biomarkers and quantifying same, for example, via a DPV technique.

In a sixth aspect of the invention, there is provided a use of a transdermal sensor system according to the third aspect in a method of transdermal electrochemical pH or bioanalyte monitoring.

In a seventh aspect of the invention, there is provided a use of a working electrode of the first or second aspects, a sensor system of the third aspect, or the wearable device the fourth aspect in the transdermal sensing or transdermal capture and/or monitoring of a bioanalyte. Desirably, this use is to capture an extract a biomarker target from ISF in skin dermal and/or epidermal layers. In an eighth aspect of the invention, there is provided a method of diagnosis or a disease or condition comprising the step of identifying the in vivo presence of a bioanalyte and/or a level of a bioanalyte in ISF or peripheral blood associated with a disease or condition using the sensor system of the third aspect or the wearable device of the fourth aspect.

In a ninth aspect of the invention, there is provided an array of solid, non-hollow microneedles at an array density of at least 2,000 microneedle/cm ² , wherein the array of microneedles is coated with a conformal nanofilm of at least one biocompatible electrically conducting, chemically inert material which supports a redox reaction, and, wherein the conformal nanofilm is modified with a monolayer of a redox mediator modified dendrimer enzyme or a monolayer of a redox mediator modified conducting nanoparticle/polymer composite modified enzyme, wherein the modified enzyme is specific for a bioanalyte of interest.

In a tenth aspect of the invention, there is provided an array of solid, non-hollow microneedles at an array density of at least 2,000 microneedle/cm ² , wherein the array of microneedles is coated with a conformal nanofilm comprising a first adhesive layer of chromium and a second conformal nanofilm of gold over the chromium layer, wherein the conformal nanofilm of gold is modified with a monolayer of a redox mediator modified dendrimer enzyme or a monolayer of redox mediator modified conducting nanoparticle/polymer composite modified enzyme, wherein the modified enzyme is specific for a bioanalyte of interest, and wherein the modified enzyme is immobilised onto the nanofilm via a self assembled monolayer.

In an eleventh aspect of the invention, there is provided an array of solid, non-hollow microneedles having a porous surface layer thereon at an array density of at least 2,000 microneedle/cm ² , wherein the array of microneedles is coated with a conformal nanofilm of carbon, wherein the carbon nanofilm of carbon is modified with a monolayer of a redox mediator modified dendrimer enzyme or a monolayer of redox mediator modified conducting nanoparticle/polymer composite modified enzyme, wherein the modified enzyme is specific for a bioanalyte of interest.

In a twelfth aspect of the invention, there is provided an array of solid, non-hollow microneedles having a porous surface layer thereon at an array density of at least 2,000 microneedle/cm ² , wherein the array of microneedles is coated with a conformal nanofilm of an electropolymerized conducting polymer which reacts with hydrogen ions for pH sending.

Desirably, the conformal nanofilm of a conducting polymer of the twelfth aspect is a polyaniline nanofilm.

Desirably, the solid, non-hollow microneedles are formed from silicon or a cured polymer having an elastic modulus of at least 1 GPa.

In a thirteenth aspect of the invention, there is provided a method of preparing a calibration curve for a bioanalyte in a medium of interest using the electrochemical sensor system of the third aspect, the method comprising the steps of: providing the electrochemical sensor system to at least one solution of a test medium of interest such that electrode surfaces of the microneedles contact the medium of interest, wherein the solution of test medium of interest comprise a known concentration of the bioanalyte; applying a steady state electrical potential to the sensor’s counter electrode, whereby the potential provides a suitable potential window for detection of a redox reaction or chain of reactions initiated by the presence of the bioanalyte; monitoring for changes in an electrochemical property occurring at the sensor’s working (WE) electrode which result from a redox reaction involving the bioanalyte of interest; associating an observed electrochemical property change with the known bioanalyte concentration; and optionally repeating the preceding steps with solutions of a test medium of interest comprising different known amount of the bioanalyte.

The step may further comprise the step of fitting the observed current data and the concentration date to a curve.

In a fourteenth aspect of the invention, there is provided a method of bioanalyte monitoring in a medium of interest comprising the steps of: providing an electrochemical sensor system according to the third aspect to a test medium of interest such that electrode surfaces of the microneedles contact the medium of interest; applying a steady state electrical potential to the sensor’s counter electrode, whereby the potential provides a suitable potential window for detection of a redox reaction or chain of reactions initiated by the presence of the bioanalyte; monitoring for changes in an electrochemical property occurring at the sensor’s working (WE) electrode which result from a redox reaction involving the bioanalyte of interest; establishing the concentration of the bioanalyte present in the medium by comparing a level of a detected electrochemical property at the sensor’s working (WE) electrode with a predetermined calibration curve for the bioanalyte.

In a fifteenth aspect of the invention, there is provided transdermal bioanalyte monitoring comprising the steps of: providing an electrochemical sensor system according to the third aspect to a test subject’s skin; aligning the sensor’s microneedle tips against the skin; applying a skin piercing pressure to the sensor to position the microneedles in skin’s dermal layer thereby allowing electrode surfaces of the microneedles to access the subject’s interstitial fluid; applying a steady state electrical potential to the sensor’s counter electrode, whereby the potential provides a suitable potential window for detection of a redox reaction or chain of reactions initiated by the presence of the bioanalyte; monitoring for changes in an electrochemical property occurring at the sensor’s working (WE) electrode which result from a redox reaction involving the bioanalyte of interest; establishing the concentration of the bioanalyte present in the interstitial fluid by comparing a detected level of the electrochemical property at the sensor’s working (WE) electrode with a predetermined calibration curve for the bioanalyte. Desirably, the electrochemical property is one or more of current, potential, resistance, impedance and capacitance.

In a sixteenth aspect of the invention, there is provided a method of diagnosis or monitoring of a disease or condition comprising the step of identifying the presence of a bioanalyte and/or a level of a bioanalyte associated with a disease or condition, using the sensor system according to third aspect. Desirably, the bioanalyte is glucose and the disease or condition is metabolic disease, insulin resistance, glucose intolerance, diabetes including Type 1 or Type 2 diabetes. Suitably, the bioanalyte is lactate and the disease or condition is sepsis. Desirably, the bioanalyte is a cancer biomarker such as an ErbB2 marker present at a concentration of greater than 15 ng/mL which is diagnostic of breast cancer.

In a seventeenth aspect of the invention, there is provided an array of solid, non-hollow microneedles at an array density of at least 2,000 microneedle/cm ² , wherein the array of microneedles is coated with a conformal nanofilm of at least one biocompatible electrically conducting, chemically inert material which supports a redox reaction, and, wherein the conformal nanofilm is modified with a monolayer of immunocapture biorecognition elements. Desirably, the conformal immunocapture biorecognition elements are antibodies against an antigen biomarker. Brief description of the drawings

Embodiments of the invention will herein be illustrated by way of example only with reference to the accompanying drawings in which:

Figure 1 is a schematic illustration of Au-Si-MNA electrode preparation. A) Si-MNA and 3D printed holder, B) attaching Si-MNA substrate to 3D printed holder, C) sputter deposition of a Au thin film, C) insulating dye application, and E) three-electrode sensing patch with Au-Si-MNA as W and C electrodes and R electrode with Au-Si-MNA coated with AgCI ink;

Figure 2 is an SEM micrograph of Au-Si-MNA; A) top view, B) tilted view, C) cross section, and D) magnified tip area (black square) of MNs showed in C. Confocal microscopy images of E) MNA after coating with Rhodamine labeled Dextran dye and F) porcine skin after MNA application;

Figure 3 is an illustration of the electrochemical characterization of Au-Si-MNA electrode. A) CV plots of Au-Si-MNA (a) and flat Au-Si (b) electrodes recorded in 0.5 M H2SO4 solution, B) CV plots of Au-Si-MNA, Au-Si-MNA and flat Au-Si modified with Fc-PAMAM, and Au-Si-MNA/Fc- PAMAM/GOx electrodes recorded in 0.1 M PBS (pH 7.4), C) CV plot of Au-Si-MNA/Fc-PAMAM electrode at different scan rates, inset; anodic and cathodic peak currents versus the square root of scan rates, and D) EIS of Au-Si-MNA (a), flat Au-Si (b), Au-Si-MNA/Fc-PAMAM (c), and Au-Si- MNA/Fc-PAMAM/GOx (d);

Figure 4 illustrates the electrochemical optimization of MNA glucose patch using 0.1 M PBS with different A) pH and B) temperature. C) Reusability of the biosensor; Figure 5 illustrates A) Interference study using 2 mM glucose and 0.1 mM AA, UA, DA, LA, and gly in 0.1 M PBS, pH 7.4 at applied potential of 0.35 V (n=3). Chronoamperometric measurements of MNA glucose patch ^' s response to glucose in B) PBS (0.1 M, pH 7.4) and C) alSF (pH 7.4), and D) calibration curves obtained in PBS (a) and alSF (b)(n=3);

Figure 6 corresponds to optical images of A) a three-electrode system used as a patch for transdermal monitoring of glucose in mice, B) application of the patch on shaved mice skin, C) mice skin after the application of the needles illustrating successful penetration of all electrodes, D) shows chronoamperometric signal recorded during application of the patch on mice skin before (a) and after (b) glucose injection;

Figure 7 illustrates a schematic diagram of the electron flow in the Fc-PAMAM/GOx biosensor;

Figure 8 illustrates a schematic diagram of the surface modifications of Au-MNA for the immobilization of GOx;

Figure 9 illustrates a schematic diagram of the stepwise modifications on the porous electrode;

Figure 10 illustrates SEM images of PSi-MNAs. A) Cross-sectional image of the MNAs; B) detail of the nanoporous surface with a 50-60 nm pore diameter;

Figure 11 illustrates a cross-sectional SEM images of PSi-MNA after electrochemical etching of a conformal nanoporous film. A-C) Detail of the nanoporous film at the base of the MNA; and D- F) detail of a microprojection cross-section for different etching times: A, D) 4 min; B, E) 6 min; and C, F) 8 min showing a porous layer of 1.6, 2.3 and 4.0 pm thickness, respectively.

Figure 12 illustrates a FT-IR spectra of TCPSi and TCPSi-ABA electrog rafted nanocomposites. Inset a zoomed in spectra.

Figure 13 illustrates a Nyquist impedance spectra of TCPSi electrodes before and after ABA electrografting obtained in a 5.0 mM Fe(CN) ₆ ^{3 /4_} solution containing 0.1 M KCI. Inset shows the equivalent (Randle) circuit used to model the impedance plots. The impedance parameters are shown in the table.

Figure 14 illustrates a cyclic voltammograms of modified TCPSi/ABA/Fc-PAMAM electrodes before (solid line) and after (dotted line) GOx enzyme immobilization.

Figure 15 illustrates CVs of TCPSi/ABA/Fc-PAMAM electrodes recorded in 0.1 M PBS solution at different scan rates: 10, 50, 100, 200, 300, 400 and 500 mV s ^_1 (left); cathodic peak (l _pc ) and anionic peak (l _pa ) vs root scan rates plot (right).

Figure 16 illustrates amperometric response towards increments of glucose concentration on TCPSi/ABA/Fc-PAMAM/GAD/GOx electrodes, n=2. (a) the glucose concentration response from 0 to 6 mM, while (c) shows the glucose concentration response from 0 to 1 mM. (b) the calibration curve for the higher concentration of glucose while (d) shows the calibration curve for the lower concentration of glucose. Figure 17 illustrates a schematic illustrations of a) MNA electrode preparation, b) surface modification and anti-HER2 monoclonal antibody immobilization on Au-Si-MNA electrode, and c) detection principle of MNA-based immunosensor.

Figure 18 illustrates a SEM micrographs of a) tilted view, b) top view (inset: top view of single microneedle) and c) cross-sectional view (inset: cross-sectional view of microneedle tip) of Au-Si- MNA.

Figure 19 illustrates characterization of MNA-based immunosensor. a) CV recorded in a 0.5 M H2SO4 solution (inset: histogram showing a comparison of the measured surface area and calculated active surface area (SD of n=3)). b) DPV and c) EIS plots obtained after each step of the electrode modification: i) Au-Si-MNA, ii) SAM, iii) after immobilization of anti-HER2 antibody, iv) PEG, and v) after ErbB2 binding (150 ng/mL) (insets: of B) histogram showing current comparison of DPV peaks (SD of n=3), and C) Randles equivalent circuit used for fitting). Measurements were performed in 5 mM K^Fe (CN)6]/K3[Fe(CN)6] (1 :1 ratio) in 0.1 M KCI. d) XPS high resolution spectra of the C 1s obtained after each functionalization step: pre-cleaned Au (black); SAM modification (red); and immobilization of anti-HER2 (blue).

Figure 20 illustrates a) DPV plots of control experiments carried out using Au-Si-MNA electrodes modified with IgG (a, b) and anti-HER2 (c, d) antibodies before (a, c) and after (b, d) incubation; inset: histogram illustrates the change in the DPV current density after incubation with ErbB2 (SD of n=3)). ErbB2 breast cancer biomarker detection and quantification with a MNA-based immunosensor. b) DPV of Au-Si-MNA/anti-HER2 sensor incubated in alSF with different concentrations of ErbB2 ranging from 10 to 250 ng/mL and c) calibration curve of Au-Si-MNA anti- HER2 sensor obtained at 0.15 V (n=3). d) Selectivity of Au-Si-MNA anti-HER2 immunosensor upon incubation in 150 ng/mL of ErbB2, 10 ⁵ pfu/mL T4 bacteriophage, 150 ng/mL insulin, alSF, 5 mM glucose, and 100 mM glycine, all prepared in alSF, 123 mM NaCI prepared in dhhO, and 0.1 M PBS (SD of n=3).

Figure 21 illustrates testing of Au-Si-MNA anti-HER2 immunosensor using skin-mimicking phantom gel. a) Calibration curve of DPV signals obtained after incubation of immunosensor in phantom gel (SD of n=3). Optical images of b) Au-Si-MNA electrode and c) 3D printed mold used for the preparation of the phantom gel. Light microscopy images of d) cross section of the phantom gel mimicking epidermis and dermis layers, and e) top view of the phantom gel after applying Au-Si- MNA electrode.

Figure 22 illustrates a A) schematic illustration of PDMS replication process of Si-MNA and B) schematic illustration of microfabrication process of OrmoComp ^® -based MNA.

Figure 23 - Figure 2. SEM micrographs of pMNA tilted (45°) view (inset shows magnified tip area of pMNA).

Figure 24 illustrates A) DPV graph of pMNA Au electrode recorded in alSF in the presence of different concentrations of ascorbic acid (AA) ranging from 100 to 400 mM. B) Calibration curve of pMNA/Au electrode derived from DPV graph shown in Fig. 3A (n=3). C) DPV graph of pMNA/Au based electrode recorded in alSF in presence of AA and uric acid (UA) (red line) and without AA and UA (black line).

Figure 25 illustrates electrochemical detection of lactate using pMNA/Au/Fc-PAMAM/LOx based biosensor. A) Amperometric response of pMNA/Au/Fc-PAMAM/LOx biosensor in 0.1 M PBS at applied potential of 0.35 V and B) calibration curve (n=3).

Figure 26 illustrates a electrochemical detection of lactate using pMNA/Au/PB-ChAuNPs/LOx based biosensor. A) Amperometric response of pMNA/Au/PB-ChAuNPs/LOx biosensor in 0.1 M PBS at applied potential of -0.15 V and B) calibration curve (n=3).

Figure 27 illustrates A) Cyclic voltammetry recorded during electro-polymerization of aniline onto pMNA surface in 0.1 M HCI and 0.2 mM NaF solution. B) Comparison of 20 ^th cycle of electro polymerization process of five different electrodes.

Figure 28 illustrates optimization of pMNA/Au coating procedure by determining optimum A) number of CV cycles performed and B) scan rate applied during electropolymerizing process (SD of n=3). Statistical analysis was performed using One-way ANOVA with P < 0.05. ns stands for not significant.

Figure 29 illustrates A) Potentiometric graph of pMNA/Au/PANI electrode in alSF with different pH ranging from 6.4 to 7.8. B) Calibration curve of pH sensor in alSF derived from Fig. 6A (SD of n=3).

Figure 30 illustrates continuous measurement of changes in aISF ^' s pH ranging from 6.5 to

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Detailed description of the invention

In addition to the aspects described above, further described herein is a working electrode and an electrochemical sensor system for pH sensing, or sensing a bioanalyte in a media of interest and/or capturing and withdrawing a bioanalyte from a tissue or body fluid such as ISF or peripheral blood for subsequent quantification.

A preferred sensor system of the invention comprises an electrochemical cell comprising the following components: a reference (RE) electrode comprising a reference electrode substrate supporting an array of electrically conductive biocompatible microneedles at a density of at least 2,000 microneedle/cm ² disposed on the reference electrode substrate; a counter (CE) electrode comprising a counter electrode substrate supporting an array of electrically conductive biocompatible microneedles at a density of at least 2,000 microneedle/cm ² disposed on the counter electrode substrate; and a working (WE) electrode comprising a working electrode substrate supporting an array of electrically conductive biocompatible microneedles at a density of at least 2,000 microneedle/cm ² disposed on the working electrode substrate; wherein each array of microneedles is coated with a conformal nanofilm of at least one biocompatible electrically conducting, chemically inert material which supports a redox reaction, and wherein the microneedles arrays on each of the reference electrode substrate, the counter electrode substrate and the working electrode substrate are simultaneously contactable with a medium of interest which serves as electrolyte in the cell, and wherein the microneedles arrays on each of the reference electrode substrate, the counter electrode substrate and the working electrode substrate are physically and electrically isolated from each other on the sensor. Physically and electrically isolated electrodes are advantageous in terms of the ability to pierce distinct areas of skin and reducing short circuits, and in some embodiments having very high MNA density, assists in facilitating skin piercing by application of finger pressure only. A finger pressure is considered as around 5 - 8 N, more preferably about 6 - 7 N. Actuators or the like to aid skin penetration are not required in preferred embodiments of the invention.

Preferred working electrodes for the sensors are defined in the first and second aspects of the invention above. Desirably, the microprojections are formed from metal, such as silicon or a polymer preferably a cured or cross linked polymer having a modulus of at least 1 GPa. Suitably, the cured or cross linked polymer is a cured hybrid polymer, such as a cured silica-polymer hybrid or cured polylactide. Preferably, the cured polymer is formed by curing in a mould generated from a microprojection array template, preferably a PDMS mould.

Desirably, the electronically conductive conformal nanofilm is in the form of (i) a metal film, (ii) a metal film comprising a surface metal and an adhesive undercoat of a metal, (iii) a porous carbonized surface film or (iv), a conducting polymer layer such as a polyaniline layer.

Preferably, the conformal nanofilm is in the form of a layer of at least one sputtered metal, more preferably in the form of a conformal nanofilm comprising a first surface metal and a second adhesive metal underlying the surface metal. The quality of the adhesive bond between the microneedle surface and the surface metal is important as it must be robust enough to support the sensing chemistry during use and particularly under the friction condition experienced when the needles pierce the skin as well as under the environmental conditions experienced when positioned in the skin for an extended period of time. Suitably the surface metal layer passes the sticky tape test where sticky tape is placed on a sample area of the surface and the surface metal layer is retained on the microneedle after the sticky tape is removed. The conformal nanofilm can also be in the form of a carbonized layer or region, for example, which can be generated from thermal carbonisation of the microneedle and/or support component by high temperature treatment in acetylene.

The conformal nature of the film is important as it enables maximum use of the high surface electrode/electrode sink regions enabled by the high density MPA arrangement. However, due to the high densities, the conformal film has to be very thin and yet robust enough and mechanically strong enough to support a sensing chemistry/system on the microneedle surface even under the challenging conditions encountered in piercing skin as well as exposure to particular chemical conditions encountered when left in place transdermally for a period of time.

In particular, it was unexpected that the sputtered metal nanofilm would be sufficient to meet these requirements. It is believed that the underlying adhesive thinner layer of metal plays a useful role in meeting these challenges.

While the description and examples described herein primarily relate to microneedles, micropillars or microprojections may also be used and in this respect, as used herein, the term ‘microneedle’ should be interpreted as including references to micropillars or other microprojections as long as they are capable of piercing the skin to access ISF and remain in an intact form.

Preferably, the sensor is adapted for continuous pH or bioanalyte monitoring or sensing. Most preferably, the sensor is adapted for continuous monitoring sensing of in vivo pH or bioanalyte monitoring, preferably continuous in vivo pH or bioanalyte monitoring of interstitial fluid residing in the skin, that is the dermis and epidermis. Desirably sensors may be used repeatedly, that is inserted into skin, and after measurement is made, if desired withdrawn from the skin for up to 10 time without losing more than 20% of the initial measurement capability or up to 5 times without losing more than 5% of the initial measurement capability such is the robustness of the device. Advantageously, the inventors have found a very high microneedle array skin pierceable MNA provided with a nanolayer or a nanofilm of a chemically inert but electrically conductive material as described for pH or bioanalyte sensing using a suitably robust sensing bioreceptor which facilitates continuous and stable bioreceptor activity upon interacting with a bioanalyte directly in interstitial fluid in a manner that involves continuous interfacing, monitoring and sensing.

The medium of interest serves as the electrolyte in the electrochemical reaction which occurs at the interface of the electrode and the electrolyte. Suitably, the medium of interest is a test solution, for example, for calibration of the sensor system, a biofluid such as blood (including peripheral blood), sweat or interstitial fluid, preferably in vivo interstitial fluid. In some embodiments, the electrolyte comprises a redox couple, for example, a ferro/ferricyanide redox couple, which is useful for sensing techniques relying on monitoring of changes in current intensities obtained from DPV measurements performed before and after binding of a target such as an antigen to an immobilized antibody on the working electrode surface. In such embodiments, if desired, the working electrode may be used for immunocapture of a target of interest in the skin ISF and withdrawal of same for subsequent electrochemical analysis. For example, antibody-antigen binding can be monitored by DPV in this way. Suitably antibodies include anti-HER2 and suitable antigens include erbB2 which is a breast cancer biomarker, whereby a level of above 15 ng/ml is indicative of breast cancer. Most preferably, the medium is in vivo interstitial fluid, that is, interstitial fluid that remains in the body. This is in direct contrast to sensing methods involving sensors which withdraw ISF from the body to be subjected in vitro or test external to the body. The medium of interest serves as the electrolyte in the electrochemical reaction which occurs at the interface of the electrode and the electrolyte. For example, when the microneedle electrodes of the sensor are embedded in the skin to access ISF, the ISF is the electrolyte completing the electrical circuit involving the cell.

The electrochemical sensor system involves an electrochemical cell whose output signal is directly related to the concentration of the bioanalyte species in the medium of interest. In short, the electrochemical sensor converts electrode transfer from the reaction between an electrode and analyte into a qualitative or quantitative electric signal. Electrochemical sensors are mainly divided into three types: potentiometric, conductometric, and amperometric/voltammetric. Desirably, the electrochemical sensor is adapted for amperometric detection of the bioanalyte, preferably, amperometric detection, more particularly chronoamperometric detection, most preferably, amperometric enzymatic detection. Potentiometric techniques involve measuring the potential difference between two electrons under no current conditions. Conductometric sensing is based on changes in the conductance of the cell. Other electroanalytic techniques may be carried out with the electrochemical sensor system. These include cyclic voltammetry, differential pulse voltammetry, square wave voltammetry and linear sweep voltammetry.

The sensor system may include other components such as a potentiostat, feedback control unit and amplification circuits, a detection unit, software interface, WIFI transmitters or receivers. It will be understood that a potentiostat is the electronic hardware which controls the three-electrode cell for electrochemical experiments. The potentiostat maintains the potential of the working electrode (WE) at a constant level with respect to the reference electrode (RE) by adjusting the current at a counter electrode (CE). The potentiostat also converts the current at the working electrode to the voltage by a transimpedance amplifier with a high gain. Preferably, at least the sensor system and a potentiostat are integrated in the wearable devices.

By ‘physically and electrically isolated from each other on the sensor’, it is meant that each of the reference electrode substrate, the counter electrode substrate and the working electrode substrate are independent of, and physically separated from, each other. In other words, each of the reference electrode substrate, the counter electrode substrate and the working electrode substrate are formed from completely independent substrate components. The reference, counter and working electrode substrates are separate and distinct components. Each electrode substrate is not provided in the form of a single or integral piece of substrate. This is in contrast to existing microneedle array- based sensors where the microneedles are provided on a single or integral piece of substrate, on which piece, discrete areas or regions of microneedles on the substrate as designated as working, reference, and counter electrode areas the substrate. The physical separation of the substrate electrically isolates each electrode from each other and may, if desired, give the sensor more flexibility depending on how the individual electrodes are positioned on a supporting boy or supporting frame to form a sensor patch for application to the skin. It is also believed that this arrangement aids transdermal application with mere finger pressure despite the high density MNAs utilised. It is considered that the three-electrode cell set up in the form of separate electrodes enables a sensing device which is improved in terms of wearability, options for addition of one or more working electrodes, facilitating detection of multiple analytes, as well as convenience for manufacturing and replacing parts.

Preferably, the electrochemical sensor system is adapted for pH sensing or glucose or lactate detection, monitoring and measurement. Preferably, the electrochemical sensor system is an electrochemical pH sensor or a glucose or lactate biosensor, although other the system can be adapted for other bioanalytes such as urea, ascorbic acid and ethanol. For example, the system can be adapted for enzymatic detection in general whereby enzymes used include glucose oxidase (GOx) for glucose sensing, lactate oxidase (LOx) for lactate sensing, urease for urea sensing and alcohol oxidase (AOx) for ethanol sensing. Desirably, the enzyme specific for the bioanalyte of interest is GOx and the bioanalyte is glucose. Desirably, the enzyme specific for the bioanalyte of interest is LOx and the bioanalyte is lactate.

Preferably, the microneedles of each electrodes project upwards in a perpendicular alignment to a horizontal plane formed by the electrode substrate.

Desirably, in some embodiments, the microneedles on one or more of the reference electrode substrate, the counter electrode substrate and working electrode substrate are solid, meaning they are not hollow and comprise a substantially smooth, defect or pore free surface. Preferably, the solid microneedles are not hollow. In other embodiments, while the microneedles are solid non hollow microneedles, they may have a porous surface layer or region on the surface of the microneedles. For example, the pores may be provided in a microneedle, particularly an Si microneedle, by anodisation or electrochemical etching. The carbon layer can be provided on top thereof. Desirably, such a porous layer is a thin layer, for example, being from about 500 nm to about 10 microns thick, preferably from about 1 to about 7.5 microns thick. Suitably, the porous microneedles have pores with a pore diameter of < 100 nm, preferably about 10 to 90 nm, most preferably about 50 nm to 60nm. Desirably, the porosity can range from about 20% to about 80%, preferably from about 30% to about 60%, preferably from about 40% to about 50%.

Suitably, the microneedles have a base diameter in the range of about 10 microns to about 150 microns, preferably about 50 microns. Desirably, the microneedles are spaced apart from each other by a distance of from about 50 to about 200 microns from centre to centre, preferably about 50 - 120 microns, most preferably about 110 microns.

Preferably, the microneedle density on one or more of the reference electrode substrates, the counter electrode substrate and working electrode substrate is at least 2,000 microneedle/cm ² , at least 7,000 microneedle/cm ² , at least 8,000 microneedle/cm ² , at least 9,000 microneedle/cm ² , at least 10,000 microneedle/cm ² , at least 15,000 microneedle/cm ² . In one Suitably, in some embodiments, the microneedle density is no greater than 20,000 or 25,000 microneedle/cm ² . Desirably, the microneedles are present at a density of at least 7,000 microneedle/cm ² . In one preferred embodiment, the density is about 9,500 microneedle/cm ² . Suitably, the maximum microneedle density is about 25,000 microneedles/cm ² , more preferably about 20,000 microneedles/cm ² . In some embodiments, where greater bioanalyte sensitivity and/or lower limits of bioanalyte detection are required, higher microneedles densities are preferred.

The array can be any regular array shape, for example, a hexagonal array or a square array.

Desirably, the microneedles have sharp tips having a tip diameter in the range of from 1 to 10 microns, preferably about 2 microns. Suitably, the microneedles have an aspect ratio of from about 2 to about 10 microns, preferably 3 microns. Preferably, the microneedles have a length in the range of from about 50 microns to about 500 microns, preferably about 250 microns.

In a particularly preferred embodiment, the MNA has needles, preferably conical shaped needles of about 250 microns height, a base diameter of about 50 microns, and needle tip diameters of about 2 microns. The sharpness is ideal for painless skin penetration whereas the shape, size and density are all suitable for analysis of epidermis and dermis layers without puncturing or damaging blood vessel of nerve bundles.

In preferred embodiments, the array has a microneedles (MN) with a density of from 7500 to 12500 MNs/cm ² and a MN height of 200 to 300 pm, a base diameter of 25 to 75 pm, and space between two adjacent cone centres of 75 - 125 pm. In one particularly preferred embodiment the array has a density of -9500 MNs/cm ² and a MN height of -250 pm, a base diameter of -50 pm, and space between two adjacent cone centres of -110 pm. Suitably, the array has a hexagonal arrangement of conically shaped microneedles or pillars with sharp defined needle tips.

Desirably, the microneedles of one or more of the electrodes comprise a biocompatible material. Biocompatible materials are preferred for wearable continuous monitoring sensors which remain on the body for extended periods of time. Preferably, one or more of the reference electrode substrates, the counter electrode substrate and working electrode substrate comprises a biocompatible material. Biocompatible materials include a metal such as silicon, gold, platinum, titanium; a polymer such as polycarbonate, polylactic acid, polylactic-co-glycolic acid, polymethyl methacrylate as polymers; and UV curable polymers such as an inorganic-organic hybrid polymer for example, Ormocomp, or an epoxy based polymer such as SU-8. Suitably, wherein the biocompatible material of the microneedles and the microneedle base are the same, preferably the microneedles and the microneedle base are in the form of an integral piece. Such a piece may arise where the microneedles are formed by etching of a precursor block of etchable material, for example silicon. Desirably, the microneedle base component of each electrode is not treated with a non specific binding coating such as a polymer or the like. Unexpectedly it was found that leaving the substrate base non-specific binding coating free (without such a coating) gave improved results this was contrary to the expectation as it was believed that interference or sensitivity loss would result where the base was not treated to reduce non-specific binding interactions. It should be noted that the substrate base is also provided with the electronically conductive material, that is sputtered metal(s) or thermal carbonized layer.

Preferably, at least the microneedles of one or more of the reference electrode substrate, the counter electrode substrate and working electrode substrate are coated with or otherwise provided with at least one thin film of an electrically conductive but chemically inert material, preferably a conformal film, more preferably a conformal nanofilm. As described above, in some embodiments, this metal layer can be a carbonised layer, a conducting polymer layer or a metal sputtered layer or a metal sputtered layer undercoated with a thinner layer of a second different metal, as an adhesive layer, such as a chromium undercoat/underlayer. In other embodiments, the nanofilm is provided a layer of carbonised, preferably thermally carbonised layer or region on the micropillars. Preferred conducting polymers include polyaniline, preferably electropolymerized onto the substrate and needles. The electrically conductive material may act as an electrode sink at the electrode surface where the electrochemical reactions occur. Preferably, the conformal and thin nanofilm or nanolayer has a thickness of from about 5 to about 500 nm, preferably about 150 nm. Suitably, the entirety of each electrode that is microneedles and substrate base is coated with the at least one film of the material. Desirably, the at least one film is a sputtered metal film, e.g., formed by sputtering or an electrodeposited metal film. Preferably, the inert material, is a biocompatible material, e.g., silicon, zinc oxide, carbon, gold or platinum, most preferably gold or carbon. In the case of solid, non-hollow microneedles, the biocompatible material is preferably a gold or platinum conformal nanofilm, most preferably electro-sputtered gold or platinum, preferably of about 150 nm thickness. In some preferred embodiments using a metal nanofilm, a further layer of a compatible material, e.g., another electrically conductive material such as titanium, nickel or chromium, is positioned between and the inert material the needles and substrate. The additional layer of compatible material is one which may improve the adhesion of the outer electrically conductive but chemically inert material to the needles and substrate. This layer may useful for facilitating the formation of the conformal outer/surface layer and for allowing it to support a sensing chemistry that is subjected to skin piercing friction on transdermal introduction as well as withstanding exposure to in vivo conditions when being worn for an extended period of time. For example, where the inert material is a film of gold, a layer of a compatible, but different metal, may be provided between the microneedles and the gold, for example, resulting in improved adhesion of the gold to the microneedles. Such additional layer of another material is a thinner layer than the electrically conductive but chemically inert material that forms of the surface of the electrodes. For example, in one embodiment, a layer of about 15 nm chromium assists in adhesion of a 150 nm nanofilm of gold. This arrangement has been found to be particularly useful to support a conformal layer of sensing enzyme.

In the case of solid, non-hollow microneedles which comprise a porous surface layer or region, the biocompatible material is preferably carbon or a carbonized surface, desirably in the form of a conformal nanolayer of carbon, preferably thermally carbonised carbon. A carbon layer is particularly preferred for microneedles having pore diameters of 100 nm, as sputtered coats cannot be used to deposit conformal layers on such nanostructures to increase the microneedle conductivity. Suitably, a conformal, nanometric carbon layer or thermal carbonised layer may be used. Such a carbon layer may be provided by thermal treatment of the microarray needles at high temperature (400-800 °C) in the presence of acetylene. A conformal, nanometric carbon layer was particularly found to provides silicon microneedles with surface porous region with enhanced electrochemical properties and outstanding chemical stability.

Desirably, the microneedles of the reference (RE) electrode are further coated or modified with a suitable reference electrode chemistry or material which has a stable and well-known electrode potential. Suitable reference materials are well-known in the art and include iridium oxide and graphite/AgCI, for example. Preferably, the reference electrode material forms an Ag/AgCI reference electrode material. During manufacture, the AgCI is preferably applied to the microneedles of the reference electrode in the form of drop cast AgCI ink which is subsequently dried to form the AgCI coating. Preferably the reference electrode material is provided as a conformal layer.

Preferably, at least the microneedles of the working (WE) electrode are coated, modified or otherwise further functionalized to detect and monitor the bioanalyte of interest. For example, surfaces of the working (WE) electrode can be provided with a sensing chemistry (for example, polyaniline for pH sensing) or a biorecognition element (enzyme, antibody or antigen) which can be immobilized onto the working electrode. Insofar as possible, the chemistry or biorecognition element is selective and specific for the bioanalyte of interest in the medium of interest. For example, in some embodiments the biorecognition element is associated with a biotransducer. A biotransducer is a recognition-transduction component consisting of at least two intimately coupled parts, a biorecognition layer and an electrochemical physicochemical transducer, which acting together, converts a biochemical signal to an electric signal. The biorecognition layer typically contains an enzyme or another binding protein such as an antibody. However, oligonucleotide sequences, sub- cellular fragments such as organelles (e.g. mitochondria) and receptor carrying fragments (e.g. cell wall), single whole cells, small numbers of cells on synthetic scaffolds may also comprise the biorecognition layer. The physicochemical transducer is typically in intimate and controlled contact with the recognition layer. As a result of the presence and biochemical action of the bioanalyte, a physicochemical change is produced within the biorecognition layer that is measured by the physicochemical transducer producing a signal that is proportionate to the concentration of the analyte.

Most preferably, the biotransducer is an amperometric transducer capable of detecting a change in current as a result of electrochemical oxidation or reduction. Typically, the bioreceptor molecule is immobilized on the working electrode. The potential between the working electrode and the reference electrode (e.g., Ag/AgCI) is fixed at a value and then current is measured with respect to time. The applied potential is the driving force for the electron transfer reaction. The current produced is a direct measure of the rate of electron transfer. The current reflects the reaction occurring between the bioreceptor molecule and bioanalyte and is limited by the mass transport rate of the analyte to the electrode.

Preferably, the sensing chemistry or the biotransducer includes a redox mediator if required. In one embodiment, the biotransducer may include an enzyme specific and selective for the bioanalyte of interest. For example, the enzyme may be glucose oxidase (GOx) which is specific for glucose or lactate oxidase (LOx) which is selective for lactate. In this case, the sensor in the three- electrode cell has an immobilized enzyme on the electrode for selective detection of glucose or lactate in the medium of interest, e.g., ISF. The current that is generated from the sensor electrode (working electrode: WE) by the enzyme reaction varies linearly with the concentration of glucose/lactate in the ISF, according to the Michaelis-Menten equation. Alternatively, the enzymes used may further include urease for urea sensing or alcohol oxidase (AOx) for ethanol sensing.

In some embodiments, a redox mediator is included as part of the sensing chemistry or biotransducer. The redox mediator can shuttle electrons to the electrode surface which can be useful where the enzyme is bulky and can shield the electrode surface from electrons released in the redox reaction. In the field of amperometric oxidoreductase electrodes, immobilized redox mediators are incorporated into the system to shuttle electrons between the enzyme’s redox centres and the electrode. A preferred redox mediator include ferrocene which undergoes a reversible one- electron oxidation at around 0.5 V versus a saturated calomel electrode (SCE). Ferrocene is a particularly successful redox mediator due to its well-behaved electrochemical properties. Other redox mediators include Prussian blue and methylene blue. Suitably, the redox mediator is retained in position between the surface and the enzyme by the dendrimer as this configuration has been found to lower the working potential and therefore reduce non-selective binding of non-target potential interfering biomolecules present.

Most preferably, the redox mediator is provided in the form of a redox mediator modified dendrimer, for example, a ferrocene-dendrimer derivative or a redox mediator modified conducting nanoparticle/polymer composite such as a Prussian Blue-chitosan-Au nanoparticle composite. The redox mediator may be provided as the focal point of the dendrimer. The dendrimer can be immobilised onto the electrically conducting surface thereby positioning the redox mediator close to the surface whereby the enzyme may be immobilised to the dendrimer in a single layer. Having the redox mediator as the focal point fixes and retains the redox mediator between the electrode surface and the enzyme layer. The dendrimer may be any generations from GO upwards, however, G2 and G3 are preferred, G2 is particularly preferred. Suitable dendrimers include hyperbranched polyglycerol dendrimers, polyethylene glycol dendrimers, glutamic acid dendrimers and PAMAM dendrimers. PAMAM dendrimers are particularly preferred as they have been extensively studied in terms of toxicology profile in the context of drug delivery.

The redox mediator modified conducting nanoparticle/polymer composite may be in the form of a redox mediator layer on the working electrode surface which a polymer, preferably a biopolymer such as chitosan, and electronically conducting nanoparticles such as AuNPs are drop casted onto the electrode surface. In such embodiments, use of polymer MNA is particularly preferred. For example, a pMNA/Au/PB-ChAuNPs/LOx biosensor may be prepared first by electrodeposition of Prussian Blue (PB) layer on a pMNA/Au electrode after which chitosan and gold nanoparticles (AuNPs) are drop casted onto pMNA/Au/PB electrode. Thereafter LOx was immobilized onto pMNA/Au/PB-ChAuNPs surface using a crosslinking agent such as glutaraldehyde. In other preferred embodiments using pMNA, the sensing layer may comprise pMNA/Au/Fc-PAMAM/LOx using the approach described elsewhere in herein for glucose sensor preparation except that instead of silicon MNA pMNA were used and instead of GOx enzyme Lactate Oxidase (LOx) is used.

For porous microneedles which are provided with a conformal nanometric carbon layer, a preferred sensing system involves initial surface functionalization of the carbonized/carbon layer, e.g., by electrografting a diazonium salt directly onto the conducting surface. A desired sensing molecule or biorecognition elements or bioreceptor can then be bound to the diazonium salt. For example, dendrimer or redox mediated modified dendrimer, such as Fc-PAMAM, may be covalently bound to the diazonium salt, e.g. by carboxyl group activation. A sensing enzyme, e.g., GOx can then be bound to the dendrimer or redox mediated modified dendrimer, e.g, by crosslinking with a suitable crosslinking agent such a glutaraldehyde (GAS). In this case immobilisation of the dendrimer onto the electrically conducting surface positions the redox mediator close to the surface whereby the enzyme may be immobilised to the dendrimer in a single layer

Suitable amperometric glucose biosensors can be fabricated by immobilization of bioreceptors, e.g., glucose oxidase (GOx), onto a redox mediated dendrimer (for example, a PAMAM-Fc dendrimer) decorated gold surface modified solid MNA electrode or a carbon coated (e.g. thermally carbonised) porous MNA electrode. Suitably, these redox active asymmetric PAMAM dendrimers have ferrocene unit as a focal point of the dendrimer.

Preferably, one or more of the reference electrode substrate, the counter electrode substrate and working electrode substrate are provided on an electrode and connector support. Desirably, the electrode and connector support is adapted to isolate an electrically conductive region of the electrode from other regions of the electrode which are non-sensing, for example, parts or regions for connection to connector, electrical contact or lead. Desirably, the support can be a flexible support that can bend and adopt a conformation that is, for example, complementary to the contours of the skin. Suitably, the electrode and connector support may be a 3D printed support using a suitable printable material. Photocurable polymers such as acrylic- and methacrylic-based, epoxy- based, or vinyl ethers-based resins can be used. Suitably, a photopolymer such as the acrylic based FullCure 720 may be used.

Desirably, each electrode substrate is affixed to, or mounted onto, its own individual support, for example, by curing using a suitable adhesive, particularly any glue-like cyanoacrylates (superglue) or curable silicones. In one embodiment, each of the electrode supports are physically distinct or independent from each other. This ensures that the electrodes are electrically isolated from each other in the sensor system.

In one embodiment, one or more of the electrode and connector supports may comprise a distal portion adapted for connection to one or more connectors, contacts or leads for transfer of electrical signal, and a proximal portion comprising the electrode substrate of the electrode. If desired, the electrode and connector support can further comprise a bridging portion disposed between the distal portion and the proximal portion adapted to transfer an electrical current from the microneedles to the distal portion. In one embodiment, the bridging portion is an elongate bridging portion whereby the support has an elongate body form whereby the electrode substrate component is at the head of the elongate body and the connectors for a connector, electrical contact or lead form of the tail of the elongate body. The elongate body embodiment is desirable where flexibility of the support is important.

Suitably, portions of the electrodes and/or electrode and connector supports which are non sensing may be coated with an electrically insulating material, for example, an electronic conduction insulator and/or an ionic conduction insulator material, such as a dielectric ink. This material is desirable as it reduces signal noise originating from non-sensing parts of the electrode. As it is the same conformal nanofilm material used on the sending MNA, it is believed to provide a more electronically robust, sensitive and flexible device without additional wires or soldered parts that may be more prone to failure or corrosion. In some embodiments, the connectors, contact or leads provided on the distal portion of the electrode support are not coated with the insulating material.

It will be understood that each electrode and connector support may be mounted on a base in the form of a fixing frame or a fixing board for mounting, positioning and fixing the individual electrode and connector supports in a desired configuration, e.g., linear, staggered, or a radial arrangement of the electrodes, as desired. The fixing board or frame can be made from a solid and ridged thermoplastic, e.g., PMMA, Polyimide, polycarbonate, or crosslinked elastomer such as silicon rubber or a printable, photocurable, rigid, semi translucent material which is suitable for 3D printing, such as FullCure 720.

The fixing frame or a fixing board, particularly where flexible, enables convenient application of the sensor system to the skin, e.g., in the form of a patch. The fixing frame or a fixing board also allows a distributed and even skin piercing pressure to conveniently be applied such that substantially all the microneedles pierce the skin to substantially the same degree. Desirably, this positions the needles optimally to ensure maximum contact with the medium of interest, i.e., ISF for transdermal monitoring. Where the supports are mounted onto a base the base can be of any desired size. The base functions to hold the three electrodes in close proximity while retaining physical separation of each of the three electrodes. The base ensures that the electrodes are electrically isolated from each other in the sensor system. Additionally, the base is considered important for wearability, ease of replacement of parts, and multiplexing where two or more analytes can be detected simultaneously or in the same device, for example.

Preferably, the sensor system is provided in the form of a sensing patch, preferably a wearable sensing patch. Desirably, the system is configured as a transdermal wearable bioanalyte sensor, e.g., and can include various retaining means for holding the sensor in position in the skin. The retaining means can include adhesive pads or films, straps, cuffs, etc.

Described herein is a wearable device for bioanalyte monitoring comprising the sensor system of the first aspect. Desirably, the device is adapted for continuous monitoring. In the continuous monitoring embodiment, the sensor system is retained in place on the skin for a period of up to about 14 days. This embodiment enables real time monitoring of the bioanalyte. In the case of at least glucose this is desirable because while the concentration of glucose in ISF is identical to that in blood in steady-state conditions, there may be a lag time immediately after glucose intake due to the kinetics of diffusion. This lag time is estimated to be about 5 to about 15 minutes.

Described herein is a use of an electrochemical sensor system as defined in the first aspect in the transdermal monitoring of a bioanalyte, preferably glucose, lactate, ethanol or urea.

Desirably, the transdermal monitoring is continuous inline, in vivo monitoring, which does not require sampling or withdrawal of any bodily fluid during testing.

Also described herein is a method for manufacturing an electrochemical sensor for sensing a bioanalyte in a medium of interest, the method comprising the steps of: forming a microneedle array of an electrically conductive biocompatible material on a reference (RE) electrode substrate at a density of at least 2,000 microneedle/cm ² , coating the microneedle array and the reference electrode substrate with a conformal nanofilm of at least one electrically conducting, chemically inert material to form a reference (R) electrode; forming a microneedle array of electrically conductive biocompatible material on a counter (CE) electrode substrate at a density of at least 2,000 microneedle/cm ² , coating the microneedle array and the counter electrode substrate with a conformal nanofilm of at least one electrically conducting, chemically inert material to form a counter (C) electrode; forming a microneedle array of electrically conductive biocompatible material on a working (WE) electrode substrate at a density of at least 2,000 microneedle/cm ² , coating the microneedle array and the working electrode substrate with a conformal nanofilm of at least one electrically conducting, chemically inert material to form a working (W) electrode; assembling the reference electrode substrate, the counter electrode substrate and the working electrode substrate in a sensing configuration, wherein the microneedles arrays on each of the reference electrode substrate, the counter electrode substrate and the working electrode substrate are configured to be contactable with the medium of interest which serves as electrolyte in the cell, and wherein the microneedles arrays on each of the reference electrode substrate, the counter electrode substrate and the working electrode substrate are physically and electrically isolated from each other on the sensor. Desirably, the MNA density is at least 7,000 microneedle/cm ² , more preferably about 9,500 microneedle/cm ² .

Suitably, the conformal film is a conformal nanofilm of at least one electrically conducting, chemically inert material is as described above, for example, a thin nanofilm of gold, particularly about 150 nm gold, preferably underlaid by an adhesive thinner layer of chromium. Most preferably, the conformal nanofilm of inert material is preferably a sputtered film. In another embodiment, the conformal, nanolayer is a carbon layer which is particularly suited to microneedles having a porous surface layer, particularly silicon porous microneedles. A preferred carbon layer is a thermally carbonized layer or region having pores therein, for example pores of <100 nm. Preferably, prior to assembly of the sensor system, the method further comprises the step of isolating the reference (RE) electrode from the counter (C) electrode and the working (W) electrode and coating at least the microneedles of the reference (RE) electrode with a suitable counter (c) electrode material. For example, the reference (RE) electrode can be modified with a film of AgCI to form an Ag/AgCI counter electrode. Preferably, the method further comprises isolating the working (W) electrode from the reference (R) electrode and the working (W) electrode and functionalizing at least the microneedles of the working (W) electrode with a sensing chemistry or a biotransducer which is specific and selective for the bioanalyte of interest in a medium of interest. Preferably, the method further comprises associating each of the reference electrode substrate, the counter electrode substrate and the working electrode substrate with its own electrode substrate support. As described above, the support functions to provide structure of the sensor while maintaining physical separation of each of the three electrodes, thereby electrically isolating each electrically conductive region of the electrode from another region of the support which is adapted to be connected to a connector or electrical contact or lead.

In the method described above, the electrode substrate may be provided onto the corresponding electrode support prior to coating with the electrically conducting chemically resistance material. Alternatively, the coated microneedles and/or substrate may be provided onto the support which must also be provided with electrically conducting path to facilitate electrical conduction from the electrode to the connectors, contact or leads provided on the distal portion of the electrode support.

The method further comprises providing portions of the electrode support which are non sensing are coated with an electrically insulating material, for example, an electric conduction insulator and/or an ionic conduction insulator material.

Suitably, the method further comprises the step of assembling the reference (RE) electrode, the counter (CE) electrode and the working (WE) electrode onto a fixing frame or a fixing board which fixes and/or spaces the individual electrodes in a desired position or spatial relationship to each other thereby forming a sensing patch suitable for transdermal application. The support body can be of any desired shape or size suitable enabling the electrodes to be provided in the form of, for example, a patch of any desired size. The support body may be rigid, or it may be flexible which can increase comfort in the case of a wearable device. Suitably, the method further comprises the step of 3D printing one or more of the electrode substrate supports. Desirably, the method further comprises the step of fixing the electrode substrate onto the substrate support, by, for example, curing using a suitable adhesive. Suitably, in the method, the step of forming sensing microneedles involves forming the sensing microneedles on two or more of the reference electrode substrate, counter electrode substrate and working electrode substrate at the same time, preferably wherein microneedles are formed on all of the reference, counter and working electrode substrates at the same time. Desirably, the step of coating the microneedles with the film of inert material involves coating the microneedles of one or more of the reference, counter and working electrodes are coated with the film of inert material at the same time, preferably all of the reference, counter and working electrode substrates are coated with the film of inert material at the same time, e.g., by spray coating. Preferably, the step of coating the non-sensing parts of the electrode with the electrical insulator involves coating the sensing microneedles on two or more of the reference, counter and working electrodes at the same time, preferably coating the sensing microneedles on all of the reference, counter and working electrodes at the same time. By ‘the same time’, it is meant simultaneously. Simultaneous coating results in manufacturing advantages over coating one at a time, which can be done but is less preferred than simultaneous coating. Desirably, the biocompatible material is silicon or a biocompatible cured polymer.

Preferably, the method comprises forming the microneedle array by etching, preferably ion etching, most preferably deep reactive ion etching, for example, of monocrystalline <100> Si wafers. Suitably, the etching involves the step of creating a sharp tip on each microneedle, for example, by isotropic SF ₆ etching. Desirably, the method further comprises the step further etching to provide microneedles of a predetermined length. Desirably, the method further comprises a tip reshaping step, whereby the tip of each microneedle is subjected to a reshaping step to sharpen the tip of the microneedle, preferably, reshaping is by a pseudo Bosch process. Where porous microneedles are desired, pores, e.g., < 100 nm diameter, can be etched into the microneedles, e.g. by electrochemical etching.

Preferably, one or more of the reference electrode substrate support, the counter electrode substrate support, and the working electrode substrate support formed by 3D-printing. Suitably, the method further comprises the step of adhering the substrate onto the support by adhesive. Preferably, where the inert material is gold, the method involved the step coating the inert material in a conformal film by a sputtering technique. Desirably, during sputtering the electrode to be coated is rotated which facilitates formation of a conformal film of the sputtered metal. In one example, where the biocompatible material is gold for example, the gold conformal nanofilm may be deposited onto the electrode by initially pre-treating the microneedles with a sputtered layer of compatible material that may assist in adhesion of the biocompatible material. The layer of compatible material may range in thickness of from about 5 to about 50 nm, preferably about 15 nm. Suitably, the electrode to be sputtered with metal may be rotated during sputtering to form a suitable conformal layer around the microneedles. Preferably, during sputtering rotation speed of the electrode is kept at constant at a rotation speed of from about 25 to about 200 rpm, preferably about 100 rpm. Desirably, during sputtering the sputter chamber vacuum is kept constant at about 10 ^_1 to about 10 ^{_ 7} mbar, preferably about 2.6 x 10 ^-3 mbar pressure. A particularly preferred layer of compatible material is a chromium nanolayer, at least in the case of a gold outer coat. Suitably, the sputter chamber vacuum and the sample rotation were kept at constant 2.6 x 10 ^-3 mbar pressure and 100 rpm, respectively. A chromium layer of about 15 nm is particularly preferred. Such a chromium layer can be applied via dual AC with power of 300 W by sputtering while rotating the electrode to be coated at a rotation speed of 100 rpm. The sensing chemistry is then applied to the working (WE) electrode.

Also described herein is a sensor system obtainable by the method of manufacture of the invention as described herein.

Described herein is a microneedle array sensor system for transdermal pH and/or bioanalyte monitoring, preferably, biomarker capture and sensing, ascorbic acid sending, glucose monitoring, lactate sensing, urea sensing or ethanol sensing.

In a bioanalyte sensor system, once the bioanalyte of interest reaches the working (sensing) electrode, an electrochemical reaction occurs involving either reduction or oxidation depending on the bioanalyte in question. Oxidation causes the flow of electrons from the working electrode to the counter electrode through the external circuit, while reduction causes the flow of electrodes to move from the counter electrode to the working electrode. In either case the direction of electron flow creates an electric current proportional to the concentration of the bioanalyte present in the medium of interest. The external circuit detect and amplify this current and scales the output accordingly with the calibration to give a reading in desired units, such as millimoles for example.

By ‘chemical redox reaction’, it is meant a process which is caused by, or causes, passage of an electrical current, typically involving a substance or material participating in the chemical redox reaction, that is a chemical reaction resulting in loss or acceptance of an electron to or from one or more chemical species.

By ‘electrochemical reaction’, during the redox reaction, reduction occurs at one electrode and oxidation occurs at another electrode, in other words the two half reactions are spatially separated. This is in contrast a chemical redox reaction where the reduction and oxidation occur at the same place. An electrochemical reaction is accompanied by the transfer of electrons (i.e. an electric current) or caused by the transfer of electrons (i.e., an electric current).

By a ‘chemically inert’ material, it is meant a material that does not undergo redox process under the operating potential applied to the sensor during normal use.

The electrochemical cell is where a chemical reaction occurs by passing electricity through the electrodes. For a better control and measurement of the current and potential going through the cell during the electricity driven chemical reaction, a three-electrode system is preferably as this reduces and compensates the potential changes caused by large currents passing through the working and counter electrodes. The chemical reactions happen at the working electrode and the counter electrode which serves as a current source, while the reference electrode acts as a reference in measuring and controlling the working electrode potential. The reference electrode has a known and stable electropotential and therefore can be used as a half-cell to determine the electropotential of the other half of the cell (the working electrode). The most commonly used reference electrode is made of Ag/AgCI with defined potentials ranging from +0.20 to +0.25 V. For metabolite sensing, biosensors with enzyme recognition elements are used where the enzyme is tethered to the working electrode by a process of enzyme immobilization. The enzyme works in the system by catalyzing a reduction-oxidation (redox) reaction to initiate an electron transfer process between the redox centre of the enzyme and the working electrode. The product concentration increase caused by the enzymatic reaction can be read using amperometry if the substances partake in redox reactions.

Amperometry is the term indicating the whole of electrochemical techniques in which a current is measured as a function of an independent variable that is, typically, time or electrode potential. Chronoamperometry is the technique in which the current is measured, at a fixed potential, at different times since the start of polarization. Chronoamperometry is typically carried out in unstirred solution and at fixed electrode, i.e., under experimental conditions avoiding convection as the mass transfer to the electrode. On the other hand, voltammetry is a subclass of amperometry, in which the current is measured by varying the potential applied to the electrode.

Described herein is a use of a microneedle array transdermal sensor system according to the invention in a method of transdermal electrochemical bioanalyte monitoring.

Described herein is a method of preparing a calibration curve for a bioanalyte in a medium of interest using the electrochemical sensor system of the invention, the method comprising the steps of: providing the electrochemical sensor system to at least one solution of a test medium of interest such that electrode surfaces of the microneedles contact the medium of interest, wherein the solution of test medium of interest comprise a known concentration of the bioanalyte; applying a steady state electrical potential to the sensor’s counter electrode, whereby the potential provides a suitable potential window for detection of a redox reaction or chain of reactions initiated by the presence of the bioanalyte; monitoring for changes in an electrochemical property occurring at the sensor’s working (WE) electrode which result from a redox reaction involving the bioanalyte of interest; associating an observed electrochemical property change with the known bioanalyte concentration; and optionally repeating the preceding steps with solutions of a test medium of interest comprising different known amount of the bioanalyte.

The step may further comprise the step of fitting the observed current data and the concentration date to a curve.

Described herein is a method of bioanalyte monitoring in a medium of interest comprising the steps of: providing an electrochemical sensor system according to the invention to a test medium of interest such that electrode surfaces of the microneedles contact the medium of interest; applying a steady state electrical potential to the sensor’s counter electrode, whereby the potential provides a suitable potential window for detection of a redox reaction or chain of reactions initiated by the presence of the bioanalyte; monitoring for changes in an electrochemical property occurring at the sensor’s working (WE) electrode which result from a redox reaction involving the bioanalyte of interest; establishing the concentration of the bioanalyte present in the medium by comparing a level of a detected electrochemical property at the sensor’s working (WE) electrode with a predetermined calibration curve for the bioanalyte.

Described herein is a transdermal bioanalyte monitoring comprising the steps of: providing an electrochemical sensor system according to the invention to a test subject’s skin; aligning the sensor’s microneedle tips against the skin; applying a skin piercing pressure to the sensor to position the microneedles in skin’s dermal layer thereby allowing electrode surfaces of the microneedles to access the subject’s interstitial fluid; applying a steady state electrical potential to the sensor’s counter electrode, whereby the potential provides a suitable potential window for detection of a redox reaction or chain of reactions initiated by the presence of the bioanalyte; monitoring for changes in an electrochemical property occurring at the sensor’s working (WE) electrode which result from a redox reaction involving the bioanalyte of interest; establishing the concentration of the bioanalyte present in the interstitial fluid by comparing a detected level of the electrochemical property at the sensor’s working (WE) electrode with a predetermined calibration curve for the bioanalyte. Desirably, the electrochemical property is one or more of current, potential, resistance, impedance and capacitance.

Described herein is a method of diagnosis or monitoring of a disease or condition comprising the step of identifying the presence of a bioanalyte and/or a level of a bioanalyte associated with a disease or condition, using the sensor system according to the first aspect or fifth aspect. Desirably, the bioanalyte is glucose and the disease or condition is metabolic disease, insulin resistance, glucose intolerance, diabetes including Type 1 or Type 2 diabetes.

Described herein is an array of solid, non-hollow silicon microneedles at an array density of at least 2,000 microneedle/cm ² , wherein the array of microneedles is coated with a conformal nanofilm of at least one biocompatible electrically conducting, chemically inert material which supports a redox reaction, and, wherein the conformal nanofilm of gold is modified with a redox mediated dendrimer enzyme specific for a bioanalyte of interest.

Described herein is an array of solid, non-hollow silicon microneedles at an array density of at least 2,000 microneedle/cm ² , wherein the array of microneedles is coated with a first adhesive layer of chromium and a second conformal nanofilm of gold over the chromium layer, wherein the conformal nanofilm of gold is modified with a redox mediated dendrimer enzyme specific for a bioanalyte of interest.

Described herein is an array of solid, non-hollow silicon microneedles having a porous surface layer thereon at an array density of at least 2,000 microneedle/cm ² , wherein the array of microneedles is coated with a conformal nanofilm of carbon which is modified with a redox mediated dendrimer enzyme specific for a bioanalyte of interest.

Detailed description of preferred embodiments The present invention is described with reference to the following examples. It is to be understood that the examples are illustrative of and not limiting to the invention described herein. The exemplary, non-limiting embodiment described herein is of a sensor system in the form of a sensing patch for application to the skin of a subject for transdermal glucose monitoring. The sensing patch of this example comprises three distinct electrodes, namely a working electrode, a counter electrode and a reference electrode. Each electrode is a standalone electrode, meaning it is physically and spatially isolated from the other electrodes to at least some degree. In this example, each electrode is made up of a sensing part comprising a silicon microneedle array provided on a silicon substrate. Each substrate adheres to its own 3D printed holder/support to form each standalone electrode. Each of three standalone electrodes (microneedles and silicone base part of substrate) are coated with a film of electronically conductive material and non-sensing/non substrate parts of each electrode are coated with an electrically insulating material. No non-specific binding coating polymer of otherwise is provided on any parts of the microneedles or the base from which the needles project. The sensing microneedle array of the reference electrode is further coated with AgCI. The sensing microneedle array of the working electrode functionalized with a sensing chemistry or biochemistry, in this case, a redox mediated sensing chemistry adapted for specifically for glucose sensing. In this example, the redox mediated sensing chemistry is a ferrocene cored dendrimer, the surfaces of which are modified with glucose oxidase. The dendrimer is bound to the electrode surface, for example, through generation of an activated electrode surface for example functionalised with carboxyl acid terminated surface to which the dendrimer can be bound, e.g., through NHS/EDC coupling. For example, a carboxyl acid terminated surface can be generated through the formation of a self assembled monolayer (SAM) of 3-mercaptopropionic acid. On binding to the surface, the ferrocene redox mediator is positioned between the surface and the enzyme which is bound to the dendrimer in a single layer. The enzyme can be bound to the dendrimer through glutaraldehyde as a crosslinking molecular. The standalone nature of the electrodes gives rise to less interference during use and thus good sensitivity, as well as several manufacturing advantages in terms of minimal fabrication steps and associated time savings and reduced costs, ease of replacement of parts, etc. As the high density MNA electrodes are separate and distinct from each other, this means they can pierce skin with finger force much more easily than were the three electrodes provided onto a single substrate. The sensing patch contains high-density Si MN and is applied in one embodiment to the transdermal electrochemical glucose monitoring. Glucose sensing patch preparation process includes the fabrication of Si MNA using deep reactive ion etching after which the MNA was used to prepare reference, working, and counter electrodes on 3D-printed holding bases. The surface of the MN was initially coated with (i) a sputtered conformal Au/Cr thin layer for solid MNAs or (ii) a nanoscale carbon layer having pores electrochemically etched therein for porous MNAs. The surface layer of the electrode is then modified with mediator-cored dendrimers and enzyme, in this example, glucose oxidase (GOx) enzyme. The skin penetration test demonstrates the capability of the presented MN patch to fully penetrate the skin despite the high density MNA used in the device. Furthermore, in vivo application of glucose sensing patch using mice, as proof of concept, demonstrated the ability of the biosensor to monitor ISF glucose. The results obtained from monitoring of ISF glucose levels in mice showed a very good correlation with the blood glucose values recorded with a commercial glucometer. Also described is an array of solid, non-hollow silicon microneedles at an array density of at least 2,000 microneedle/cm ² , wherein each microneedle of the array of microneedles is coated with a conformal nanofilm of thermally carbonised carbon having a porous surface or region having pore diameters of <100 nn, more preferably diameters from 50 to 60 nm. Preferred needled densities and other components and characteristics are described above in detail.

Experimental section - Materials and instrumentations - Glucose oxidase from Aspergillus Niger, D-(+)-glucose, potassium hexacyanoferrate (III), potassium hexacyanoferrate (II) trihydrate, 1-ethyl- 3-[3-dimethylaminopropyl] carbodiimide hydrochloride (EDC), N-hydroxysuccinimide (NHS), 2-(N- morpholino) ethanesulfonic acid hydrate (MES), 3-mercaptopropionic acid, 2-mercaptoethanol, uric acid, ascorbic acid, 4-(2-hydroxyethyl)-1 -piperazineethanesulfonic acid (HEPES), KCI, MgSC , NaCI, CaC , NaOH, sodium salt, sucrose, hydrogen peroxide (H2O2), rhodamine labeled dextran dye, and glutaraldehyde were purchased from Sigma Aldrich Chemical Co. Disodium hydrogen phosphate and sodium dihydrogen phosphate monohydrate were purchased from Merck. Hydrochloric acid and sulfuric acid were purchased from J.T. Baker, and ammonia solution was obtained from AJAX. All chemicals were of analytical grade and were used without further purification. Aqueous solutions were prepared using Milli Q water (18.2 MQ.cm at 25 °C). Ferrocene cored poly(amidoamine) dendrimers (Fc-PAMAM) were synthesized as previously reported. Electrochemical measurements were carried out using electrochemical analyzer CHI 650E (CH Instruments, USA). The Objet Eden 260V 3D printing system was used to prepare MNA electrode holders, Au nanolayer deposition was preformed using the sputter deposition system Intlvac Nanochrome AC/DC, and O2 plasma treatment was performed using loN Wave 10 Plasma Asher. Field-emission gun scanning electron microscope (FEG-SEM) was used for morphological analysis of MN. Confocal imaging was carried by Nikon A1 Rsi.

Example 1 - MNA fabrication and preparation of MNA glucose patch - The sensing microneedle arrays (MNA) were fabricated by Deep Reactive Ion Etching (DRIE) of mono-crystalline <100> Si wafers. With this method, the size, length, and density of MNA as wells as the sharpness of MN can be tuned to any particular requirements. The first approach consists of a three-step process. First, a preliminary isotropic SF ₆ etching was used to create a sharp tip. This step was followed by a continued Bosch anisotropic etching until obtaining the desired length and then a reshaping step was performed by using a pseudo Bosch process with the purpose of creating a sharper conical tip. Exemplary resulting MNA arrays have length of approximately 250 pm, a base diameter of ~50 pm and a sharp tip with ~2 pm diameter, although these parameters may be varied as desired.

The exemplary sensing patch is designed as a three-electrode system, where each electrode is made of MNA. MNA patch fabrication starts with developing a holder/support for the Si-MNA substrate in this case using 3D printing technology although metals could also be used depending on the conducting nanofilm to be used (Figure 1 A). The Si-MNA holding base was 3D printed using Full Cure 720 (FC720) high-resolution, rigid and translucent printing material. After the fabrication, the 3D printed holder was immersed in 20% NaOH solution in order to remove the remaining residues of supporting material used during the printing. If not completely removed, this supporting material can decrease the adhesion of Au and subsequently have a negative effect on the biosensor performance. Next, Si-MNA substrate was attached to 3D printed holder using fast curing adhesive glue (Figure 1B). In the next step, a thin nanofilm of Au was deposited using an AC/DC sputter coater (Figure 1 C). Prior the Au nanofilm deposition, Si-MNA electrode were cleaned with O2 plasma for 1 min after which it was placed in AC/DC sputtering instrument containing Cr and Au targets powered by dual AC and DC power supplies for Cr and Au, respectively. First, 15 nm of Cr was deposited using dual AC with power of 300 W, which served as an adhesive layer for the subsequent Au layer. 150 nm of Au on Si was deposited using DC with target voltage of 100 V. During the deposition of both Cr and Au layers, the chamber vacuum and the sample rotation were kept at constant 2.6 x 10 ^-3 mbar pressure and 100 rpm, respectively. The resulting Au-coated electrodes are referred as Au-Si-MNA. After the sputtering process, a sticky tape test was performed on Au-Si- MNA electrode in order to assess the adhesion of Au layer. An insulating dye was applied on some parts (non-sensing, non connector parts) of Au-coated 3D holder/support (Figure 1D) in order to obtain an electrochemical signal only from the sensing MNA region only. The same electrode preparation procedure was used for the fabrication of the working (WE) and counter (CE) electrodes. The procedure for the reference (RE) electrode involves one extra step which is drop-casting of AgCI ink on the Au-Si-MNA electrode surface and drying at 60°C for 30 min. Finally, Au-Si-MNA electrodes were assembled into a single three-electrode glucose sensing patch by mounting on a frame (Figure 1 E) which is then suitable for electrochemical measurements, as well as in vivo application involving skin piercing. Prior to functionalization of the WE electrode, the Au-Si-MNA surface was cleaned in order to remove any remaining impurities and/or organic contaminants from electrode surface. First cleaning step was performed by thoroughly rinsing Au-Si-MNA electrode with ethanol and acetone. Next, the Au-Si-MNA was immersed in preheated base Piranha solution prepared using 20% ammonia solution, dhhO, and 35% H2O2 for 1 min at 80°C, and then thoroughly rinsed with dhhO. Surface functionalization was performed by incubating the Au-Si-MNA electrode in 2 mM 3- mercaptopropionic acid overnight, allowing self-assembling of thiol groups on the Au surface. Further on, the electrode was thoroughly rinsed with dhhO and the backfilling of any exposed Au surface was achieved by immersing the electrode into 2 mM 2-mercaptoethanol solution for 30 min. In the next step, electrode is incubated for 45 min in a mixture of 1 :1 10 mg/mL EDC and 15 mg/mL NHS solution, both prepared in 10 mM MES buffer pH 5. The Au-Si-MNA electrode was then immersed in 10 mM Fc-PAMAM solution prepared in 10 mM PBS, for 4 h. Subsequently, the electrode was thoroughly rinsed with dH ₂ 0 and incubated in 2.5% glutaraldehyde for 45 min. In the final step, functionalized Au-Si-MNA/Fc-PAMAM electrode was immersed in 15 mg/mL GOx solution prepared in 10 mM PBS and left overnight at 5°C. Finally, before proceeding with electrochemical measurements, the Au-Si-MNA/Fc-PAMAM/GOx electrode was once more rinsed with dhhO to remove unbound GOx. When not in use, Au-Si-MNA/Fc-PAMAM/GOx electrode was stored in 10 mM PBS at 5°C.

In vivo testing of the MNA glucose patch - In vivo testing and application of the exemplary three- electrode system glucose sensing patch was performed on mice anesthetized through inhalation of 1 .2% isoflurane in 2 L/min O2. Hair on the dorsal side of the mice was shaved off using hair removal cream after which shaved area was thoroughly cleaned with dH ₂ 0. Later, the MNA glucose patch (Figure 1E) was applied onto the dorsal side of mouse skin by thumb pressure and then electrochemical signal was recorded. In vivo electrochemical measurements were performed by recording amperometric signals before and after intramuscular injection of 120 pL of 30 wt% glucose solution. In parallel, glucose level in mouse blood was measured using a commercial glucometer (CareSens N). Obtained blood glucose level values from both glucose patch and glucometer were compared.

Results and Discussion - Morphological characterization of MNA and skin penetration testing

- Morphological characterization of the exemplary Au-Si-MNA was carried out using SEM. Figure 2A-D show scanning electron micrographs of Au-coated Si needles with detailed morphological features. Micrographs of top (Figure 2A) and tilted (Figure 2B) views show well-ordered MNA in hexagonal arrangement with density of -9500 microneedles/cm ² , where the space between two adjacent cone centres is -110 pm. A cross section image of MNA (Figure 2C) reveals a height of -250 pm, with a base diameter of -50 pm. At increased magnification, the sharpness of individual needle tip with the diameter of -2 pm was apparent (Figure 2D). Skin is composed of three layers with epidermis being the outermost layer with an average thickness of 150 pm, dermis being the second layer with a thickness ranging from 500 to 2000 pm and finally hypodermis. The sharpness of the presented MNs is ideal for the painless penetration inside the skin whereas the shape, size and density of the MNA are all suitable for the analysis of epidermis and dermis layers, according to the reported studies, without puncturing of blood vessels or damaging nerve bundles. This MNA was used for the preparation of the three-electrodes system glucose sensing patch. Furthermore, to demonstrate that MN can successfully and uniformly penetrate the skin, MNA was applied to porcine skin by finger pressure. Before applying to porcine skin, MNA was spin-coated with Rhodamine labelled Dextran dye (Figure 2E) and dried for 24 h. As shown in Figure 2F, MNA could generate uniform pores in arrangement that matches MNA patch on the surface of the porcine skin. Electrochemical characterization of MNA glucose patch - The electrochemical characterization of MNA glucose patch was performed using cyclic voltammetry (CV) and electrochemical impedance spectroscopy (EIS) methods. First characterization was carried out by potential cycling of Au-Si- MNA and flat Au-Si electrodes in 0.5 M H ₂ SO ₄ solution. Figure 3A shows CV plot of Au-Si-MNA and flat Au-Si electrodes, where the oxidation peaks of deposited Au nanolayer can be observed within 1 V to 1.2 V, and reduction peaks within 0.7 V to 0.9 V potential range. Charge of the reduction peaks can then be used to calculate the active surface area of the electrodes, assuming that a specific charge required for Au reduction is 390 pC crrr ² . Inset of the Figure 3A shows comparison of active surface area of the Au-Si-MNA and flat Au-Si electrodes which is 0.97 crrr ² and 0.33 crrr ² , respectively. This increase in active surface area is due to the high density of MNA.

Further characterization was carried out to monitor the surface modifications of Au-Si-MNA electrode with Fc-PAMAM and GOx. The functionalization of MNA towards glucose sensing includes the self-assembly of 3-mercaptopropionic acid to generate a carboxylic acid-terminated surface. Upon activation of -COOH groups, the -NFh-terminated Fc-PAMAM was immobilized. Then, GOx enzyme as the biological recognition element was attached to the surface using glutaraldehyde as linking molecule. Figure 3B shows CV plot of each electrode modification step starting with bare Au- Si-MNA electrode where no oxidation-reduction peaks were observed. After the immobilization of Fc-PAMAM on Au-Si-MNA and flat Au-Si electrodes, the oxidation-reduction peaks appeared at 0.45 V and 0.35 V, as a result of the redox activity of Fc core of the dendrimers. Besides, the difference in the intensity of oxidation-reduction peaks of Au-Si-MNA electrode and flat Au-Si modified with Fc- PAMAM can be observed, the Au-Si-MNA electrode exhibits a current peak intensity ~3 times higher than that of flat Au-Si electrode. This is due to the larger surface area of MNA (see Figure 3A) that results into more available sites for the Fc-PAMAM immobilization. The intensity of Fc current peaks drastically decreased when GOx enzyme was immobilized onto Fc-PAMAM as expected since GOx has shielding effect on the Fc molecule. In Figure 3C, CV plot of Au-Si-MNA/Fc-PAMAM electrode at different scan rates is illustrated. Increase in scan rate increased both anodic and cathodic peaks without shifting which shows reversable electron transfer ability of the electrode. The inset of the Figure 3C shows the anodic and cathodic peak currents versus the square root of scan rates. This linear behaviour shows the diffusion-controlled electron transfer property of the Au-Si-MNA/Fc- PAMAM electrode. Further characterization was performed by EIS measurements recorded in redox probe K ₄ Fe(CN) ₆ /K ₃ Fe(CN) ₆ (5 mM, 1 :1 ratio) containing 0.1 M KCI solution. Figure 3D shows EIS Nyquist plot of bare Au-Si-MNA (a) and flat Au-Si (b), and Au-Si-MNA electrodes modified with Fc- PAMAM (c), and GOx (d). EIS data was fitted using Randles electrical equivalent circuit. According to results represented in Figure 3D, the charge transfer resistance (Ret) values for bare Au-Si-MNA electrode is 20 W and flat Au-Si electrode Ret is 340 W. After the immobilization of Fc-PAMAM, Ret increases to 1 .22 kW and after GOx immobilization, the Ret increases to 6.17 kW. Obtained EIS and CV results complement each other as well as showing successful immobilization of Fc-PAMAM and GOx enzyme on Au nanolayer surface.

Biosensor optimization and glucose detection - Electrochemical optimization of MNA glucose patch was performed by testing its current response to glucose under different pH and temperature conditions. Current response to glucose is represented as relative response (%) with 100% being maximum obtained response. The optimization of the working pH for the MNA glucose patch was investigated by recording the response to glucose in buffer with different pH values ranging from 6.2 to 8.2 (0.1 M PBS). Free GOx enzyme operates under relatively broad pH range (pH 4 to 7) and its optimum working pH is 5.5, yet after immobilization, the optimum pH values of the enzyme can change. According to the results represented in Figure 4A, there was no significant response to glucose at pH 6.2 however, with the increase of pH, response to the glucose also increased reaching its maximum at pH 7.4. Further increase in pH resulted in decrease of sensor’s performance. Th determined optimum working pH for MNA glucose patch is 7.4. Determination of optimum working temperature of the glucose patch was carried out in 0.1 M PBS (pH 7.4) under different temperatures ranging from 25°C to 55°C. Recorded electrochemical response of biosensor to glucose, represented in Figure 4B, showed relatively good responses from 25°C to 35°C, with a maximum response at 35 °C. At higher temperatures, the response drastically decreased with temperature increase which can be due to the denaturation of GOx enzyme. The optimum working temperature for the glucose sensing MNA patch is 35°C. Furthermore, reusability of the sensing patch was studied. Experiments were performed in 0.1 M PBS by measuring the sensor response to glucose for 10 consequent times using the same patch (Figure 4C). The glucose biosensor kept 95% of its initial response during first five consecutive measurements after which the response started to slowly decrease to finally lose 20% of its initial response after 10 measurements.

The selectivity of the glucose biosensor is of high importance especially when working with ISF due to possible interference of other molecules present in ISF like uric acid (UA) and ascorbic acid (AA) with the electrochemical response of the enzyme electrode. In order to prevent this, Fc mediator was integrated into an enzyme immobilization matrix which did not just serve as an electron carrier during enzymatic reaction but also decreased the working potential of glucose patch. The use of a high working potential can lead to oxidation of electroactive compounds on the electrode surface leading to false positive response. For this reason, we conducted an interference study of the MNA glucose sensing patch using 2 mM glucose and 0.1 mM AA, UA, dopamine (DA), lactic acid (LA) and glycine (gly) as interfering species. AA, UA, DA and LA are one of the most important electroactive interferons in ISF because of their ability to self-oxidize on the electrode surface and thus effect sensors response. As seen in Figure 5A, 0.1 mM of AA, UA, DA, LA, and gly had no significant effect on the electrochemical signal of glucose patch. These optimization and interference study data show that the MNA glucose patch exhibits a suitable performance in relevant physiological pH and temperature with no significant interference from physiologically relevant interferons.

Analytical characteristics of the MNA glucose patch were validated in PBS (0.1 M pH 7.4) and in artificial ISF (alSF) (pH 7.4), respectively. The operating potential of MNA glucose patch is dictated by Fc mediator which was used together with PAMAM dendrimers (Fc-PAMAM) during the preparation of working electrode. The main purpose of using PAMAM dendrimers is to provide large number of functional groups for GOx enzyme immobilization and in the same time to keep Fc mediator located between the electrode surface and enzyme layer. At this specific position, Fc serves as mediator which shuttles the electron towards electrode surface and decreases working potential of the biosensor. Figure 5B shows a chronoamperometric graph obtained from the MNA glucose patch responding to the increasing glucose concentrations in 0.1 M PBS. The biosensor exhibits well defined response to glucose in PBS having linear dynamic range from 1 to 11 mM with sensitivity of 0.58 mA/ (mM cm ² ), detection limit of 0.45 mM and correlation coefficient of 0.9996 (Figure 5D). Since MNA glucose patch will be used for transdermal glucose analysis from ISF, it is important to examine the performance of biosensor in an environment similar to ISF. For this experiment, we used artificial ISF (alSF) prepared in dhhO using 10 mM HEPES, 123 mM NaCI, 7.4 mM saccharose, 2.2 mM CaCh, 3.5 mM KCI, 0.7 MgSC , and 1 .5 mM NaFhPC with pH adjusted to 7.4. Figure 5C shows a chronoamperometric graph obtained with the MNA glucose patch for the increasing concentration of glucose in alSF. The linear dynamic range of biosensor is characterized by a sensitivity of 0.1622 mA/ (mM cm ² ), linear range and detection limit of 1 to 9 mM and LOD of 0.66 mM (with correlation coefficient of 0.9995), respectively. Notice that glucose concentration found in healthy individuals ranges from 3.6 to 6 mM. The observed differences in sensor’s response (Figure 5D) in PBS and alSF can be due to different compositions of testing solutions. alSF is a more complex medium which can affect the enzyme activity however, the sensing performance obtained with alSF remains good enough to be used for the further in vivo testing.

In vivo glucose monitoring - After the evaluation of the analytical performance of glucose biosensor in vitro, as a proof of concept, we performed in vivo testing of the MNA glucose patch on mice. Prior to the in vivo measurements, the mouse was anesthetized and hair on dorsal side of the mice was shaved off after which MNA glucose patch was applied. Figure 6A shows three-electrode MNA glucose patch with ~1 .2 cm size composed of C and R electrodes, and W electrode previously modified with Fc-PAMAM and GOx enzyme. On-mouse application of the patch was gently supported with finger (Figure 6B) and after removal of MNA glucose patch, the indentations created by the MNA patch can be clearly seen with no sign of bleeding (Figure 6C). Figure 6D displays electrochemical results of on-mouse MNA glucose patch obtained from three different mice. First measurements were recorded for five minutes in order to ensure a stable electrochemical signal and at the same time to recorded base line for 60 s before the glucose injection. After obtaining base line for glucose measurements, intramuscular injection of 120 mI_ of glucose solution (30% w/v) was performed allowing ~30 min in order for the blood glucose level to reflected in the ISF. Note that glucose concentration found in blood is similar to ISF except that increase in glucose concentration in ISF has slight lag time due to low capillary density in dermis layer and slow metabolic activity in epidermis layer. This ISF lag time is particularly important for MNA glucose patch because the sensing process is carried out in dermis and epidermis layers, for this reason we had to allow enough time for diffusion of injected glucose into ISF. Later on, the amperometric measurements were performed once again. In the same time, blood glucose level of the mice was recorded using commercially available glucometer which was measured for several times in parallel with electrochemical measurements. Blood glucose measurements using glucometer allowed us to follow increase in blood glucose starting after injection of glucose solution until electrochemical recording using MNA glucose patch. A distinct increase of current was observed before and after glucose injection (Figure 6D). Inset of Figure 6D shows the concentration change (representing total amount of glucose increase after glucose solution injection) of glucose recorded with MNA sensing patch, and with commercial glucometer. From the obtained results, it is possible to conclude that the ISF glucose level detected by MNA patch and glucose level in blood recorded with glucometer correlate very well. Even though intramuscular injection of glucose solution was exactly the same for all three mice used during in vivo experiments, their glucose levels vary possible due to their different metabolism, difference in the diet and stress developed during the experimental procedure. Overall, in vivo results suggest that the glucose in ISF detected by MNA patch corresponds to the glucose level found in blood which illustrates the successful proof of concept application of the glucose MNA sensing patch.

Example 2 - Porous silicon microneedle arrays for glucose sensing - Preparation of microneedles - Silicon microneedle arrays (Si-MNA) were fabricated as previously described by means of UV photolithography and DRIE.

Porosification of Si-MNA of porous silicon layer - Porous silicon microneedle arrays (PSi-MNA) were prepared by electrochemical etching from Si-MNA. The reduced size of the pores (< 100 nm) prevents the deposition of metal coatings inside them, therefore sputter-coating of Au cannot be used to deposit conformal conductive nanofilms layers on nanostructures to increase the conductivity of PSi-MNA and enable their use as electrodes. Therefore, a thermal treatment at high temperature (525-800 °C) in the presence of acetylene was employed to generate a conformal, nanometric carbon (carbonised) layer that provides PSi-MNA with enhanced electrochemical properties and outstanding chemical stability. The carbon layer provides key characteristics required for electrochemical biosensing.

The electrochemical etching was performed by applying a current between highly doped p- type Si-MNA (anode) and a Pt wire (cathode) using HF-based solutions as electrolytes. Typically, a porous silicon layer was etched at 40 mA cm ^-2 in HF/EtOH 1 :1 for a time that was adjusted between 30 s and 8 min.

Chemical modification - thermal carbonization - Si-MNA were modified by thermal carbonization with acetylene. The treatment method was a two-step carbonization at 525 and 800 °C. The freshly- etched, hydrogen-terminated PSi-MNA were inserted in a quartz tube under N2 flow (1 L min ^-1 ) for 45 min at room temperature to remove oxygen and adsorbed moisture. In the first step of carbonization treatment, an acetylene flow of 1 L min ^-1 was added for 15 min, after which the quartz tube was placed in a tube furnace at 500 °C for 15 min under 1 :1 N2~acetylene flow. After this first carbonization treatment, the tube was allowed to cool back to room temperature under N2flow. For the second step, acetylene flow (1 L min ^-1 ) was added for 10 min at room temperature followed by annealing of the sample for 10 min at 800 °C under N2 flow. Finally, the thermally carbonized electrode (TCPSi-MNA) was allowed to cool back to room temperature under N2 flow. Electrografting of 4-carboxyphenyi diazonium salt - 4-carboxyphenyl diazonium salt was generated in situ by adding NaNC>2 to an acidic solution of aminobenzoic acid (ABA). 20 mM ABA was prepared in 10 mL (0.5 M) HCI solution and was left to degas in N for 10 min. After, 50 mI_ of 100 mM NaNC> were added into the ABA solution and continued degassing for 30 min in an ice bath. The carbonized PSi-MNA were electrog rafted with the generated diazonium salts by cyclic voltammetry scanning between -0.60 to 0.60 V for five cycles at 100 mV s ^_1 . The electrog rafted TCPSi was then rinsed with water, followed by cyclic voltammetry in 0.1 M PBS (10 cycles, between -0.20 to 0.60 V at 100 mV s ^_1 ) (see Figure 1).

Immobilization of Fc-PAMAM and enzyme - Fc-PAMAM was covalently bound to the end of the electrog rafted PSi-MNA surfaces by activating the terminal carboxyl group using EDC/NHS chemistry. Then, glutaraldehyde was used as a cross-linker between the Fc-PAMAM and GOx. PSi- MNA were incubated in 2.5% glutaraldehyde for 45 min, then the electrodes were rinsed with PBS to remove any unattached compounds. Finally, the PSi-MNA electrodes were incubated in a GOx solution overnight at 4 °C, to enable enzyme immobilization on the electrode surface (Figure 9). Results and discussion

Morphological characterization of PSi-MNA - SEM was employed to morphologically characterized the formation of the porous layer on MNAs. Figure 10 shows SEM micrographs of the resulting structures after electrochemical etching. The nanoporous MNAs retain their morphological properties after electrochemical etching (Figure 10A). The nanoporous film grew conformally around the projections and on the base of the substrate. A high-magnification image (Figure 10B) revealed the size of the nanopores, which were calculated to have an average diameter of 50-60 nm on the outer surface. The thickness of the porous layer can be controlled with the electrochemical etching conditions, as the thickness is proportional to the etching time. Several etching times were applied and the resulting PSi-MNA samples were characterized by SEM. Figure 11 shows cross sectional images of the PSi-MNAs after applying three different etching times, i.e. 4, 6 and 8 min. In Figure 11A-C the formation of a thin porous layer can be observed, which increased with increasing etching time. To confirm the conformality of the nanoporous film, MPAs were mechanically fractured to facilitate the imaging of the microprojection core. Figure 11D-F show a detail of fractured MPAs where conformal porous layer and a solid Si core can be observed. The thicknesses of these films were similar to those measured at the bottom of the MPA, confirming conformal growth of the nanoporous film.

Characterization of thermal carbonised layer (TCPSi) - Fourier transform infrared (FTIFt) spectroscopy characterization of TCPSi before and after diazonium salt electrografting was performed as shown in Figure 12. FT-IR spectrum for carbonized pSi (TCPSi) shows absorption bands at 1018 cnr ¹ (Si-O), 3050 cnr ¹ and 2870 - 2970 cnr ¹ (C-H), and 1601 cnr ¹ (C=C) (50). TCPSi- ABA displayed characteristic bands of carbonyl group (C=0) vibration at 1709 cnr ¹ , two band peaks of C=C vibration of the aromatic ring at 1601 cnr ¹ and 1541 cnr ¹ , and a mixture of diazo (=C=N=N) compounds, C-H and C-N stretching at 1387 cnr ¹ . The TCPSi-ABA electrog rafted nanocomposites shared identical characteristic bands at 1018 cnr ¹ (Si-0 bending) and 781 cnr ¹ (Si-C) with the TCPSi electrode. This demonstrates that ABA was successfully electrografted onto the porous layer. EIS was used to confirm probe the features of TCPSi films before and after ABA electrografting. The resulting Nyquist impedance spectra is shown in Figure 13. The data was fitted to a Randles circuit model (inset Figure 13), from which from which charge transfer resistance (R _et ), double layer capacitance (C _di ), solution resistance (R _s ), and Warburg element (diffusion impedance) (Z _w ) were extracted. The value of the R _ct after ABA electrografting is approximately four times higher than that of the bare TCPSi (33.17 W TCPSi to 122.20 W TCPSi- ABA). Therefore, this indicates ABA was successfully electrografted on the TCPSi layer.

CV is another method to characterise the stepwise modification of the electrode. The advantage of CV over EIS is the ability to study the electrocatalytic effects of the modified electrodes. Since CV provides information about the reduction and oxidation peak potentials and current intensities ascribed to the embedded mediator, it was the chosen electrochemical technique to characterize the TCPSi/ABA/Fc-PAMAM electrode before and after GOx immobilization. The information obtained from CV was then used to determine the optimal applied voltage for the amperometric measurement of glucose. CVs were recorded by scanning the potential from 0 to 0.7 V at 100 mV/s in 0.1 M PBS (pH 7.4). The redox peaks assigned to ferrocene in the Fc-PAMAM were mostly supressed after the immobilization of the enzyme (Figure 14). The enzyme blocks the Fc-PAMAM mediator from undergoing oxidation and reduction on the electrode surface. These CVs indicate that GOx was immobilized and still active upon covalent binding to the electrode surface, and Fc was available to perform its role as redox mediator.

The CV plots can determine the optimal applied potential for amperometric measurements. In the case of oxidation reactions, the potential to be applied should be equal or larger than the potential of the oxidation peak. The ideal applied potential for amperometric current response was found to be 0.4 V. As seen in Figure 15, CVs at fast scan rates lead to a decrease in the diffusion layer; as a consequence, higher currents are observed. The data show the electron transfer of the redox mediator on the electrode is a reversible process, characterized by a similar E _p at every scan rate. The representation of l _pc and l _pa with respect to the square root of scan rate shows a linear dependency, indicating that the electron charge propagates in a diffusion-like process with a fast electron transfer rate. The fast electron transfer of TCPSi electrode might be attributed to the carbon properties and the pores which allow for more electron-surface interactions. The high surface area of the TCPSi may increase and promotes the electron transfer rate, but this was not realized on the pillars electrode. The electrodes can efficiently facilitate the electron transfer process between the solution and the electrode surface via the ferrocene core mediator. These results further support that Fc-PAMAM is covalently bound to the electrode surface.

Electrochemical measurement of glucose - A current-time amperometric response was used to determine glucose concentrations. It is expected that the electrochemical response will be proportional with every increase of glucose concentration. Measurements for glucose concentrations between 0 to 10 mM with increments of 1.0 mM concentration were recorded. The upper glucose concentration range was explored to study the electrode linear working range. Current values plotted to obtain the calibration curve were calculated by subtracting the difference between the current measured and the initial baseline. Figure 16 shows the amperometric response exhibited by increasing glucose concentrations using a TCPSi/ABA/Fc-PAMAM/GAD/GOx modified electrode. This electrode was found to have a limit of detection of 24 nM, a linear sensing range from 0.6 to 6 mM, a response time of 20 s, and a sensitivity of 1 .33 mA mM ^-1 cm ^-2 .

The electrochemical etching of highly doped p-type Si-MNA enables the formation of nano sized architectures within the microstructures. The results demonstrate the formation of a nanoporous film that grows conformally around the MNAs. The dimensions of this film (e.g. pore diameter preferably of <100 nm, most preferably 40 - 60 nm; film thickness, preferably of 15 microns, most preferably of from 0.1 to 7.5 microns) can be controlled by adjusting the electrochemical etching parameters such as current density and etching time. The nanoporous architectures can be used as electrochemical transducers. As a proof of concept, PSi electrodes were engineered for the construction of novel enzymatic sensors. Due to the small size of the nanopores, metallization via sputter-coating was not a feasible option for generating conductive PSi-MPAs. A thermal carbonization treatment in the presence of acetylene was used instead. Electrochemical characterization of the TCPSi electrodes revealed a good electrochemical behaviour, with fast electron transfer kinetics. The TCPSi electrodes were chemically modified via the electrografting of in situ generated 4-carboxyphenyl diazonium, which introduced carboxylic acid functionality. The enzyme, GOx, was immobilized on the TCPSi electrode through a redox mediator dendrimer conjugate (e.g., Fc-PAMAM) covalently bound to the surface of TCPSi/ABA surfaces. Amperometry was applied to measure glucose concentration at 0.40 V. Results demonstrated a detection range from 0.6 to 6 mM, a detection limit of 24 nM, a response time of 20 s and a sensitivity of 1 .33 mA mM ¹ cnr ² . The large surface area of PSi electrodes together with the optimal electrochemical behaviour of the carbonized layer are believed to be responsible for the excellent sensitivity. These results demonstrate that PSi-MNA configurations show great potential as biosensing platforms for skin application and detection of bioanalytes in the dermal and epidermal skin layers.

Example 3 - Electrochemical Immunosensor for Breast Cancer Biomarker Detection Using High-density Silicon Microneedle Array - In this work, high-density gold coated silicon microneedle arrays (Au-Si-MNA) were simultaneously used as biomarker extraction platform and electrochemical transducer, enabling the selective immunocapture of epidermal growth factor receptor 2 (ErbB2), a key breast cancer biomarker, and its subsequent quantification. Healthy individuals have blood ErbB2 concentrations ranging from 2 to 15 ng/mL whereas ErbB2 concentration in breast cancer patients is above 15 ng/mL. The analytical performance of the device was tested in artificial interstitial fluid exhibiting linear response in a wide concentration range from 10 to 250 ng/mL, with a detection limit of 4.8 ng/mL below the biomarker levels expected in breast cancer patients. As a proof of concept, the immunosensor demonstrated its ability to successfully extract ErbB2 from model a mimicking the epidermis and dermis layers, and subsequently quantify it showing a linear range from 50 to 250 ng/mL and a detection limit of 25 ng/mL. The uniqueness of this sensing platform combining direct biomarker transdermal extraction and quantification opens up new avenues towards the development of highly performing point-of-care devices.

Materials - 3-Mercaptopropionic acid, 2-mercaptoethanol, 2-(N-morpholino)ethanesulfonic acid hydrate (MES), 1 -ethyl-3-[3-dimethylaminopropyl] carbodiimide hydrochloride (EDC), N- hydroxysuccinimide (NHS), polyethylene glycol) (PEG) bis(amine) (Mw: 3000), potassium hexacyanoferrate (III) (K3[FeCN6]), potassium hexacyanoferrate (II) trihydrate (K ₄ [FeCN ₆ ]-3H ₂ 0), 4- (2-hydroxyethyl)-1-piperazineethanesulfonic acid (HEPES), NaCI, NaOH, MgSC , CaC , KCI, sucrose, gelatin, agar, glycerol, insulin, and goat IgG were purchased from Sigma Aldrich Chemical Co. H2O2, Na ₂ HP0 ₄ and NaH ₂ P0 ₄ were purchased from Merck. Ammonia solution was obtained from AJAX. H2SO4 and HCI were purchased from J.T. Baker. All chemicals were of analytical grade and were used without further purification. Anti-HER2 antibody and ErbB2 were kindly supplied by Dr Adams from CSIRO. The humanized anti-HER2 monoclonal antibody, trastuzumab (Herceptin), was a gift from Dr. Sam Greenall (Hudson Institute of Medical Research, Melbourne). Human recombinant ErbB2 ectodomain corresponding to the mature, N-terminal 623 amino acids and incorporating a C-terminal Flag epitope tag (DYKDDDDK) was purified by immunoaffinity and size- exclusion chromatography from the supernatant of HEK293-EBNA cells stably transfected with the pAPEX-3P mammalian expression vector encoding ErbB2623.

Immunosensor Preparation - Si-MNA-based electrodes (Figure 17a) were fabricated using procedures previously reported. Here briefly, MNA were fabricated using UV-photolithography and deep reactive ion etching (DRIE) (Plasmalab 100 ICP380) of a mono-crystalline <100> Si wafer. The Si wafer was pre-patterned using AZ 40XT photoresist (MicroChemicals) and a chromium mask via UV exposure, after which the photoresist was developed in AZ 726 MIF (MicroChemicals). The DRIE etching process consisted of controlled pseudo-Bosch anisotropic etching of the pre-patterned Si wafer mixing SF6 and C4F8 gases which was performed until the desired height of MNs was obtained. Sharpening of MNs and removal of the remaining resist was carried out by applying a second pseudo-Bosch anisotropic etching process. In the next step, the Si-MNA substrate was mounted onto an electrode support in the form of a 3D-printed holder made of Full Cure 720 (FC720) high-resolution, rigid and translucent 3D printing material (Objet Eden 260V 3D printing system) using fast curing adhesive gel (Figure 17a(i)) obtaining the Si-MNA electrode. A thin film of Au was deposited using AC/DC sputtering instrument (Intlvac Nanochrome AC/DC) which firstly deposited 15 nm of Cr as adhesive layer and then 150 nm of Au (Figure 17a(ii)). This electronically conducting nanofilm was provided on the sensing MNA substrate needles/projections and the non sensing support regions including the non-sensing connector heads. Later, an insulating ink was applied on the Au-coated holder, particularly the non-sensing parts of thereof, in order to limit the source of electrochemical signal only to the sensing Au-coated Si-MNA (Au-Si-MNA) region (Figure 17a(iii)). A similar electrode preparation procedure was used for the fabrication of the counter and reference electrodes, with an additional step in the reference electrode preparation which involved drop-casting AgCI ink on the electrode surface and drying it at 60°C for 30 min. Au-Si-MNA as used herein means a PSi MNA-Cr-AU electrode.

Prior to functionalization (Figure 17b), the Au-Si-MNA electrode surface was cleaned first by thoroughly rinsing the electrode with ethanol and acetone, and then the electrode was immersed in base Piranha solution composed of aqueous solution of 1 .5 wt% ammonia, and 5.4 wt% H2O2, for 1 min at 80°C, followed by rinsing with dhhO. The electrode was then incubated in a 2 mM aqueous solution of 3-mercaptopropionic acid overnight, allowing the formation of a self-assembled monolayer (SAM) via reaction of thiol groups with the Au surface. After rinsing the electrode with dhhO, it was immersed in a 2 mM aqueous solution of 2-mercaptoethanol for 30 min for the purpose of the backfilling of any exposed Au surface which is believed to reduce non-specific binding onto the surface reducing false positive results. Next, the electrode was immersed in a 1 :1 mixture of 10 mg/mL EDC and 15 mg/mL NHS prepared in 10 mM MES buffer pH 5 for 45 min. The NHS ester functionalized Au-Si-MNA electrode was then immersed in 25 pg/mL anti-HER2 antibody solution prepared in 10 mM PBS for 4 h. The electrode was then rinsed with 10 mM PBS in order to remove unbound anti-HER2. The residual NHS sites were blocked by reaction with NH2 functionalized PEG (1 mg/mL solution) for 30 min. This step is also believed to reduce non-specific binding onto the surface reducing false positive results. Finally, the Au-Si-MNA anti-HER2 electrode was once more rinsed with 10 mM PBS and, when not in use, it was stored in 10 mM PBS at 5°C.

Electrochemical Detection - A CHI 650E (CH Instruments, USA) electrochemical analyzer was used for all electrochemical measurements. Differential pulse voltammetry (DPV) measurements were performed by scanning the potential from -0.3 to 0.5 V in an electrochemical setup including Au-Si-MNA anti-HER2, Au-Si-MNA, and Au-Si-MNA modified with AgCI ink, as working, counter and reference electrodes, respectively. All measurements were performed in a solution of 5 mM K ₄ [Fe(CN) ₆ ] and 5 mM K3[Fe(CN)6] in 0.1 M KCI using MNA based working, counter, and reference electrodes. The anti-HER2-modified Au-Si-MNA working electrode was incubated during 1 h in artificial interstitial fluid (alSF) spiked with known concentrations of ErbB2 ranging from 10 to 250 ng/mL. alSF consisted of 10 mM HEPES, 3.5 mM KCI, 123 mM NaCI, 1.5 mM NaH ₂ PC> ₄ , 7.4 mM sucrose, 2.2 mM CaC , and 0.7 mM MgSC prepared in dH ₂ 0 and the pH was adjusted to 7.4. The antibody-antigen binding was monitored through the changes in the current intensities obtained from DPV measurements performed before and after incubation of alSF spiked with ErbB2 (Figure 17c). Control experiments were carried out using a working electrode modified with a goat IgG, following the same immobilization method used for the anti-HER2-modified Au-Si-MNA electrode. Error bars illustrate the standard deviation (SD) of three measurements (n = 3) unless otherwise stated. Preparation of Skin-Mimicking Phantom Gel - Phantom gels simulating epidermis and dermis layers were prepared as follows. The epidermis mimicking layer is composed of 10 wt% gelatin and 5 wt% of glycerol prepared in alSF, after which 0.1 wt% glutaraldehyde was added to the mixture. Dermis mimicking layer is composed of 24 wt% gelatin and 1 wt% agar prepared in alSF. The phantom gel was fabricated using 3D printed moulds. The moulds used for epidermis and dermis layer fabrication had a height of 120 miti and 1 .5 mm, respectively. The mixture used for dermis layer fabrication was initially heated up to ~70°C under constant mixing until all constituents completely dissolved. The final solution was cooled to 40°C and then it was casted onto the mould and left to cool for 2 h at room temperature. The gel samples used for the calibration plot experiments were prepared by spiking them with known concentrations of ErbB2 biomarker (10-250 ng/mL) prior to pouring the solution in the mould. Later, the epidermis layer mixture, also heated up to ~70°C and cooled to 40°C under constant stirring, was poured on top of the solidified dermis layer and left to cool down at room temperature. Detection of ErbB2 using Au-Si-MNA/anti-HER2 immunosensor was performed by extracting the biomarker from the phantom gel and analyzing the changes observed in the DPV signal caused by the binding of ErbB2. The immunosensor was gently inserted in the phantom gel, incubated for 1 h, thoroughly washed with PBS after removing it from the phantom gel, and finally used for DPV analyses in K ₄ [Fe(CN) ₆ ]/K ₃ [Fe(CN) ₆ ] solution.

Results and Discussion - Morphological Characterization of MNA-based Immunosensor - A field-emission gun scanning electron microscope (FEG-SEM, FEI NovaNano SEM 430) was used to characterize the Au-Si-MNA electrode. Figure 18 illustrates SEM micrographs of tilted (Figure 18), top (Figure 18b), and cross-sectional (Figure 18c) views of Au-Si-MNA. Micrographs show hexagonal arrangement of MN with a density of -9500 MNs/cm ² with MN height of -250 pm, a base diameter of -50 pm, and space between two adjacent cone centres of -110 pm. At increased magnification represented in the inset of Figure 18c, the sharpness of individual needle tips of -2 pm diameter was apparent. Density, shape, size, and sharpness of fabricated MNA are ideal for the painless penetration of stratum corneum and interfacing the ISF in the epidermis and dermis layers of the skin, since the thickness of the epidermal layer can range from 36 to 61 pm on human forearm and 100 pm on the wrist, with the dermis layer ranging from 500 to 2000 pm.

Electrochemical Characterization - Au-Si-MNA electrodes were electrochemically characterized using cyclic voltammetry (CV), DPV and electrochemical impedance spectroscopy (EIS) using Au- Si-MNA as working and counter electrodes, and Au-Si-MNA modified with AgCI ink as reference electrode. First, a CV in 0.5 M H ₂ SO ₄ was applied to determine the active surface area of the working electrode (Figure 19a). The active surface area can be calculated using the charge of the reduction peak divided by the charge per microscopic unit area of Au which was previously determined to be 390 pC cm ^-2 . In Figure 19a, the oxidation peaks of the deposited Au layer can be observed at 0.93 and 1.20 V, and reduction of oxidized Au can be seen at 0.7 V. The active surface area of Au-Si- MNA electrode is illustrated in the inset of Figure 19a and was found to be 0.97 cm ² , four times higher than the geometric surface area. This increase in active surface area is attributed to the high density of MNA which allows to fit -9,500 MNs/cm ² .

The surface modification of the Au-Si-MNA electrode was electrochemically followed to confirm the immobilization of anti-HER2 antibody on MNA and binding of ErbB2 antigen. DPV and EIS measurements, represented in Figure 19b and Figure 19c, respectively, confirm successful step-by-step modification of MNA by SAM formation and immobilization of anti-HER2 antibody, and binding of ErbB2 antigen. As expected, the current intensity decreased, and charge transfer resistance (Ret) increased for the Au-Si-MNA electrode upon Au surface functionalization with SAM and immobilization of anti-HER2 antibody. To demonstrate the feasibility of the Au-Si-MNA/anti- HER2 sensor to detect ErbB2, ErbB2 antigen was incubated on the surface, causing further decrease in the peak current and increase in Ret which indicates successful antigen-antibody interaction. Both DPV and EIS results confirm immobilization of anti-HER2 on the electrode surface as well as successful binding of ErbB2 antigen. Furthermore, the redox properties of the Au-Si- MNA/anti-HER2 sensor were studied using CV by applying different scan rates ranging from 10 to 150 mV/s. The CV shows well-defined oxidation-reduction peaks of the ferro/ferricyanide redox couple and, as scan rate increased, potentials of oxidation and reduction peaks shifted in positive and negative directions, respectively. Increase in scan rate resulted in increase of oxidation and reduction peaks, indicating aquasi-reversible process. In CV plots of oxidation and reduction current peak values show the linearity of the plot for both oxidation and reduction peaks shows a diffusion- controlled electron transfer process occurring at the MNA electrode surface.

Spectroscopic Analysis - X-ray photoelectron spectroscopy (XPS) analysis was employed to characterize the surface modification of Au-Si-MNA working electrode at each step of the functionalization process: pre-cleaned Au, SAM of 3-mercaptopropionic acid, and immobilization of anti-HER2 antibody (Figure 19d). The atomic composition of Au decreased as the functionalization progresses, as represented by the significant decrease in the Au/C ratio, which was consistent with the modification of the bare Au-Si-MNA surface with SAMs and the attachment of the anti-HER2 antibody. The efficient coverage of pre-cleaned Au-Si-MNAs with carboxylic acid terminated alkanethiols was confirmed by the increase in sulfur and oxygen atomic concentrations. The S/C and O/C atomic ratio after SAM functionalization are distinctly above the ratios of untreated Au-Si- MNA, confirming the presence of -COOH terminated SAMs. The high-resolution C1s XPS curves (Figure 19d) shows a main peak centered at -285.0 eV (C1 component), which can be ascribed to aliphatic carbons from alkyl chains. The carbon present on pre-cleaned Au-Si-MNAs most likely indicates the presence of adventitious carbon. After SAM modification, a significant increase in the C4 component (representative of 0-C=0- functionalities centered at -288.5 eV) becomes evident. The shoulder centered at -286.5 eV (C2 component) represents the carbon next to the sulfur atoms. Following EDC/NHS activation and incubation with anti-HER2, the bioconjugated MNA displayed a significantly higher carbon and nitrogen content compared to SAM-modified Au-Si-MNA. After anti- HER2 immobilization, the C1 s spectrum was significantly altered. While the C1 component can still be attributed to aliphatic carbons, the higher binding energy peaks can be assigned to the various C-N and C-0 based functional groups in the polypeptide (C2) and the characteristic peptide bonds of the immobilized antibody (C3). The latter components (C2, C3) were observed to increase significantly in intensity, confirming the presence of the antibody. The carboxylic acid signal (C4) originating from the SAM is no longer detectable, indicating an efficient immobilization of the antibody onto the activated SAM layer. Optimization of the MNA-Based Immunosensor - In order to maximize the performance of the immunosensor, the experimental parameters such as the concentration of anti-HER2 used for the electrode modification and ErbB2 incubation time were optimized. To evaluate the immobilization of anti-HER2 antibody, the MNA electrode was incubated in 10 mM PBS solutions containing different concentrations of antibody for 1 h. Antibody concentrations used in this study ranged from 4 to 75 pg/mL. An increase in current density change ( Al) with increase of the antibody concentration was observed. D/ is the difference of the oxidation current density measured before (h) and after incubation (/) of Au-Si-MNA/anit-HER2 with ErbB2 biomarker (D/=/o - / ). Current density change reached a plateau for the immunosensor prepared with 25 pg/mL anti-HER2 antibody which indicates saturation of the electrode surface with antibody. Thus, 25 pg/mL was selected as the optimum concentration for the fabrication of the immunosensor. Various incubation times were tested to determine the optimum time to maximize the binding of ErbB2 to immobilized anti-HER2. The change in current density of the recorded DPV of Au-Si-MNA/anti-HER2 incubated with the same concentration of ErbB2 biomarker (150 ng/mL) for different time intervals ranging from 10 min to 2 h and drastically increased during the first 40 min of incubation and later reached its maximum at 60 min. Further increase in the incubation time did not have significant effect on the current change. Therefore, 60 min was selected as optimum incubation time.

Detection of ErbB2 Breast Cancer Biomarker - A set of control experiments was performed using goat IgG in order to confirm the specificity of developed immunosensor towards ErbB2. Figure 20a shows DPV plots of Au-Si-MNA/anti-HER2 and Au-Si-MNA/lgG electrodes, recorded before and after incubation in alSF containing 150 ng/mL of ErbB2 biomarker. The current density change for IgG modified electrode (control) was insignificant comparing to the anti-HER2 electrode (inset of the Figure 20a). This result confirms that Au-Si-MNA/anti-HER2 sensor has high specificity towards ErbB2.

The analytical performance of MNA immunosensors was studied by incubating Au-Si- MNA/anti-HER2 sensor in alSF spiked with different concentrations of ErbB2 biomarker. The sensing principle of the MNA immunosensor is based on determining the changes in the diffusion of the redox probe towards the electrode surface as a result of the ErbB2 biomarker binding to the antibody modified Au-Si-MNA. Figure 20b shows the DPV signal obtained by scanning the potential from -0.3 to 0.5 V after incubating the working electrode in alSF with different concentrations of ErbB2 ranging from 10 to 250 ng/mL for 1 h. Increase in ErbB2 concentration gradually decreased the intensity of the current peak produced by the oxidation of K ₄ [Fe(CN) ₆ ]. This correlation is demonstrated in Figure 20c, where a linear increase in the current density change is observed with increasing ErbB2 concentrations. The change in current density ( Al) represented in Figure 20c corresponds to the difference of the oxidation current density measured before and after incubation of Au-Si-MNA anit-HER2 with ErbB2 biomarker (D/=/ o -I) at 0.15 V. The dynamic range of the immunosensor ranged from 10 to 250 ng/mL with a correlation coefficient of 0.9987 and linear regression equation of Al = 0.3978[ErbB2] + 10.282. This linear range of the MNA-based immunosensor is suitable for the detection of ErbB2 in breast cancer patients since the ErbB2 concentration in serum ranges from 33.2 to 166.6 ng/mL and in cancer tissue can reach 65.38 ng/mL. The limit of detection of the MNA-based immunosensor is 4.8 ng/mL which is in the range of the ErbB2 concentration found in blood of healthy individuals (2-15 ng/mL).

The MNA-based immunosensor was tested against potentially interfering components prepared in alSF such as glucose (4 mM), glycine (100 mM), insulin (150 ng/mL), T4 bacteriophage (10 ⁵ pfu/mL), and only NaCI (123 mM) prepared in dhhO. Selected potentially interfering molecules such as glucose, glycine, and insulin are likely to be found in ISF, whereas T4 bacteriophage was selected based on its size and likeliness that similar sized species potentially could affect the sensors performance. Figure 20d shows a column plot derived from the obtained DPV signal from which no significant changes are observed for components other than ErbB2, suggesting that the MNA-based immunosensor has an excellent selectivity for ErbB2 biomarker when compared to potential interfering molecules at high concentrations. Furthermore, the effect of different types of media such as 0.1 M PBS (pH 7.4) and alSF (pH 7.4) on the performance of the immunosensor was tested. As shown in Figure 20d no significant change in current response was observed showing that the functionality of the immunosensor does not depend on the medium which also confirms its stability. Detection of ErbB2 Biomarker in Skin-Mimicking Phantom Gel - The performance of Au-Si- MNA/anti-HER2 sensor was studied using a skin-mimicking phantom gel (Figure 21a-e) that was fabricated using a 3D printed mould (Figure 21c). The phantom gel was prepared containing ErbB2 biomarker at a 10-250 ng/mL concentration. Due to the differences in the chemical composition of the phantom model (see section entitled preparation of skin-mimicking phantom gel) we could model both epidermis and dermis layers, with the mock epidermis layer (100 ± 40 pm thickness) being clearly distinguished on top of the mock dermis layer in the fabricated phantom gel (Figure 21d). First, this gel was used for testing the penetration of the MNA-based sensor. The footprint of the MNA in the form of hexagonally arranged circular microscale craters can be observed in Figure 21e, demonstrating successful penetration of the MNA. Figure 21a shows the calibration curve derived from the DPV plots that was obtained after ErbB2 extraction from the phantom gel and subsequent electrochemical detection. The linear range of the immunosensor in skin-mimicking phantom gel is from 50 to 250 ng/mL with a detection limit of 25 ng/mL, allowing the detection of ErbB2 biomarker in breast cancer tissue. When compared to the linear range and detection limit obtained in alSF, the aforementioned values of the immunosensor in phantom gel are lower, possibly due to the difference in antibody-antigen binding environment. In other words, alSF solution allows for a free diffusion of the antigen towards the antibodies immobilized on the electrode surface. This diffusion is limited in phantom gel because of its higher viscosity.

Overall, these results indicate the potential of the MNA-based immunosensor to be used as a point-of-care diagnostic tool for the extraction and quantification of bioanalytes including ErbB2 biomarker obviating the need for further sample treatment. In the literature, the detection of biomarkers from ISF has been performed using two different MN-based approaches that involve the extraction and quantification of the biomarkers. The first approach uses hollow MN to extract complete ISF whilst the second one employs antibody-modified MN to extract biomarkers only. Both methods require post extraction quantification of the biomarker. A disadvantage of the ISF extraction approach is the requirement of additional sample pre-treatment steps due to the biological complexity of the ISF, and currently the major drawback of the direct extraction of biomarkers using MN is the requirement for long and labor-intensive analysis techniques such as ELISA to quantify them. The Au-Si-MNA/anti-HER2 electrochemical immunosensor can provide efficient solutions for both problems mentioned above. High-density MNA not only serves as a method for extraction of biomarkers from the skin that is completed within 1 h, but also as an electrochemical transducer which simplifies the quantification of the biomarker providing response in only ~1 min and avoids the complexity of ELISA-like methods.

Conclusions - Here we present a high-density MNA-based electrochemical immunosensor for the detection of the ErbB2 breast cancer biomarker. For the first time, high-density Au-Si-MNA were simultaneously used as a biomarker extraction platform and in situ electrochemical transducer for the quantification of the captured biomarker. The size, shape, and conformation of MNA were characterized by means of SEM while the surface functionalization and the antibody immobilization were analysed using XPS and electrochemical methods. As a proof of concept, the MNA-based immunosensor was tested using a carefully designed phantom gel that mimics the epidermis and dermis layers from where ErbB2 biomarker was extracted and quantified via DPV. MNA-based immunosensors allow for the detection of ErbB2 biomarker over a wide concentration range from 10 to 250 ng/mL with a detection limit of 4.8 ng/mL in alSF, and 50 to 250 ng/mL and a detection limit of 25 ng/mL in the phantom gel, and analysis time of 60 min for extraction and electrochemical measurement of ~1 min. The analytical performance of the immunosensor is sufficient for detecting ErbB2 biomarker in ISF and tissue of breast cancer patients, and distinguish healthy and cancer patient cohorts. This MNA-based sensing platform is suitable for the modification with various biological recognition elements and thus can be applied for the sensing of different cancer biomarkers.

Example 4 - Polymeric Microneedle Arrays (pMNA) based sensors - Fabrication of pMNAand electrode preparation - Prior to fabrication of pMNA polydimethylsiloxane (PDMS) mould (Figure 22A) was prepared using Si-MNA previously fabricated for electrochemical glucose sensing as described above. Before pouring PDMS, Si-MNA were silanized overnight in gas phase using tricholoro (1 H, 1 H, 2H, 2H-perfluorooctyl)- silane in vacuum. After silanization, Si-MNA were covered with PDMS (10:1 , polymer: curing agent) and degassed for 1 h in order to remove entrapped air bubbles. Next, PDMS was cured in an oven at 60°C for 4 h. Then, cured PDMS was demolded from the Si-MNA substrate.

OrmoComp ^® (a UV curable silica-polymer hybrid) based MNA were fabricated using the PDMS mould. OrmoComp ^® was poured onto the PDMS mold and degassed in desiccator until all entrapped air in the polymer was removed. Afterwards, PDMS/OrmoComp ^® mould was exposed to UV light for 90 s and then PDMS mould was carefully removed from the crosslinked OrmoComp ^®

(Figure 22B).

SEM of fabricated pMNA - Morphological characterization (Figure 23) of pMNA was carried out by mans of scanning electron microscopy (SEM). Electrode preparation for electrochemical sensing was preformed using same protocol applied in Si-MNA-based glucose sensor described above. Detection of ascorbic acid in artificial interstitial fluid (alSF) using pMNA - The electroanalytical performance of pMNA was investigated in the detection of ascorbic acid (AA) with an electrode configuration consisting of pMNA-coated with sputtered nanolayer of Au and undercoat of Cr. The initial results of the calibration plot are shown in Figure 24A and B. Figure 24 shows differential pulse voltammetry (DPV) graph recorded in alSF in the presence of AA where current peak observed at -0.3 V increases with increasing concertation of AA. Calibration curve of pMNA-based sensor (Figure 24B) shows linear range from 100 mM to 400 mM of AA with sensitivity of 0.68 mA/mM. Figure 24C shows DPV plot of pMNA/Au electrode in the absence and presence of AA and uric acid (UA) in alSF. Current peaks for AA and UA were observed at -0.3 V and -0.5 V, respectively.

Detection of lactate using pMNA based biosensors - Sensor preparation - First approach: pMNA/Au/Fc-PAMAM/LOx sensor was prepared with exactly same approach used for glucose sensor preparation described above except that instead of silicon MNA, pMNA were used and instead of GOx enzyme Lactate Oxidase (LOx) is used. Second approach: pMNA/Au/PB- ChAuNPs/LOx biosensor was prepared first by electrodeposition of Prussian Blue (PB) layer on the pMNA/Au electrode after which chitosan and gold nanoparticles (AuNPs) were drop casted onto pMNA/Au/PB electrode. LOx was then immobilized onto pMNA/Au/PB-ChAuNPs surface using glutaraldehyde (2.5%).

Lactate Detection - Figure 25 shows amperometric response (Figure 25A) and calibration curve (Figure 25B) of the pMNA/Au/Fc-PAMAM/LOx biosensor recorded in 0.1 M PBS. Biosensor has linear range from 200 to 800 mM with sensitivity of 2.38 mA/mM. Results for pMNA/Au/PB- ChAuNPs/LOx are in Figure 26A and 26B. Biosensor has linear range from 200 to 1200 mM with sensitivity of 0.46 mA/mM.

Example 5 - Potentiometric pMNA-based pH sensor - Sensor preparation - pMNA-based pH sensor was fabricated by electro-polymerization of aniline onto MNA surface. Figure 27A shows the cyclic voltammogram (CV) obtained during electro-polymerization of polyaniline on pMNA-based electrode. Figure 27B shows electrode comparison of the 20 ^th cyclic voltammogram obtained when using five different electrodes illustrating good reproducibility of poly(aniline) film deposition on pMNA surface.

Optimization of electropolymerized poly(aniline) layer - The thickness of deposited polyaniline film was optimized by Potentiometry with the results shown in Figure 28A. The potential obtained with 20 and 30 cycles of polymerization are significantly higher (P < 0.05) than the other tested values. In order to decrease the electrode fabrication time, 20 cycles were selected despite of the insignificant potential response obtained with 30 cycle of polymerization. The effect of scan rate on the potentiometric response was also investigated. A comparison between the results shown in Figure 28B indicate that the maximum potential response on pH change can be obtained when using a scan rate of 75 mV/s (P < 0.05).

Analytical performance of pMNA based pH sensor - The pH range of artificial interstitial fluid (alSF) that the fabricated pMNA-based sensor can operate is shown in Figure 29A and B, indicating a linear range of pH 6.4 to 7.8. Testing of such short pH range is because normal change in ISF pH will be from 6.60 to 7.60. As seen from Figure 29B, the pH sensor provided a linear Nernstian response (sensitivity) of 67.2 mV/pH with high correlation coefficient (R) of 0.9992 across pH range of 6.4 - 7.8. An illustration of the application of the pMNA-based pH sensor for the continuous monitoring of pH in alSF is shown in Figure 30.

Overall summary - Si/Cr/Au/MNA/Fc-PAMAM/GAD/GOx modified electrode enabled direct glucose detection from ISF conveniently without need for extraction or collection, and then external analysis of the sample. The MNA glucose patch in alSF showed very good selectivity, with sensitivity of 0.1622 mA/ (mM cm ² ) in working range of 1 to 9 mM glucose, and detection limit of 0.66 mM. Successful in vivo application of the biosensor on mouse shows the MNA patch has the ability to detect changes in the ISF glucose levels which correlated very well with the blood glucose values obtained with commercial glucometer.

Similarly, TCPSi/ABA Fc-PAMAM/GAD/GOx modified electrode was found to have a limit of detection of 24 nM, a linear sensing range from 0.6 to 6 mM, a response time of 20 s, and a sensitivity of 1.33 mA mM ¹ cnr ² .

High-density silicon MNA (Si-MNA) was used to build a sensing platform for the simultaneous extraction and quantification of breast cancer biomarker ErbB2. The Si-MNA platform consists of -9500 MNs/cm ² , conformally coated with a 150 nm thick gold (Au) layer, incorporated into support in the form of a 3D printed holder to facilitate skin piercing MNA application and use for extraction and electrochemical quantification of ErbB2 biomarker. The Au-Si-MNA electrode was functionalized by self-assembling of 3-mercaptopropionic acid which carboxylic groups were activated via carbodiimide chemistry to covalently bind anti-HER2 antibody. The Si-MNA-based immunosensor was able to detect ErbB2 in artificial ISF (alSF) with a linear range from 10 to 250 ng/mL and a detection limit of 4.8 ng/mL which is sufficient to cover the ErbB2 concentration range in both healthy individuals and cancer patients. Furthermore, a skin-mimicking phantom gel simulating epidermis and dermis layers of the skin was used to assess penetration efficiency and ErbB2 extraction. The Si-MNA-based immunosensor showed excellent skin model penetration ability as well as promising analytical performance in extraction and electrochemical quantification of ErbB2 biomarker. This type of technology shows promise for minimally invasive transdermal diagnosis and monitoring of cancer biomarkers and thus has high potential in early cancer detection.

Given the above, the proposed MNA sensing patch provides an alternative transdermal diagnostic tool to the invasive existing techniques.

Embodiments Embodiment 1 : An electrochemical sensor system for sensing a bioanalyte in a media of interest, the sensor system comprises the following components: a reference (RE) electrode comprising a reference electrode substrate supporting an array of skin piercing solid microneedles at a density of at least 5,000 microneedle/cm ² disposed on the reference electrode substrate; a counter (CE) electrode comprising a counter electrode substrate supporting an array of skin piercing solid microneedles at a density of at least 5,000 microneedle/cm ² disposed on the counter electrode substrate; and a working (WE) electrode comprising a working electrode substrate supporting an array of skin piercing solid microneedles at a density of at least 5,000 microneedle/cm ² disposed on the working electrode substrate; wherein each microneedle in the array of microneedles is coated with a conformal nanofilm of at least one biocompatible electrically conducting, chemically inert material which supports a redox reaction, and wherein the solid microneedles arrays on each of the reference electrode substrate, the counter electrode substrate and the working electrode substrate are simultaneously contactable with a medium of interest which serves as electrolyte in the cell, and wherein each of the reference electrode substrate, the counter electrode substrate and the working electrode substrate are physically and electrically isolated from each other on the sensor system.

Embodiment 2: The sensor system of embodiment 1 , wherein the redox mediated dendrimer enzyme is specific for a bioanalyte selected from glucose, lactate, ethanol or urea.

Embodiment 3: The sensor system of embodiment 1 or embodiment 2, wherein the redox mediator is the focal point of the dendrimer.

Embodiment 4: The sensor system of any one of the preceding embodiments, wherein the redox mediator is retained between the conducting material and the enzyme by the dendrimer. Embodiment 5: The sensor system of any one of the preceding embodiments, wherein the dendrimer is immobilised onto the conformal nanofilm of biocompatible electrically conducting, chemically inert material.

Embodiment 6: The sensor system of any one of the preceding embodiments, wherein enzyme molecules are bound to the dendrimer in a single layer.

Embodiment 7: The sensor system of any one of the preceding embodiments, wherein the dendrimer is selected from hyperbranched polyglycerol dendrimers, polyethylene glycol dendrimers, glutamic acid dendrimers and PAMAM dendrimers.

Embodiment 8: The sensor system of any one of the preceding embodiments, wherein the dendrimer is a GO generation dendrimer or upwards, preferably, a PAMAM dendrimer.

Embodiment 9: The sensor system of any one of the preceding embodiments, wherein the dendrimer is a G2 PAMAM dendrimer. Embodiment 10: The sensor system of any one of the preceding embodiments, wherein the redox mediator is ferrocene.

Embodiment 11 : The sensor system of any one of the preceding embodiments, wherein the enzyme is GOx, the redox mediator is a ferrocene-dendrimer derivative, and the bioanalyte is glucose. Embodiment 12: The sensor system of any one of the preceding embodiments, wherein the sensitivity of the sensor system allows detection of mM variation in bioanalyte levels, particularly glucose levels.

Embodiment 13: The sensor system of any one of the preceding embodiments, wherein the redox mediated dendrimer enzyme is selective for glucose in a media comprising glucose and one or more of from uric acid (UA), ascorbic acid (AA), dopamine (DA), lactic acid (LA) and glycine (gly). Embodiment 14: The sensor system of any one of the preceding embodiments, wherein the microneedles of one or more of the electrodes are formed from silicon.

Embodiment 15: The sensor system of any one of the preceding embodiments, wherein the microneedles of one or more of the reference electrode substrate, the counter electrode substrate and working electrode substrate are coated with a sputtered nanolayer of an electrically conductive but chemically inert material, such as a conformal nanofilm of gold or platinum.

Embodiment 16: The sensor system of any one of the preceding embodiments, wherein the solid microneedles of one or more of the electrodes comprise an adhesive layer of chromium under a nanofilm of gold, wherein the chromium layer is thinner than the gold layer.

Embodiment 17: The sensor system of embodiment 16, wherein the dendrimer is bound to the surface of the gold layer and enzyme molecules are bound to the dendrimer.

Embodiment 18: An electrochemical sensor system for in vivo sensing of a bioanalyte in interstitial fluid, the sensor system comprising an electrochemical cell comprising the following components: a reference (RE) electrode comprising a reference electrode substrate supporting an array of solid, non-hollow silicon microneedles at an array density of at least 5,000 microneedle/cm ² disposed on the reference electrode substrate; a counter (CE) electrode comprising a counter electrode substrate supporting an array of solid, non-hollow silicon microneedles at an array density of at least 5,000 microneedle/cm ² disposed on the counter electrode substrate; and a working (WE) electrode comprising a working electrode substrate supporting an array of solid, non-hollow silicon microneedles at an array density of at least 5,000 microneedle/cm ² disposed on the working electrode substrate; wherein each microneedle of the array of microneedles is coated with a first adhesive layer of chromium and a second conformal nanofilm of gold over the chromium layer, and wherein surfaces of the material of the conformal nanofilm of the gold of the working electrode are modified with a ferrocene mediated G2 PAMAM dendrimer GOx enzyme specific for a glucose bioanalyte of interest, wherein the ferrocene is the focal point of the dendrimer and the ferrocene is retained between the gold and the GOx enzyme by the G2 PAMAM dendrimer; and wherein the solid microneedles arrays on each of the reference electrode substrate, the counter electrode substrate and the working electrode substrate are simultaneously contactable with in vivo interstitial fluid which serves as electrolyte in the cell, and wherein the reference electrode substrate, the counter electrode substrate and the working electrode substrate are physically and electrically isolated from each other on the sensor. Embodiment 19: The sensor system of embodiment any one of embodiments 1 to 14, wherein the at least a portion of the microneedles on one or more of the reference electrode substrate, the counter electrode substrate and working electrode substrate are solid microneedles having a carbonized porous surface layer, preferably having pore diameters of <100 nm.

Embodiment 20: The sensor system of embodiment 19, wherein the carbonized porous surface layer comprises a conformal nanofilm of carbon, preferably a thermally carbonised nanofilm.

Embodiment 21 : The sensor system of embodiment 20, wherein the carbonized porous surface layer is obtained by thermal treatment of the microarray needles at high temperature of from 400-800 °C in the presence of acetylene.

Embodiment 22: The sensor system of any one of embodiments 19 or 21 , wherein the dendrimer is immobilised into a diazonium salt grafted on the carbonised nanofilm and enzyme molecules are covalently bound to the dendrimer in a single layer.

Embodiment 23: An electrochemical sensor system for in vivo sensing of a bioanalyte in interstitial fluid, the sensor system comprising an electrochemical cell comprising the following components: a reference (RE) electrode comprising a reference electrode substrate supporting an array of solid silicon microneedles having porous surface layer thereon, at an array density of at least 2,000 microneedle/cm ² disposed on the reference electrode substrate; a counter (CE) electrode comprising a counter electrode substrate supporting an array of solid silicon microneedles having porous surface layer at an array density of at least 2,000 microneedle/cm ² disposed on the counter electrode substrate; and a working (WE) electrode comprising a working electrode substrate supporting an array of solid silicon microneedles having porous surface layer at an array density of at least 2,000 microneedle/cm ² disposed on the working electrode substrate; wherein each microneedle of the array of microneedles is coated with conformal nanofilm of thermally carbonised carbon having a porous surface layer or region having pore diameters of <100 nm, , wherein surfaces of the material of the conformal nanofilm of the working electrode are modified with a ferrocene mediated G2 PAMAM dendrimer GOx enzyme specific for a glucose bioanalyte of interest, wherein the ferrocene is the focal point of the dendrimer and the ferrocene is retained between the thermally carbonised carbon and the GOx enzyme by the G2 PAMAM dendrimer; and wherein the solid microneedles arrays on each of the reference electrode substrate, the counter electrode substrate and the working electrode substrate are simultaneously contactable with in vivo interstitial fluid which serves as electrolyte in the cell, and wherein the reference electrode substrate, the counter electrode substrate and the working electrode substrate are physically and electrically isolated from each other on the sensor. Embodiment 24: The sensor system of any one of the preceding embodiments, wherein the medium of interest is a calibration fluid, an in vitro test fluid, a biofluid, such as blood including peripheral blood, sweat or interstitial fluid, preferably in vivo interstitial fluid.

Embodiment 25: The sensor system of any one of embodiment 15 to 18, wherein the electrode is a solid non-porous silicon microneedle array (MNA) coated with a conformal gold nanolayer and exhibiting a limit of detection of 0.66 mM, a linear sensing range from 1 to 9 mM, and a sensitivity of 0.6122 pA mM- ¹ cnr ² .

Embodiment 26: The sensor system of any one of embodiments 19 to 23, wherein the electrode is a solid non-porous silicon coated with conformal nanofilm of thermally carbonized carbon having a porous surface layer having pore diameters of <100 nm and exhibiting limit of detection of 24 nM, a linear sensing range from 0.6 to 6 mM, and a sensitivity of 1 .33 mA mM ^-1 cnr ² .

Embodiment 27: The sensor system of any one of the preceding embodiments, in the form of a sensing patch, preferably a wearable sensing patch, preferably configured as a transdermal wearable bioanalyte sensor.

Embodiment 28: A wearable device for continuous bioanalyte monitoring comprising the sensor system of any one of embodiments 1 to 27.

Embodiment 29: Use of a sensor system of any one of embodiments 1 to 27 or the wearable device of embodiment 28 in the transdermal monitoring of a bioanalyte, preferably glucose, lactate, ethanol or urea.

Embodiment 30: Use of a microneedle array sensor system according to any one of embodiments 1 to 27 or the wearable device of embodiment 28 in transdermal bioanalyte monitoring. Embodiment 31 : A method of bioanalyte monitoring in a medium of interest comprising the steps of: providing an electrochemical sensor system according to any one of embodiments 1 to 27 to a test medium of interest such that electrode surfaces of the microneedles contact the medium of interest; applying a steady state electrical potential to the sensor’s counter electrode, whereby the potential provides a suitable potential window for detection of a redox reaction or chain of reactions initiated by the presence of the bioanalyte; monitoring for changes in an electrochemical property occurring at the sensor’s working (WE) electrode which result from a redox reaction involving the bioanalyte of interest; establishing the concentration of the bioanalyte present in the medium by comparing a level of a detected electrochemical property at the sensor’s working (WE) electrode with a predetermined calibration curve for the bioanalyte.

Embodiment 32: A method of transdermal bioanalyte monitoring comprising the steps of: providing an electrochemical sensor system according to any one of embodiments 1 to 27 to a test subject’s skin; aligning the sensor’s microneedle tips against the skin; applying a skin piercing pressure to the sensor to position the microneedles in skin’s dermal layer thereby allowing electrode surfaces of the microneedles to access the subject’s interstitial fluid; applying a steady state electrical potential to the sensor’s counter electrode, whereby the potential provides a suitable potential window for detection of a redox reaction or chain of reactions initiated by the presence of the bioanalyte; monitoring for changes in an electrochemical property occurring at the sensor’s working (WE) electrode which result from a redox reaction involving the bioanalyte of interest; establishing the concentration of the bioanalyte present in the interstitial fluid by comparing a detected level of the electrochemical property at the sensor’s working (WE) electrode with a predetermined calibration curve for the bioanalyte.

Embodiment 33: A method according to any one of embodiments 31 or embodiment 32, wherein the electrochemical property is one or more of current, potential, resistance, impedance and capacitance. Embodiment 34: A method of diagnosis or a disease or condition comprising the step of identifying the presence of a bioanalyte and/or a level of a bioanalyte associated with a disease or condition, using the sensor system according to any one of embodiments 1 to 27, preferably, wherein the bioanalyte is glucose and the disease or condition is insulin resistance, glucose intolerance, diabetes including Type 1 or Type 2 diabetes.