GB2539224A - Method of forming a chemical sensor device and device - Google Patents

Abstract

The present disclosure relates to a method for detecting simultaneously the pH and the concentration of chemical substances in a solution (e.g. a biological fluid). The chemical may be glucose, lactate, or carbon dioxide. The method utilises a multi-electrode sensing device 100 with two or more electrodes, wherein a first electrode is used for pH detection, and wherein further working electrodes are enzyme-coated electrodes for sensing the chemical (e.g. glucose or lactate), wherein the measurements from the enzyme electrodes are compensated using the measurements from the pH electrode. The enzyme electrode may comprise glucose dehydrogenase, lactate dehydrogenase, or carbonic anhydrase. The method is suitable for use both in stand-alone sensor devices, where the solution is deposited or conveyed by fluidic transport means to the sensing layers, and in implantable and skin patch-like sensor devices, where the device is applied for measuring the pH and the concentration of given substances in situ. The measurements may be poteniometric.

Description

METHOD OF FORMING A CHEMICAL SENSOR DEVICE AND DEVICE

BACKGROUND

Technical Field

The present disclosure generally relates to a method for detecting simultaneously the pH value of a solution, such as but not limited to a biological fluid, and the concentration of chemical substances in the solution, such as, but not limited to, glucose, lactate, carbon dioxide and the like, and the method to develop a device performing as self-calibrating, oxygen-independent, electrochemical sensor for such substances. The method is suitable for use both in stand-alone sensor devices, where the solution is deposited or conveyed by fluidic transport means to the sensing layers, and in implantable and skin patch-like sensor devices, where the device is applied for measuring the pH and the concentration of given substances in situ, and where the solution corresponds to one selected from the group consisting of the following: a water-containing epidermis's stratum corneum of skin, sweat, blood, cerebrospinal fluid, tears, plasma, saliva, or another biological fluid of the like.

More specifically, without limiting the use of other enzymatic chemical reactions, through the mentioned methods, in one embodiment of the present disclosure, a chemical sensor is disclosed which allows to measure independently the pH of the aforementioned solution and the concentration of chemical substances in the same solution, such as glucose and lactate, using at least two enzymes, such as lactose dehydrogenase (LDH) and glucose dehydrogenase (GDH), catalyzing the reaction of L-Lactate and Glucose in the presence of oxidized nicotinancide adenine dinucleotide (NAD+) as co-enzyme to produce respectively, pyruvate and gluconic acid together with reduced nicotinatnide adenine dinucleotide (NADH) and protons (H+), with no dependence on oxygen concentration in the solution.

In another embodiment of the present disclosure, another enzyme such as carbonic anhydrase, catalyzing the conversion of carbon dioxide (CO2) and water (H20) to bicarbonate (HCO3) and protons (H+), is used in combination with the same pH sensing method, with or without the use of other enzymes, as the ones mentioned in other embodiments, in order to detect such metabolic analytes in combination with the concentration of carbon dioxide, with no dependence on oxygen concentration in the solution.

In yet another embodiment, the pH level of the solution is measured amperometrically rather than potentiometrically, by means of suitable materials and device structures, also detailed in the present disclosure.

Moreover, the present disclosure teaches a method to form a sensor device of the mentioned substances, based on the measuring of the hydrogen ions formed during the enzymatic reactions, which is performed through a device with at least four terminals and corresponding electrodes, of which at least one is built to act as pH sensitive electrode, another one is built to perform as reference electrode for potentiometric or amperometric reading, at least one of the other electrodes is built to perform as working electrode to measure the concentration of a given substance, such as glucose, lactate, carbon dioxide, and the like, and another electrode of the device is built to perform either as counter-electrode, for amperometric reading of pH according to one embodiment, or as yet another working electrode to measure the concentration of another substance among the ones aforementioned and the like, according to another embodiment of the present disclosure.

The method disclosed is advantageous when compared to conventional systems, such as the ones involving chemical reactions based on oxidase enzymes, because the chemical reactions at the basis of the sensing mechanism, produce more acidic species, and hydrogen ions, which are responsible of a greater shift in the pH locally, in the working electrode areas, than known enzymatic systems based on oxidase (i.e. glucose oxidase and lactate oxidase).

Another advantage of the present disclosure is that the detection method for the chemical substances mentioned above does not involve oxygen in the catalytic reactions, which makes the sensor device able of operating in anaerobic environments.

Moreover, the present invention discloses the architecture of a device composed by a sensing part, for detecting the afore mentioned chemical substances in a solution, and a reference part, for detecting the pH level of the same solution, which allows to reduce the need for complex calibration of the device, and teaches a method for fabricating a self-calibrating architecture and using the same, and preferred embodiments, as examples of physical implementation, non-limiting other forms of physical implementation of the same method.

Furthermore, the present disclosure teaches a method for manufacturing sensor devices, in which at least two sensing layers are combined to perform a multi-analyte detection of target substances.

An application of devices thereof is envisaged for monitoring the concentration of at least one of the substances between glucose and lactate in a biological fluid, such as sweat, tears, saliva, blood, cerebrospinal fluid, and others. It is anticipated that such monitoring will ultimately allow to develop wearable non-invasive devices for continuous monitoring of the mentioned chemical substances, finally resulting in applications for the health care, sport and fitness of a human being, and for implantable devices for monitoring the aforementioned substances from various tissues and sites inside the body.

Description of the Related Art

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

Similar approaches have been developed for monitoring other metabolic analytes that are present in biological fluids and tissues of living organisms, such as, for example, lactate, glutamate, phosphate, creatinine, creatine, cysteine, and the like, where biological fluids include, for example, blood, tears, sweat, cerebrospinal fluid, plasma, and the like. The publication by Wilson, et al. "Biosensors for Real-Time in Vivo Measurements" in Biosensors and Bioelectronics (2005) pp. 1-16; Elsevier B.V. provides a comprehensive picture of the different approaches and state of the art at the time of publication.

For instance, the significance of continuous monitoring lactate concentration levels in biological fluids and/or in tissues is linked to the detection of a number of diseases and abnormal conditions associated with e.g. sepsis, kidney or liver disorders, cardio-pulmonary diseases and diabetes. Elevation of resting blood lactate concentration is not only associated with survival risk, but it can also be used as an indicator of the patient oxygen supply. Accordingly, rapid determination of lactate is particularly important in special care units.

Moreover, physiological lactate levels are related to the status of anaerobic metabolism associated with muscle contraction. Under resting conditions, healthy persons have lactate concentrations between 0.6 and 2 mM, but during strong physical activity the value of this parameter can rise up to 20 or 30 mM. Accordingly, lactate quantification is particularly important in sports medicine, since athletes have to stop their physical activity when they reach their lactate threshold. After this limit, the concentration of lactate rises exponentially and the athlete may have metabolic disorders and injured tissues.

More broadly, the quantification of L-lactate by amperometric hiosensors is gaining importance not only for sport and clinical purposes, but also for food and wine quality assessment.

Lactate oxidase (LOD) is widely used in lactate hiosensors because of its simple reaction and easy biosensor design configuration. LOD catalyzes the oxidation of lactate to pyruvate. In presence of dissolved oxygen, the enzyme can be re-oxidized, releasing hydrogen peroxide. This last product can he oxidized at the electrode surface restoring the former concentration of oxygen and giving a current proportional to the amount of dissolved lactate.

In some clinical situations, simultaneous monitoring of two or more analytes is desirable. Because of the complex interrelationship between glucose and other metabolic analytes it is often desirable to simultaneously detect glucose and lactate, possibly with some electrolytes such as for example potassium (K+), sodium (Na+) or calcium (Ca2) ions, as well as oxygen, carbon dioxide, and the like. Simultaneous monitoring of glucose, lactate and/or oxygen levels in the brain provides a comprehensive picture of complementary energy supply to the brain in response to acute neuronal activation. Levels of glucose and glutamate in Cerebral Spinal Fluid (CSF) are important in the control of diseases such as meningitis.

Dual electrode systems for continuous monitoring of a given analyte have been disclosed (US Pat. Appl. No. 2008/0083617A1 Apr 10, 2008), with the purpose of reducing the noise in amperometric measurement of the given analyte (e.g. glucose), which rely on red-ox electrochemical reactions that occur in presence of oxygen, with known limitations as described in the following paragraphs.

Other published works have disclosed methods and significance of combining enzymatic sensors for simultaneous monitoring of both glucose and lactate in diagnostics and therapeutics. See for instance the work of N.Sato, et al "Amperometric Simultaneous Sensing System for D-Glucose and L-Lactate Based on Enzyme-Modified Bilayer Electrodes" Analytica Chimica Acta (2006) pp. 250-254; vol. 565; Elsevier B.V.

More recently, F. Papadimitrakopoulos et al. (US 8,608,922 B2, Dec. 17, 2013) developed a method to manufacture multi-analyte enzymatic biosensors based on a three-electrode device: a Reference, a Counter and a Working electrode, respectively, where the latter is coated with either a sequentially covered plurality of enzyme-containing porous sections or permeability-adjusting spacers devoid of enzymes and functioning by controlling diffusion of various metabolites through adjacent enzyme-containing sections and/or inner and outer surfaces of the measuring device.

Currently, most of the electrochemical sensors used for the specific detection of lactate, glucose, and the like, employ analyte-specific enzymes, and are based on amperometric detection.

For example, first generation glucose sensors employ the glucose oxidase enzyme (GOD), immobilized on top of a Working electrode. This enzyme catalyses the oxidation of glucose to gluconolactone, as shown in reaction (1) below: Glucose + (GOD)", -> gluconolactone + (GOD)rea 02+ (GOD)rea -> (GOD)", + H202 (1) The generated hydrogen peroxide is then amperometrically assessed on the surface of a Working electrode according to reaction (2) below: H202 02 ± 2H+ + 2e- (2) Similar enzymatic reactions are adopted for monitoring the concentration of lactate, according to known methods in prior art, the scheme being based on the following chemical reactions, with lactate oxidase immobilized on top of the Working electrode: Lactate + (LOD)ox pyruvate + (LOD)red (3) 02+ (LOD) red (LOD)ox + H202 The generated hydrogen peroxide is then amperometrically assessed on the surface of a Working electrode with the same method as for Glucose, according to reaction (2) above.

In some other known methods, the reaction takes pace via a bi-enzymatic method, employing two enzymes both immobilized on the same working electrode. For instance this is the case of bi-enzymatic sensors for detection of creatine, which employ creatinase and sacrosine oxidase. First creatine is enzymatically converted by creatinase to sacrosine and urea, the former of which is subsequently converted to glycine and hydrogen peroxide by the action of sacrosine oxidase enzyme and with the addition of oxygen. Similar to glucose sensors, the generated hydrogen peroxide is amperometrically assessed on the surface of working electrode by relating the current to creatine concentration.

The main problem existing in methods in prior art described above arise by the need of having oxygen involved as part of the enzymatic reactions used for detecting the concentration of target analytes.

Consequently optimum sensor performance can only be attained when the ratio of the target analyte concentration (i.e. glucose) to oxygen is less than I. If this ratio is greater than or equal to 1, the lack of oxygen renders reactions oxygen-limited and this results in inaccurate readings by the hiosensor.

For instance, in the case of implantable sensors for continuous reading of glucose concentration in blood, a 0.18 mM oxygen concentration in the subcutaneous tissue is substantially lower than the 5.6 mM of physiological concentration of the target analyte (i.e. glucose/oxygen ratio of ca. -30). This leads to signal saturation at higher concentrations of the target analyte.

This limitation has been addressed by the use of diffusion limiting membranes that provide a greater permeability resistance to the larger sized molecules (i.e. glucose, and the like) as opposed to oxygen. As a result of this modification, semipermeable membranes based on NAFIONO, polyurethane, cellulose acetate, epoxy resins, polyether-polyether sulfone copolymer membranes, and layer by layer (LBL) assembled polyelectrolytes and/or multivalent cations have been extensively investigated in prior art.

However, in the use of semi permeable membranes it is desirable to have strict control over the thickness and uniformity of the outer membranes and this methodology comes at the expense of decreased sensitivity and increased sensor response time. Furthermore, the accumulation of exogenous reagents within these outer membranes (i.e., calcification, biofouling etc.) lead to sensor drift and therefore to their eventual failure. Finally, foreign body reaction after some time tends to lower the concentration of oxygen in the site of implant over time, again leading to signal saturation at higher concentration of the target analyte (i.e. glucose).

In another attempt to minimize the effects of oxygen concentration on the measurement of the analyte, further approaches based on a different set of chemical reactions, generally referred to as second-and third-generation hiosensors, employ redox mediators. However, in the case of mediators, their toxicity and hiocompatihility along with the possibility to leach out from the device to the surrounding tissue present a major problem, especially for application in implantable devices.

Also, known methods of the related prior art where the sensing mechanism is based on oxygen-dependent chemical reactions are generally affected by calibration needs and related challenges in order to improve their specificity to detect one or more of the target anal ytes, due to the presence of interferences with exogenous species that oxidize at the same potential as hydrogen peroxide.

For example, in voltages of about 0.6 to about 0.7 volts (V) many endogenous species such as bilirubin, creatinine, L-cystine, glycine, ascorbic acid (AA), acetaminophen (AP), uric acid (UA), and the like, also get oxidized (leading to an erroneous electrochemical signal according to known methods).

In order to increase confidence in sensing accuracy, anionic charged membranes (e.g., NAFION®, polyester sulfonic acid, cellulose acetate, and the like) have shown to exclude interferences from anionic species like ascorbic acid, uric acid, and so on, based on the principle of charge repulsion. These methods, however, inevi tably impede permeation of negatively charged analyte species (e.g., lactate, pyruvate, glutamate, and the like), and render their detection challenging. In addition, the large response time associated with the diffusion of analytes through these membranes require long equilibration times in order to attain steady state performance between the inner and outer membrane, which is an additional drawback.

Another approach to eliminate interference signals from endogenous species has been the use of inner, ultra-thin, electropolymerized films between the Working electrode and the enzyme layer. These films have been seen to partially screen analytes and analyte sensors from the interference agents. However, while these electropolymerized films minimize the contributions to signal from the endogenous species, they do not completely eliminate them.

In another approach, secondary enzymes (for example, ascorbate oxidase, which converts ascorbic acid to dehydroascorbate and water) have been incorporated in the outer membrane of the sensor to eliminate the particular species from reaching the electrode surface and contributing to the amperometric current. These secondary enzymes, do however, require oxygen and therefore have the potential of depleting the sensor of oxygen, which can negatively impact the operation of the primary enzyme.

Another major problem with the current state of the art enzymatic sensors for continuous monitoring of lactate according to reaction (3) above, the generated hydrogen peroxide is amperometrically assessed on the surface of a working electrode by applying a positive potential, as shown above in reaction (2). However, because of the application of positive bias onto the working electrode, the negatively charged pyruvate in reaction (3) tends to electrostatically adsorb on its surface leading to (i) taint the working electrode and subsequent loss of sensor sensitivity and (ii) inhibition of the reaction of lactate oxidase in reaction (3) with subsequent erroneous readings. To this end, higher applied potentials, double pulsed amperometry or pulsed amperometric detection have been the common strategies to renew the surface of the Working electrode even though such techniques are complex to be applied for miniaturized sensors and implantable sensors with driving electronics.

Because of its role to every metabolic activity of the body, the level of glucose is expected to vary following trauma, fever, exercise and/ or another physical activities. Implantable glucose sensors can be made more reliable only when one takes into consideration the local and physiological variations in various metabolites that are in relation to glucose. These metabolites include lactate, and carbon dioxide.

BRIEF SUMMARY

According to the disclosed embodiments, a method to form a multi-functional electrochemical sensor, which is able to detect the pH level of a solution and at least one of the substances between glucose, lactate, carbon-dioxide, and other metabolite and /or electrolytes of interest in low oxygen concentration, or anaerobic, conditions, and an electrochemical sensor are provided that are free from drawbacks of the known art.

In one embodiment of the disclosed method, a device is composed by a four-terminal unit, with four terminals connected by conductive tracks to electrodes, so-called first working (WEI), second working (WE2), third working (WE3) and reference electrode (RE), respectively, which are in contact with skin or in a solution, such as a biological fluid containing the chemical substance to be detected, with given concentration.

Several functional materials are suitable for being used to form the first working (WEI), second working (WE2), third working (WE3), and reference (RE) electrodes in the electrochemical sensor element, and the like, including semiconducting polymers, metals, metal nanoparticles, carbon nanotubes or graphene and other I D and 2D nanomaterials of the like. in one preferred embodiment the reference electrode (RE) is formed by a conductive layer coated by silver and silver cloride (Ag/AgC1), the first working electrode (WEI) is formed by a pH sensitive conductive layer coated with a membrane selective for protons, such as but not limited to p-xylene (with chemical formula C6H4(CH3)3) or its polymeric form known as parylene, for potentiometric reading, the second and third working electrodes (WE2, WE3) are formed by conducting surfaces coated with a pH sensitive layer, further coated with a membrane selective for protons and containing specific enzymes, the realization of membranes thereof being possible in form of a hydrogel, or sol gel, or polymer, in such a way to produce an excess of hydrogen ions, as effect of given enzymatic chemical reactions with chemical substances of interest, than the ones produced in the first working electrode (WE1) thereof, the latter being used to detect the base pH value of the solution, without being exposed to the effect of such enzymatic reactions for detecting the specific substances present in the solution.

In order to limit the diffusion in the solution of protons in excess produced in correspondence to any of the given working electrodes WE2 and WE3 to other unwanted electrodes, a semipermeable membrane can be adopted in another embodiment, according to known methods and materials in prior art, such as NAFION®, poly(urethane)-based coatings, solgel membranes such as (aminoethylaminomethyl)phenethyltrimethoxysilane or methyltrimethoxysilane from Sigma-Aldrich and the like, various hydrogels such as those based on Poly(Ethylene Glycol) (PEG), including polyethylene glycol diacrylate (PEG-DA) or polyethylene methyl diacrylate, those based on polyvinyl alcohol (PVA) such as polypyrrol-based PVA (PPy-PV A), those based on 2-hydroxyethyl methacrylate (HEM A) such as 2-hydroxyethyl methacrylate -diethylene glycol dimethacrylate (HEMA-DEGDMA) or 2-hydroxyethyl methacrylate -ethylene glycol dimethacrylate (HEMA-EGDMA) and its co-polymer derivatives such as Poly(2-hydroxyethyl methacrylate-co-methyl methacrylate) (PHEMA-co-PMMA), or those based on Hyarulonic Acid such as Methacrylate derivatized hyaluronic acid (Met-HA), and other ion-selective coatings of the like.

These electrode surfaces and device architectures will have a given morphology and chemical composition, as reported in Fig. 1-2, for one preferred embodiment, and show specific operating modes and electrical characteristics, as reported in Fig. 4, by a non-limiting example of other possible forms of realization, with reference to the simplified electrical circuit scheme in Fig. 3, which is used for reading out the potentiometric response of the sensor, and actuate a compensation of the signals generated by the enzymatic reactions occurring on working electrodes WE2 and WE3, when detecting specific substances from solution, such as glucose and lactate, with the signal detected by working electrode WEI, corresponding to the pH level of the solution.

Generally described, one or more embodiments of the present disclosure is directed to techniques for measuring simultaneously the pH of a solution and the concentration of two metabolic analytes, such as glucose and lactate, in the same solution, in absence of oxygen, for instance by means of the following enzymatic chemical reactions, in presence of nicotinamide adenine dinucleotide (NAT') catalyzed by either L-lactate dehydrogenase (LDH),

LDH

L -lactate + NAD+ pyruvate + NADH + H+ (4) or by Glucose dehydrogenase (GDH),

CHM

glucose + NAD+ -> Gluconic acid + NADH + (5) Also, generally described, another set of embodiments of the present disclosure is directed to techniques for measuring simultaneously the pH of a solution and the concentration of carbon dioxide dissolved in the same solution, by means of the following enzymatic chemical reaction, in presence of water and catalyzed by the carbonic anhydrase (CAM enzyme:

GAIT

CO2 + H2 0 -> HCO3 + H+ (6) To the end of implementing integrated multi-functional sensing devices as embodiments of the present disclosure according to one or more of the above chemical reactions, without limiting other forms of realization which those skilled in the art will obviously appreciate by a number of changes and variants made to the same embodiments without departure from the scope of the present disclosure, a device with at least four electrodes is developed, which is organized according to different designs and geometrical layout, such as the ones reported in Figure 1, Figure 2 and Figure 3 respectively, as examples of embodiments of potentiometric reading of the generated electrical signals, referred to the simplified electrical circuit scheme in Figure 4, or the ones reported in Figure 6, Figure 7 and Figure 8 respectively, as examples of embodiments of amperometric reading of the generated electrical signals, referred to the simplified electrical circuit scheme in Figure 9.

According to one aspect of the present disclosure a method is provided to simultaneously read a signal corresponding to the concentration of the given substances in the aforementioned solution and to actuate a compensation of the corresponding signal with respect to the pH fluctuations in the solution, as illustrated for example in figure 5.

In another aspect, according to the forms of implementation in Figure 6, Figure 7 and Figure 8, a device for amperometric, rather than potentiometric, reading in absence of oxygen in order to he free of limitations in prior art, where the electrode 7 operates as counter electrode (CE) and is connected together with reference electrode 9 (RE) to the electrical circuit as in the simplified scheme of figure 9, and first and second working electrodes (WEL WE2).

Other forms of implementation of the present disclosure that those skilled in the field will appreciate on the basis of the methods and device structures outlined in the present disclosure, for instance on the basis of Figures 1, 2 and 3, include device structures with a different number of working electrodes, corresponding terminals and coating layers, such as membranes embedding specific enzymes that catalyze one or more of the chemical reactions (4), (5) or (6), and the like, or with a different shape and physical arrangement, or other ways of combining the signals generated by the corresponding sensor device in order to actuate a compensation method of the signals corresponding to the concentration of given substances in solution to the variation of unwanted interferences, such as those generated by fluctuations in the pH level of the solution thereof, according to a method and architecture similar to the one illustrated for example in Figure 5.

It is also obvious for a person skilled in the art to modify the number of electrodes and corresponding terminals, their composition and layout, in the device structure illustrated in Figures 6, 7 and 8, for amperometric reading of the generated signals, and actuating the compensation method teach by the present disclosure to the variation of unwanted interferences, such as those generated by pH fluctuations in solution, in such a way to he advantageous with respect to known techniques in prior art, and to operate well in anaerobic conditions, as it may be required for use in wearable and implantable sensor devices.

BRIEF DESCRIPTION OF THE FIGURES

Further illustrative embodiments of the present disclosure are also defined in the appended claims and in the description, which is to be studied in combination with the figures, where like parts are numbered alike: Figure I schematically illustrates a planar 4-terminal device for potentiometric detection of multiple parameters characterizing a solution, such as the pH level of the solution and two more substances detected through an enzymatic reaction producing an excess of protons in the relevant electrodes. Figure la illustrates the top view while Figure lb shows a cross section of the same device.

Figure 2 schematically illustrates another form of implementation of a planar 4-terminal device for potentiometric detection of multiple parameters characterizing a solution, such as the pH level of the solution and two more substances detected through an enzymatic reaction producing an excess of protons in the relevant electrodes. Figure 2a illustrates the top view while Figure 2b shows a cross section of the same device.

Figure 3 schematically illustrates yet another form of implementation of a cylindrical 4-terminal device for potentiometric detection of multiple parameters characterizing a solution, such as the pH level of the solution and two more substances detected through an enzymatic reaction producing an excess of protons in the relevant electrodes. Figure 3a illustrates the 3D view while Figure 3b shows a cross section of the same device.

Figure 4 schematically illustrates a simplified electrical circuit for reading the potentiometric signals generated by a four terminal device as the ones illustrated in Figures I, 2 or 3. Figure 4a illustrates the simplified block scheme adopting 3 differential instrument amplifiers, while Figure 4b shows one of the possible design of the circuit implementing a single differential instrument amplifier.

Figure 5 illustrates an example of an electronic circuit used to acquire the signals generated by the reading circuit in Figure 4 and process them in order to actuate a compensation algorithm by means of a microcontroller. The figure shows by example the corresponding signals for a simultaneous reading of the glucose and lactate concentration of a solution and its pH level, which allows to obtain the compensated signals, also illustrated I the same figure, by filtering the data corresponding to the concentration of glucose and lactate from the signal corresponding to the pH level fluctuations in the solution thereof.

Figure 6 schematically illustrates a planar 4-terminal device for amperometric detection of multiple parameters characterizing a solution, such as the pH level of the solution and one more substance detected through an enzymatic reaction producing a current flow between the relevant working and current electrodes when the working electrode is maintained at a fixed potential than the reference electrode. Figure 6a illustrates the top view while Figure 6b shows a cross section of the same device.

Figure 7 schematically illustrates another form of implementation of a planar 4-terminal device for amperometric detection of multiple parameters characterizing a solution, such as the pH level of the solution and one more substance detected through an enzymatic reaction producing a current flow between the relevant working and current electrodes when the working electrode is maintained at a fixed potential than the reference electrode. Figure 7a illustrates the top view while Figure 7b shows a cross section of the same device.

Figure 8 schematically illustrates yet another form of implementation of a cylindrical 4-terminal device for amperometric detection of multiple parameters characterizing a solution, such as the pH level of the solution and one more substance detected through an enzymatic reaction producing a current flow between the relevant working and current electrodes when the working electrode is maintained at a fixed potential than the reference electrode. Figure Sa illustrates the 3D view while Figure 8h shows a cross section of the same device.

Figure 9 schematically illustrates a simplified electrical circuit for reading the amperometric signals generated by a four terminal device as the ones illustrated in Figures 6, 7 or 8.

DETAILED DESCRIPTION

The invention will now he described in more details with reference to the accompanying drawings, in which various embodiments are shown. This invention may, however, be embodied in many different forms, and should not be construed as limited to the embodiments set forth herein. Rather, these embodiments are provided so that this disclosure will be thorough and complete, and will fully convey the scope of the invention to those skilled in the art. Like reference numerals refer to like elements throughout.

Figure I a schematically illustrates a top view of a 4-terminal planar sensor device and the 4 electrodes of the same device. For better clarity, further illustration of the same device structure is represented in Figure lb with a view in cross-section of the device, outlining different layers of the same device.

According to a preferred embodiment of the present disclosure as schematically illustrated in Figure la, for a top view, and Figure lb, for a view in prospective, the sensor device 100 is obtained by fabricating four electrodes composed by sensitive layers 6, 7, 8 and 9 on top of patterned conductive areas 2, 3, 4 and 5 on the substrate 1, preferably made of a flexible film such as polyethylene nafthalate (PEN), without limiting the use of other substrate materials, including bio-compatible, and /or flexible, and /or conformable to biological tissues, and/or at least partly biodegradable film, and the like which a person skilled in the art can easily identify and use according to the methods disclosed by the present invention, all being considered obvious equivalent means to implement the same method. For instance the substrate layer 1 can equivalently being made of other materials such as polyester terephth al ate (PET), polyimmi de (PI), polydimethylsiloxane (PDMS), parylene, hyarulonic acid-based compounds or poly(l acti c-cogl ycolic acid) (PLGA) substrates, and the like.

Figure 2a schematically illustrates a top view of a 4-terminal planar sensor device and the 4 electrodes of the same device, according to another embodiment of the present disclosure. For better clarity, further illustration of the same device structure is represented in Figure 2h with a view in cross-section of the device, outlining different layers of the same device.

According to the forms of implementation in Figure I, Figure 2, a first set of terminals and the related patterned conductive areas 2, 3, 4 and 5 are first deposited and patterned on the substrate 1. These patterned conductive areas are preferably made of a metallic layer, such as Platinum or Gold having a thickness comprised between 50 nanometers and 10 micrometers, normally in combination with a thin adhesive layer, such as Titanium or Chromium, with thickness between 10 and 100 nanometers, which is interposed between said metallic layer and the substrate, in order to ensure the adhesion between the metallic conductive layers and the substrate itself.

These patterned conductive areas can he formed by a plurality of known methods of either subtractive or additive type. For instance according to known photolithographic technique, the following process steps can be implemented on the initial substrate layer I in order to form the patterned conductive areas 2, 3, 4, and 5 in figure 1 and 2: a) cleaning and activation of the surface of the substrate 1, for instance by an oxygen plasma treatment, b) deposition of a thin adhesive layer, such as Titanium of Chromium, followed by a deposition of a subsequent layer of Platinum or Gold on the substrate by sputtering or by thermal evaporation, optionally followed by a treatment of the surface with adhesion promoter, c) deposition of a positive photoresist material, for instance by spin coating, d) soft bake at temperatures compatible with the maximum temperatures allowed by the substrate material and exposure of the photoresist to a UV radiation source through a photolithographic mask, e) post-exposure bake at temperatures compatible with the maximum temperatures allowed by the substrate material and development of the photoresist layer in portion of areas that will have to be removed in order to form the aforementioned patterned conductive areas, f) dry etching, for instance by reactive ion etching technique, of the metallic and adhesive layers uncovered by the photoresist in order to obtain the patterned conductive areas 2, 3 and 4 or wet etching of the metallic and adhesive layers in areas uncovered by the photoresist by double immersion in given wet etch solutions specific for the given metallic and the thin adhesive layers, g) removal of the remaining photoresist layer (photoresist strip) from areas of conductive patterns 2, 3, 4 and 5.

Similarly, another subtractive technique, known as lift-off photolithographic technique, can be adopted to obtain the same result, which consists in the following process steps: a) cleaning and activation of the surface of the substrate I, for instance by an oxygen plasma treatment, optionally followed by a treatment of the surface with an adhesion promoter layer, b) deposition of a negative photoresist material, for instance by spin coating, c) soft bake at temperatures compatible with the maximum temperatures allowed by the substrate material and exposure of the photoresist to a UV radiation source through a photolithographic mask, d) post-exposure bake at temperatures compatible with the maximum temperatures allowed by the substrate material and development of the photoresist layer to remove it in portion of areas where the aforementioned patterned conductive areas will have to be formed, e) optional deposition of a thin adhesive layer, such as Titanium of Chromium, followed by a deposition of a subsequent layer of Platinum, Gold or Silver on the substrate for instance by evaporation techniques f) removal of the negative photoresist layer by wet etching with consequent removal of the conductive metal layer (and the adhesive layer) sitting on top of the photoresist layer, obtaining the patterned conductive areas.

Other techniques for obtaining the aforementioned patterned conductive areas are known both subtractive, such as for instance by (excimer) laser scribing, and additive, such as by contact and non-contact printing of metal inks from solution and curing either by baking at temperatures compatible with the maximum temperatures allowed by the substrate material or by UV, or IR exposure.

A person skilled in the art can easily identify other materials, deposition and patterning methods, for the same scope and use them according to the methods disclosed by the present invention, all being considered obvious equivalent means to implement the same method. For instance the conductive areas 2, 3, 4 and 5 can be made of one or more electrically conductive materials, such as Copper or Aluminium, in combination with thin adhesion layers between the metal tracks and the substrate or not, and /or different forms of carbon-based conductive materials, including graphene, carbon nanotubes, conductive composites, conductive polymers and others. Also in order to increase the thickness of conductive tracks a further step of plating, such as electroplating or electro-less plating, can be adopted in combination with the aforementioned process steps after metallization and patterning, or the same can he obtained with other known deposition methods such as printing, coating or transfer methods of conductive layered materials.

There are several possible materials for the realization of the pH sensitive layer 6 to form the devices as schematically reported in Figures 1, 2, 3, 6, 7, and 8. Known materials to act as pH sensing layers 6 are for example compounds from the groups of quinones, hydroquinones, polyphenols, antioxidants, in combination with pH affecting enzymes (penicillinase, urease, oxaloacetate, decarboxylase), polymers such as parylene C, Poly(3,4-ethylenedioxythiophene) (PEDOT) or its poly-sulphonate form PEDOT:PSS, Polyaniline, copolymers such as 2-hydravethyl Inethacrylate -dunethylaminoethyl methatrylate (HEMA-co-DMAEMA) (9:1), and the like. Other materials suitable to perform as pH sensing layers 6, according to other forms of realization of the present disclosure, include the combination of metallic electrode layer 2 with its oxide, such as Pt/Pt02, W/W203, Pb/Pb02, Ir/Ir02, and the like, or carbon nanotubes, graphene, and other 1D and 2D nanomaterials of the like.

There are also several process available for manufacturing the realization pH sensitive layer 6 as a coating layer to the corresponding electrodes. For instance, according to one preferred embodiment, the sensing layer 6 is composed by a thin layer of parylene C, a pH sensitive biocompatible material, which is the polymer form of p-x ylene, deposited by vapor process in form of conformal coating on the conductive track of electrodes 2, 3 and 4, as schematically represented in Figures 1, 2 and 3, or on the conductive track of electrodes 2 and 4, as schematically reported in Figures 6, 7, and 8, with or without masking processes, and optionally after the corresponding electrode surfaces are pre-activated by oxygen plasma or similar treatments in order to improve the adhesion of the coating layer 6 and correspondingly the sensitivity of such electrodes to hydrogen ions.

According to another preferred embodiment, a layer of PEDOT can he formed by electrodeposition methods consisting of immersion of the electrically conductive wires in a electrolytic solution containing the precursor molecule EDOT 0.1M with LiC1O4 0.01M in Acetonitrile within a standard electrochemical cell and performing several cycles of cyclic voltammetry, between the electrodes 2, 3, and 4 of devices 100, 101 or 102 or electrodes 2 and 4 of devices 200, 201, or 202 and other counter and reference electrodes of the electrochemical cell. More in details, since the EDOT has a nominal electrical potential of I.13eV, a person skilled in the technique, will appreciate the cyclic voltammetry to he done by sweeping the voltage forth and hack, from a value lower than 1.13V, for example 0.8V, to a value higher than 1.13V, for example 1.5V, at a given rate between 10 and 100 mV/s and doing it for a number of cycles, typically between 2 and 10, sufficient to form a thickness of the PEDOT layer, from 100 nm to 500 nm, or more. Other polymeric coating can be formed with the same method or by other methods consisting for instance of immersing the electrodes 2, 3, and 4 of devices 100, 101 or 102 or electrodes 2 and 4 of devices 200, 201, or 202 into a solution of the targeted polymer or polymer precursors and promoting polymerization and cross-linking in the surface of the wire by either temperature, UV exposure, or potentiostatically by application of a continuous electrical field, having a voltage VDc between 1 and 2V to bias the cell, and a superimposed alternating electrical field of smaller value VAC of about 0.1V.

Alternatively the active pH sensing layer 6 can be formed on the surface of the corresponding electrodes, by thermal evaporation (TE), physical vapor deposition (PVD) and the like, by exposing the electrode surface through a shadow mask, when forming the planar devices 100, 101, 200, or 201, or by exposing the metal wire electrodes without masks, to form the cylindrical devices 102 or 202. The latter method is suitable to form for instance layers of Pt/Pt02, W/W203, Pb/Pb02, Ir/Ir02, and the like as well to deposit layers of organic materials such as parylene C, acting as both encapsulant and pH sensing layer.

Figure 3a schematically illustrates a 4-terminal cylindrical sensor device and the 4 electrodes of the same device, according to another embodiment of the present disclosure. For better clarity, further illustration of the same device structure is represented in Figure 3h with another cross-section of the device.

As opposite to the manufacturing process flows indicated for the fabrication of devices according to Figures 1 and 2, and according to Figures 4 and 5, the manufacturing of device according to Figure 3, and 8 is preferably obtained starting from conductive wires, for instance Platinum, Gold or Silver wires forming the electrodes 2, 3, 4 and 5, that are coated conformably with various functional layers.

Specifically, according to preferred embodiments of device 102 in Figure 3 and device 202 in Figure 8, the cylindrical wires corresponding to electrodes 2, 3 and 4 of device 102, and the cylindrical wires corresponding to electrodes 2 and 4 of device 202 are first coated with a pH sensing layer 6, using known pH sensing materials, by the aforementioned formation methods.

In figures 1, 2, and 3, the electrodes 7, 8, and 9 are formed by coating the surface of the electrodes with a given functional membrane material, in form of hydrogel or sol-gel of given formulation and containing the specific enzyme to govern the reaction to take place at the corresponding working electrode.

For instance, according to the embodiment of a multianalyte potentiometric sensor for monitoring pH, glucose, lactate and CO2 of a solution, the electrode 7, 8 and 9 of figure I to 3 are coated with ion-sensitive membranes in form of either polymer matrix, such as poly(urethane)based coatings, sol-gel membranes such as (aminoethylaminomethyl)phenethyltrimethoxysilane or methyltrimethoxysilane from Sigma-Aldrich and the like, or hydrogels such as those based on Poly(Ethylene Glycol) (PEG), including polyethylene glycol diacrylate (PEG-DA) or polyethylene methyl diacrylate, those based on polyvinyl alcohol (PVA) such as polypyrrol-based PVA (PPyPVA), those based on 2-hydroxyethyl methacrylate (HEMA) such as 2-hydroxyethyl methacrylate -diethylene glycol dimethacrylate (HEMA-DEGDMA) or 2-hydroxyethyl methacrylate -ethylene glycol dimethacrylate (HEMA-EGDMA) and its co-polymer derivatives such as Poly(2-hydroxyethyl methacrylate-co-methyl methacrylate) (PHEMA-co-PMMA), or those based on Hyarulonic Acid such as Methacrylate derivatized hyaluronic acid (Met-HA), and the like.

These hydrogel membranes are formed in such a way that they contain the molecules and enzymes specific to promote the enzymatic reactions, such as nicotinamide adenine dinucleotide (NAD+) and L-lactate dehydrogenase (LDH), or Glucose clehydrogenase (GDH), according to reactions (4) and (5) for monitoring the concentration of lactate and glucose, respectively, or Carbonic anhydrase (CAH), according to reaction (6) for monitoring the concentration of carbon di-oxide in solution.

For instance, according to a possible method of formation of a hydrogel membrane this can be done by: a) preparing a pre-polymer solution with 2-HEMA and DMAEMA having a molar ratio of 95/5, and solvent containing ethylene glycol and water in equimolar amounts, adding to the monomers a cross-linker such as TEGDMA, and optionally a photoinitiator, such as DMPAP to further promotes the cross-linking of polymer chains by UV exposure, the mole ratio between the monomers and the solvent being in the range of 1 to 1.2; b) developing the pre-polymer in DI water in order to remove the unpolymerized chains; c) adding to the hydrogel a solution containing the specific enzyme (LDH, GDH, or CAH) in PBS buffer, d) stirring it at low temperatures, around 4°C; e) using the hydrogel locally to further coat the electrodes 3, 4 and 5, pre-coated with the pH sensing layer, for instance by printing, or by any other additive or subtractive deposition and patterning technique, on top of the corresponding electrodes in devices of figure I and 2, or by immersing the given electrically conducting wire in the hydrogel, for the device of figure 3, and t) promoting further cross-linking for conformal coating by further UV exposure and for by temperature increase, and repeating the above steps until obtaining the specific coating layer membranes for sensing the specific analytes via the enzymatic reaction given by (4), (5) and (6), by potentiometric sensing according to the scheme of figure 4.

Also, there are several possible materials and methods to foiiii the reference electrode composed by the layer 9 on conductive electrode 5, in all the devices 100, 101, 102, 200, 201, and 202. According to a preferred embodiment this layer 9 is composed by a silver electrode or a silver coating layer on top of a conductive electrode layer 5, which is partially converted to silver chloride to form Ag/AgC1, for use as reference electrode with a fixed potential in electrochemical sensing, according to the electrical circuit schemes of Figures 4 and 9. For instance this can be obtained by soaking the silver or silver coated electrode in a solution of concentrated hydrogen chloride or sodium chloride for few seconds, between 2 and 10 seconds, until the surface of silver is chloridi zed to form a Ag/AgCI reference electrode, and checking its potential by cyclic voltammetry.

Similar manufacturing processes are adopted to develop the devices for amperometric sensing, according to the embodiments corresponding to figures 6, 7 and 8.

However, according to the principle of operation of amperometric reading of electrochemical enzymatic sensors, figures 6, 7 and 8 refer to devices implementing other embodiments of the present disclosure characterized by the fact that they require one of the electrodes acting as counter electrode.

For instance, without limiting the number of additional electrodes, their coating layer materials and use, in a 4-terminal device as in the example of devices 200, 201, and 202 of Figures 6, 7 and 8, there are only two electrodes accessible as working electrodes, for sensing pH (electrode 2, with layer 6) and another analyte (electrode 4, with layers 6 and 8), respectively, while the electrode 5 with layer 9 is used as reference electrode and the electrode 3, with layer 7, is used as counter electrode.

In one preferred embodiment the layer 7 of the counter electrode is omitted or made of the same material of electrode layer 3, for instance gold, by electroplating technique, and act as counter electrode for the corresponding devices 200, 201 and 202, according to the amperometric reading scheme of Figure 9.

Also, according to the preferred embodiments of devices 102 and 202 in Figures 3 and 8, the cylindrical wires, coated with the above mentioned materials and techniques, are mechanically assembled within a cylindrical rigid or flexible hollow needle 10 of a given diameter by means of an inert polymeric material 11 which is preferably selected among the materials that are permeable or semipermeable to both the hydrogen ions and chemical substances present in the solution, formed for instance by poly(dimethylsiloxane) (PDMS), and directly exposed in the terminal end of the wires to the solution.

Figure 4a outlines an equivalent circuit scheme showing a preferred embodiment of the present disclosure for the potentiometric reading of the multianalyte sensor unit, as referred for example to the aforementioned sensor devices 100, 101 or 102, in previous figures. The circuit is designed in order to operate a simultaneous detection of the potentiometric voltage signal detected between each of the working electrodes 2, 3 and 4 of the sensor device 100, 101 or 102, and the reference electrode 5 of the same device, according to the functioning principles of the sensor device resulting from the electrochemical and enzymatic reactions explained in the present invention disclosure.

For instance, when the sensor device is fabricated with materials such that the electrode 2 and its coating layer 6 is functioning as working electrode for pH detection, the electrode 3 and its coating layers 6 and 7 is functioning as working electrode for detection of glucose on the basis of the reaction (5), the electrode 4 and its coating layers 6 and 8 is functioning as working electrode for detection of lactate on the basis of the reaction (4), and the electrode 5 and layer 9 is functioning as reference electrode according to a potentiometric readout scheme, the interface circuit 13 is designed so that he provides as output three voltage signals proportional to: a) the pH level of the solution the sensor device is in contact with (output signal OUTI), b) the concentration level of glucose in the solution (output signal OUT2) and c) the concentration level of lactate in the same solution (output signal OUTS), respectively, not yet compensated against the variations of the pH level occurring during the 2 enzymatic reactions (4) and (5), or due to any undesired change of the pH level in the solution.

Similarly by replacing one of the layers 7 or 8 with a layer containing the carbonic anhydrase enzyme according to the reaction (6), by means of the materials and methods explained before, it is possible to measure the concentration of carbon dioxide in the same solution, yet uncompensated against the variations of pH in the solution.

More specifically, according to this preferred embodiment, the electronic sensor interface circuit 13 is composed by 3 instrumentation amplifiers 14, 15, 16, each one formed by a circuit as shown by an example non limiting other forms of realizations, circuit schemes and use of electronic components, in Figure 4b, by means of 3 operational amplifiers 27, 28 and 29 and 7 resistors.

It should be appreciated that the four 10k52 resistors indicated in the schematic of figure 4b are precision matched resistors obtained by internal laser trimming of thin film transistors and monolithically integrated with the output difference amplifier 29 for a good common-mode rejection ratio (CMRR), according to known circuit design techniques.

Finally, the method outlined in Figure 4a is not limited in the number of input and output signals and can easily be extended by adding one or more instrumentation amplifiers for potentiometric reading of additional sensor electrodes and generating corresponding output signals, uncompensated against the pH variations.

Figure 5 outlines a simplified scheme of the method as preferred embodiment of the present disclosure, which is adopted in order to compensate the above mentioned variations of pH and correct the measurements of the multiple analytes detected by the sensor devices and processed by the circuit interface 13 accordingly. More specifically the circuit scheme in figure 5 shows a possible way of implementing a compensation algorithm consisting of comparing the signals obtained by processing the potentiometric levels detected on working electrodes 3 (WE2, glucose) and 4 (WE3, e.g. lactate) with the signal obtained by processing the potentiometric level detected on electrode 2 (WEI, i.e. pH).

It can be appreciated that the corresponding potentiometric levels at electrodes 3 (WE2 e.g. glucose) and 4 (WE3, e.g. lactate), are the result of hydrogen ions generated from reactions (5) and (4), respectively, by means of the specific enzyme containing membrane layers 7 and 8, and detected by the same pH sensing mechanism occurring in the electrode 2 (WED) where the specific enzyme containing membrane layers 7 and 8 are missing.

A simple compensation is therefore implemented in the case of potentiometric reading according to the following algorithm: OUT2 compensated (e.g. glucose) = OUT2 (e.g. glucose) -OUTI (pH)(7) OUT3 compensated (e.g. lactate) = OUT2 (e.g. lactate) -OUTI (pH) (8) According to the circuit in Figure 5, and without limiting other forms of realization of the same method according to the present disclosure, the voltage signals generated by the interface circuit block 13, and corresponding to the data plots 19, 20 and 21, are converted in digital data by means of known methods and circuit blocks such as analog to digital converters (ADC) and analog or digital multiplexers (MUX), circuit block 17, and processed by a microcontroller unit (MCU), circuit block 18, to generate the compensated signals, corresponding to the data plots 22 and 23.

Figure 9 includes a simplified schematic of a preferred embodiment of the sensor interface circuit 30 used to read the devices 200, 201, or 202, which is composed by one operational amplifier 31 having the role of controlling the voltage at the counter electrode and keeping it at a constant bias voltage \TIN BIAS with respect to the potential of the reference electrode 5, and two instrumentation amplifiers high input impedance and high current sensitivity 32, detecting and amplifying the current signal generated by assessing amperometrically the hydrogen ions produced through the pH sensing layer 6 in the surface of the electrode 2 and 33, detecting and amplifying the current signal generated by assessing amperometrically the hydrogen ions produced through the layers 6 and the enzymatic membrane layer 8 on electrode 4.

The various embodiments described above can be combined to provide further embodiments. These and other changes can be made to the embodiments in light of the above-detailed description. In general, in the following claims, the terms used should not be construed to limit the claims to the specific embodiments disclosed in the specification and the claims, but should he construed to include all possible embodiments along with the full scope of equivalents to which such claims are entitled. Accordingly, the claims are not limited by the disclosure.

Claims (30)

1. CLAIMS: 1. A method for detecting simultaneously the pH value of a solution and the concentration of one or more chemical substances in the same solution, suitable for operating in anaerobic conditions, comprising: forming a multi-electrode sensing device with two or more electrodes, said first working electrode, second, and further working electrodes, formed by electrically conductive materials coated with a pH sensing layer producing a potential difference referred to another electrode of the same device, said reference electrode, which corresponds to a potentiometric measure of hydrogen ions concentration onto the surface of said first, second, and further working electrodes; further coating said second, and further working electrodes with a membrane layer containing a specific enzyme catalyzing a reaction of the chemical substance of interest in the solution in order to produce an excess of hydrogen ions that are assessed potentiometrically in the surface of said second, and further working electrodes by measuring the potential difference of the same electrodes referred to said reference electrode; and determining a compensated measure of the concentration of said chemical substances involved in said enzymatic reaction by filtering the corresponding signal generated across said second, and further working electrodes with respect to the signal generated by said first working electrode.
2. 2. The method of claim 1, wherein said first working electrode is formed by electrically conductive materials coated with a pH sensing layer producing a potential difference referred to said reference electrode, which corresponds to a potentiometric measure of hydrogen ions concentration onto said first working electrode; and wherein said second, and further working electrodes are formed by electrically conductive materials coated by a ion sensitive membrane layer containing a specific enzyme catalyzing a reaction of the chemical substance of interest in the solution in order to produce an excess of hydrogen ions that are assessed potentiometrically in the surface of said second, and further working electrodes by measuring the potential difference referred to said reference electrode; and determining a compensated measure of the concentration of said chemical substances involved in said enzymatic reaction by filtering the corresponding signal generated across said second, and further working electrodes with respect to the signal generated by said first working electrode.
3. 3. The methods of claims 1 and 2, wherein said first working electrode is foamed by electrically conductive materials coated with a p1-1 sensing layer selected among the ones suitable for potentiometric reading, such as several polymeric compounds as PEDOT, parylene, polyanyline, and the like, or co-polymers, such as (HEMA-co-DMAEMA), or metal oxides such as Pt/Pt02, W/W203, Pb/Pb02, 1r/1102, and the like, or carbon nanotubes, graphene, and other 1D and 2D nanomaterials of the like.
4. 4. The methods of claims from 1 to 3 wherein said reference electrode is formed by an electrically conductive material coated with a material having known and stable potential for electrochemical sensing, such as Ag/AgCl.
5. 5. The methods of claim 4 wherein said reference electrode is formed by soaking a silver electrode in a solution containing concentrated hydrogen chloride or sodium chloride for few seconds and checking its potential by cyclic voltammetry.
6. 6. The methods of claims from I to 3 wherein said reference electrode is formed by an electrically conductive noble material having known and almost stable or predictable potential for electrochemical sensing, such as Platinum, Gold or Silver.
7. 7. The methods according to any of the previous claims from I to 6 wherein said first, second and further working electrodes, and said reference electrode are formed on conductive areas on a substrate that are deposited, preferably after treatment of the substrate to clean and activate its surface, for instance by oxygen plasma, and deposition of a thin metal adhesion layer, for instance Chromium or Titanium and the like, and patterned according to known methods, such as subtractive methods, namely photo-lithography and the like, or additive methods, namely printing and the like, and wherein said first, second and further working electrodes and said reference electrode are formed by coating said patterned conductive areas on the substrate with given functional layers, such as pH sensing layers, enzyme-containing membranes and reference electrode layer, deposited and patterned by either subtractive, namely photolithography and the like, or additive methods, namely printing, electro-deposition methods and the like.
8. 8. The methods of claims from 1 to 6 wherein said first, second and further working electrodes, and said reference electrode are formed by electrically conductive wires of given materials, e.g. Platinum, Silver, Gold and the like; and wherein the electrically conductive wires of said first, second and further working and reference electrodes are coated with given functional layers, such as pH sensing layers, enzyme-containing membranes and reference electrode layer, deposited by either conformal coating techniques, such as thermal evaporation, physical or chemical vapor deposition, electro-deposition or electro-plating, soaking in an electrolyte solution; and wherein the electrically conductive wires coated with the above mentioned materials and techniques are mechanically assembled within a cylindrical rigid or flexible hollow needle of a given diameter by means of an inert polymeric material which is preferably selected among the materials that are permeable or semipermeable to both the hydrogen ions and chemical substances present in the solution, formed for instance by poly(dimethylsiloxane) (PDMS), and directly exposed in the terminal end of the wires to the solution.
9. 9. The methods according to the claim 8 wherein one or more electrodes are formed by uncoated wires of a given material such as gold, platinum and the like, carbon, and the like.
10. I O. The methods to read the signals generated by the multi -analyte sensing devices formed according to any of the methods indicated in claims from I to 9, wherein a set of instrumentation amplifiers are connected between said working electrodes and said reference electrode of said multi-analyte sensing devices in such a way to generate a number of analogue signals, namely voltage signals, that correspond respectively to: the pH level of a solution, assessed potentiometrically onto the surface of said first working electrode; the hydrogen ions generated by a first enzymatic reaction and potentiometrically assessed onto the surface of one of said working electrodes after the first one, for instance measuring the glucose level of the solution when the enzyme-containing membrane is made for being selective to glucose according to an oxygen-independent chemical reaction, such as in the reaction (5), where glucose reacts with nicotinamide adenine dinucleotide (NAD+) in presence of glucose dehydrogenase (GDH), producing gluconic acid, nicotinamide adenine dinucleotide in its reduced form (NADH) and hydrogen ion (Hi); and /or the hydrogen ions generated by a second enzymatic reaction and potentiometrically assessed onto the surface of one of said working electrodes after the first one, for instance measuring the lactate level of the solution when the enzyme-containing membrane is made for being selective to lactate according to an oxygen-independent chemical reaction, such as in the reaction (4), where lactate reacts with nicotinamide adenine dinucleotide (NAD+) in presence of L-lactate dehydrogenase (LDH), producing pyruvate, nicotinamide adenine dinucleotide in its reduced form (NADH) and hydrogen ion (Hi).
11. I I. The methods of cl aim 10 further compri sing an instrumentation amplifier connected between a working electrode and said reference electrode of said multi-analyte sensing devices in such a way to generate an analogue signal, namely a voltage signal, that corresponds to the hydrogen ions generated by an enzymatic reaction and potentiometrically assessed onto the surface of one of the working electrodes after the first one, for measuring the level of carbon dioxide molecules in the solution when the enzyme-containing membrane is made for being selective to carbon dioxide according to an enzymatic chemical reaction, such as in the reaction (6), where carbon dioxide (CO2) reacts with water (H2O) in presence of the carbonic anhydrase (CAH) enzyme, producing hydrogen carbonate ion (HCO3) and hydrogen ions (H+).
12. 12. The methods of claims 10 and 11 wherein the output signals generated by the instrumentation amplifiers, and corresponding to the readout levels of pH and two or more chemical substances dissolved in the solution where the multi-analyte sensor device is used, are further elaborated in order to implement a given compensation algorithm, to eliminate the dependence of the read-out signals corresponding to the concentration of target chemical substances in the solution from the pH level of the same solution, which is assessed in the surface of the first working electrode.
13. 13. The method of claim 12 wherein said compensation algorithm is implemented by means of a code within a programmable microcontroller device and wherein the uncompensated signals generated by the set of instrumentation amplifiers are introduced in the microcontroller by means of an analog to digital converter and multiplexer circuit block and data processed to obtain the compensated signals, after filtering the variations of pH in the solution according to the signal generated by the first working electrode and preprocessed by the corresponding instrumentation amplifier.
14. 14. A method for detecting simultaneously the pH value of a solution and the concentration of one or more chemical substances in the same solution, suitable for operating in anaerobic conditions, comprising: forming a multi-electrode sensing device with at least one reference electrode, one counter electrode and two or more electrodes, said first, second and further working electrodes, said working electrodes being formed by electrically conductive materials coated with a pH sensing layer, the second and further working electrodes being further coated with enzyme-containing membranes for specific analyte sensing in anaerobic conditions; assessing the concentration of hydrogen ions in the solution by measuring the current flow from said first working electrode, when a fixed bias potential difference is applied between a counter electrode and a reference electrode, according to known amperometric measurement techniques; assessing the concentration of other chemical substances in the solution according to the specific enzymatic reaction catalyzed by said enzyme-containing membranes, by measuring the current flow from said second and further working electrodes, when a fixed bias potential difference is applied between a counter electrode and a reference electrode, according to known amperometric measurement techniques of the excess hydrogen ions produced onto the surface of said second and further working electrodes; and determining a compensated measure of the concentration of said chemical substances involved in said enzymatic reaction by filtering the corresponding signal generated across said second, and further working electrodes with respect to the signal generated by said first working electrode.
15. 15. A method for detecting simultaneously the pH value of a solution and the concentration of one or more chemical substances in the same solution, suitable for operating in anaerobic conditions, comprising: forming a multi-electrode sensing device with at least one reference electrode, one counter electrode and two or more electrodes, said first, second and further working electrodes, said first working electrode being formed by an electrically conductive material coated with a pH sensing layer, the second and further working electrodes being coated with ion-sensitive enzyme-containing membranes for specific analyte sensing in anaerobic conditions; assessing the concentration of hydrogen ions in the solution by measuring the current flow from said first working electrode, when a fixed bias potential difference is applied between a counter electrode and a reference electrode, according to known amperometric measurement techniques; assessing the concentration of other chemical substances in the solution according to the specific enzymatic reaction catalyzed by said ion-sensitive enzyme-containing membranes, by measuring the current flow from said second and further working electrodes, when a fixed bias potential difference is applied between a counter electrode and a reference electrode, according to known amperometric measurement techniques of the excess hydrogen ions produced onto the surface of said second and further working electrodes; and determining a compensated measure of the concentration of said chemical substances involved in said enzymatic reaction by filtering the corresponding signal generated across said second, and further working electrodes with respect to the signal generated by said first working electrode.
16. 16. The methods of claims 14 and 15, wherein said first working electrode is formed by electrically conductive materials coated with a pH sensing layer selected among the ones suitable for amperometric reading, such as several polymeric compounds as PEDOT, parylene, polyanyline, and the like, or co-polymers, such as (HEMA-co-DMAEMA), or metal oxides such as Pt/Pt02, W/W203, Pb/Pb02, Ir/Ir02, and the like, or carbon nanotubes, graphene, and other 1D and 2D nanomaterials of the like.
17. 17. The methods of claims from 14 to 16 wherein said reference electrode is formed by an electrically conductive material coated with a material having known and stable potential for electrochemical sensing, such as Ag/AgCI.
18. 18. The methods of claim 17 wherein said reference electrode is formed by soaking a silver electrode in a solution containing concentrated hydrogen chloride or sodium chloride for few seconds and checking its potential by cyclic voltammetry.
19. 19. The methods of claims from 14 to 16 wherein said reference electrode is formed by an electrically conductive noble material having known and almost stable or predictable potential for electrochemical sensing, such as Platinum, Gold or Silver.
20. 20. The methods according to claims from 14 to 19 wherein said first, second and further working electrodes, and said reference and counter electrodes are formed on conductive areas on a substrate that are deposited, preferably after treatment of the substrate to clean and activate its surface, for instance by oxygen plasma, and deposition of a thin metal adhesion layer, for instance Chromium or Titanium and the like, and patterned according to known methods, such as subtractive methods, namely photo-lithography and the like, or additive methods, namely printing and the like, and wherein said first, second and further working electrodes and said reference electrode are formed by coating said patterned conductive areas on the substrate with given functional layers, such as pH sensing layers, enzyme-containing membranes and reference electrode layer, deposited and patterned by either subtractive, namely photolithography and the like, or additive methods, namely printing, electro-deposition methods and the like.
21. 21. The methods according to claims from 14 to 19 wherein said first, second and further working electrodes, and said reference and counter electrodes are formed by electrically conductive wires of given materials, e.g. Platinum, Silver, Gold and the like; and wherein the electrically conductive wires of said first, second and further working electrodes, said counter and reference electrodes are coated with given functional layers, such as pH sensing layers, enzyme-containing membranes, reference electrode and counter electrode layers, deposited by either conformal coating techniques, such as thermal evaporation, physical or chemical vapor deposition, electro-deposition or electro-plating, soaking in an electrolyte solution; and wherein the cylindrical wires coated with the above mentioned materials and techniques are mechanically assembled within a cylindrical rigid or flexible hollow needle of a given diameter by means of an inert polymeric material which is preferably selected among the materials that are permeable or semipermeable to both the hydrogen ions and chemical substances present in the solution, formed for instance by poly(dimethylsiloxane) (PDMS), and directly exposed in the terminal end of the wires to the solution.The methods according to the claim 21 wherein one or more electrodes are formed by uncoated wires of a given material such as gold, platinum and the like, carbon, and the like.
22. 22. The methods to read the signals generated by the multi -analyte sensing devices formed according to any of the methods indicated in claims from 14 to 22, wherein a sensor interface circuit is composed by a set of operational amplifiers and resistors in their feedback loop perform as current to voltage converter for measuring the current flows generated in said first, second and further working electrodes of the multi-analyte sensing device when its counter electrode is maintained at a fixed potential with respect to the reference electrode by means of a differential amplifier connected in a voltage follower configuration and converting the measured current flows into voltage signals by means of transconductance amplifiers in such a way to generate a number of analogue signals, namely current signals, that correspond respectively to: the pH level of a solution, amperometrically assessed onto the surface of said first working electrode; the hydrogen ions generated by a first enzymatic reaction and amperometrically assessed onto the surface of one of said working electrodes after the first one, for instance measuring the glucose level of the solution when the enzyme-containing membrane is made for being selective to glucose according to an oxygen-independent chemical reaction, such as in the reaction (5), where glucose reacts with nicotinamide adenine dinucleotide (NAD+) in presence of glucose dehydrogenase (GDH), producing gluconic acid, nicotinamide adenine dinucleotide in its reduced form (NADH) and hydrogen ion (Hi); and /or the hydrogen ions generated by a second enzymatic reaction and amperometrically assessed onto the surface of one of said working electrodes after the first one, for instance measuring the lactate level of the solution when the enzyme-containing membrane is made for being selective to lactate according to an oxygen-independent chemical reaction, such as in the reaction (4), where lactate reacts with nicotinamide adenine dinucleotide (NAD+) in presence of L-lactate dehydrogenase (LDH), producing pyruvate, nicotinamide adenine dinucleotide in its reduced form (NADH) and hydrogen ion (Hi).
23. 23. The method of claim 23 further comprising an operational amplifier and a resistor in its feedback loop in order to perform as current to voltage convener of the current flow measured onto one of the working electrodes after the first one and to generate an analogue signal, namely a voltage signal, that corresponds to the hydrogen ions generated by an enzymatic reaction and amperometrically assessed onto the surface of one of the working electrodes after the first one, for measuring the level of carbon dioxide molecules in the solution when the enzyme-containing membrane is made for being selective to carbon dioxide according to an enzymatic chemical reaction, such as in the reaction (6), where carbon dioxide (CO2) reacts with water (I-120) in presence of the carbonic anhydrase (CAH) enzyme, producing hydrogen carbonate ion (HCO3) and hydrogen ions (H).
24. 24. The method of claims 23 or 24 wherein the output signals corresponding to the readout levels of pH and two or more chemical substances dissolved in the solution where the multi-analyte sensor device is used, are further elaborated in order to implement a given compensation algorithm, to eliminate the dependence of the read-out signals corresponding to the concentration of target chemical substances in the solution from the pH level of the same solution, which is assessed in the surface of the first working electrode.
25. 25. The method of claim 25 wherein said compensation algorithm is implemented by means of a code within a programmable microcontroller device and wherein the uncompensated signals generated by the sensor interface circuit of claim 23 are introduced in the microcontroller by means of an analog to digital converter and multiplexer circuit block and data processed to obtain the compensated signals, after filtering the variations of pH in the solution according to the signal generated by the first working electrode and preprocessed by the sensor interface circuit therein.
26. 26. A multi-electrode sensing device realized according to any of the methods of claims from 1 to 7, or claims from 14 to 20 where the corresponding electrodes are formed on a flexible and /or stretchable and /or biocompatible and/or biodegradable substrate selected among the different classes of materials such as polyester terephthalate (PET), polyethylene nafthalate (PEN), polyimmide (P1), poly(dimethylsiloxane) (PDMS), Polyether ether keton (PEEK), poly(lactic glycolic acid) (PLGA), Hyarulonic Acid (HA)-based films, parylene, several hydrocolloid films and the like, in such a way to be implemented in form of a skin-compatible patch device for monitoring the skin pH level and concentration of specific analytes from sweat.
27. 27. A multi-electrode sensing device according to claim 27 where said device is fabricated on substrates or coated by hiocompatible and/or biodegradable materials in order to be embedded into an implantable device for monitoring the pH level and concentration of specific analytes in situ within the body of a living organism, such as subcutaneously, sub-parenchima, sub-dura, in the sub-arachnoid space, or other location inside the body for monitoring the said analytes and pH of the biological fluids present in the corresponding regions inside the body, such as blood, plasma, cerebrospinal fluid, swab, and the like.
28. 28. A device as according to claims 27 or 28 wherein the device is designed and manufactured in form of a contact lens for monitoring said pH level and specific analytes from tears.
29. 29. A multi-electrode sensing device realized according to any of the methods of claims from 1 to 6 and claims from 8 to 9, or claims from 14 to 19 and claims from 21 to 22 where the corresponding electrodes are formed and encapsulated within a cylindrically shaped needle device or another elongated structure, composed by a rigid, flexible, or semi-flexible hiocompatible and /or biodegradable substrate selected among the different classes of materials suitable for the fabrication of implantable devices such as stainless steel, titanium, glass, polycarbonate, polyurethane, polyethylene, poly(dimethylsiloxane), PEEK, and the like and designed and manufactured for use in implantable devices such as various catheters for sub-parenchymal introduction used in traumatic brain injuries, or implants for monitoring and treatment of spinal cord injuries, embedded or combined with various fluidic devices used for release of drugs in situ and /or with microdyalisis probes for sampling of biological fluids.
30. 30. A multi-electrode sensing device according to any of the claims from 27 to 30 wherein the device is combined with the corresponding sensor interface electronics according to the methods of claims from 10 to 13, when said multi-electrode sensing device is formed according to any of the methods of claims from 1 to 9, or according to the methods of claims from 23 to 26, when said multi-electrode sensing device is formed according to any of the methods of claims from 14 to 22.