Flexible microneedle electrode for biopotential monitoring, method for constructing flexible microneedle electrode and patch electrode comprising flexible microneedle electrode

Abstract

The invention discloses a flexible microneedle electrode for biopotential monitoring, a method for constructing the flexible microneedle electrode and a patch electrode comprising the flexible microneedle electrode. The method comprises the following steps: providing a negative mold with a plurality of microneedle structures; depositing at least one layer of conductive material on the negative mold; and stripping the at least one layer of conductive material from the negative mold to obtain the flexible microneedle electrode comprising a plurality of microneedle structures.

Description

Technical Field

[0001] The present invention relates to a flexible microneedle electrode for biopotential monitoring, a method for constructing a flexible microneedle electrode, and a patch electrode including the flexible microneedle electrode, in particular an ultra-thin and flexible microneedle electrode for accurate and long-term biopotential monitoring. Background Art

[0002] In the digital healthcare era, the demand for wearable electrodes is increasing continuously, and there is an urgent need for more efficient and accurate biopotential monitoring technologies. Surface biopotentials are crucial for evaluating body functions and diagnosing diseases and are widely used in clinical practice, such as electrocardiogram (ECG), electromyogram (EMG), electroencephalogram (EEG), and electrooculogram (EOG). With the progress of bioelectric acquisition technologies and artificial intelligence (AI) algorithms, these biopotentials are also increasingly used in various applications such as sports training, prosthetic control, human-computer interaction, and virtual and augmented reality. However, existing electrodes on the market usually require skin pretreatment, which may cause discomfort and skin irritation. Over time, the signals will degrade due to gel evaporation, and there are limitations in terms of material cost, adhesion, flexibility, and production scalability. Summary of the Invention

[0003] According to a first aspect of the present invention, there is provided a method for constructing a flexible microneedle electrode for biopotential monitoring, including the following steps: providing a negative mold having a plurality of microneedle structures; depositing at least one layer of conductive material on the negative mold; and peeling off at least one layer of conductive material from the negative mold to obtain a flexible microneedle electrode including a plurality of microneedle structures.

[0004] According to the first aspect, the negative mold includes a mold substrate manufactured using nanoimprint lithography.

[0005] According to the first aspect, the negative mold is defined as having a plurality of microneedle structures in a pyramid shape, a cone shape, or a cylindrical shape.

[0006] According to the first aspect, the height range of the plurality of microneedle structures is from 20 to 200 micrometers, the spacing range is from 50 to 500 micrometers, and the length range is from 20 to 300 micrometers.

[0007] According to the first aspect, the step of providing a negative mold having a plurality of microneedle structures includes the following steps: providing a molding material to replicate a plurality of microneedle structures from a positive mold; and curing the molding material to provide a mold substrate.

[0008] According to the first aspect, the positive mold includes a PDMS positive mold defined with a plurality of microneedle structures.

[0009] According to the first aspect, the molding material includes a polymer molding material.

[0010] According to a first aspect, the molding material is ultraviolet curable.

[0011] According to a first aspect, the negative mold is conductive, and at least one layer of conductive material is deposited on the negative mold by electrodeposition.

[0012] According to a first aspect, the negative mold includes an indium tin oxide (ITO) layer covering the mold substrate.

[0013] According to a first aspect, the step of providing a negative mold having a plurality of microneedle structures further includes the step of depositing an ITO layer on the mold substrate to provide the negative mold.

[0014] According to a first aspect, the ITO layer is deposited on the mold substrate by sputtering.

[0015] According to a first aspect, the thickness of the ITO layer is about 250 nanometers.

[0016] According to a first aspect, at least one layer of conductive material includes gold (Au) and nickel (Ni).

[0017] According to a first aspect, the gold layer and the nickel layer are sequentially electrodeposited on the negative mold by a step current source.

[0018] According to a first aspect, the thicknesses of the gold layer and the nickel layer are 500 nanometers and 5 micrometers, respectively.

[0019] According to a second aspect of the present invention, there is provided a flexible micro-needle electrode (MNE) for biopotential monitoring, including at least one layer of conductive material made by the method of the first aspect and having a plurality of microneedle structures.

[0020] According to a third aspect of the present invention, there is provided a patch electrode for biopotential monitoring, including the flexible micro-needle electrode (MNE) for biopotential monitoring according to the second aspect; and a conductor configured to electrically connect the flexible micro-needle electrode to a biopotential monitoring device.

[0021] According to the third aspect, the patch electrode is a dry electrode suitable for patient wear.

[0022] According to the third aspect, the flexible micro-needle electrode is suitable for being worn for at least twenty-four hours without losing performance.

[0023] The present invention provides a solution with improved user comfort, reduced skin irritation, higher signal-to-noise ratio during movement, and a more cost-effective manufacturing process.

[0024] The present invention provides a remarkable and simple method for fabricating highly conductive, flexible, and ultra-thin microneedle electrodes (MNE) for precise and imperceptible biopotential monitoring, utilizing a unique micro / nano electrodeposition technique. To achieve a cost-effective and scalable manufacturing process, a metal layer is electrodeposited on an indium tin oxide (ITO) substrate with a microneedle structure, such that the electrodeposited metal thin film exhibits the same microstructure as the underlying ITO substrate. This method uses magnetron sputtering of indium tin oxide (ITO) on a microneedle mold and subsequent electrodeposition of a metal layer to produce ultra-thin, flexible, and highly conductive MNE. The performance of these MNE is superior to that of electroplated planar electrodes and wet silver / silver chloride (Ag / AgCl) electrodes. BRIEF DESCRIPTION OF THE DRAWINGS

[0025] The foregoing and other objects and advantages of the present invention will become more apparent and understandable when considered in conjunction with the following detailed description and the accompanying drawings, in which like reference numerals represent like components in the various views, and in which:

[0026] Figure 1A is a flow chart of a method for constructing a flexible microneedle electrode (MNE) for biopotential monitoring according to an embodiment of the present invention.

[0027] Figure 1B is an optical image of a flexible MNE fabricated according to an embodiment of the present invention.

[0028] Figure 1C is Figure 1B an optical microscopic image of the MNE in

[0029] Figure 1D is Figure 1B an optical microscopic image of the MNE in

[0030] Figure 2A is a diagram of the microneedle structure of an MNE fabricated according to an embodiment of the present invention.

[0031] Figure 2B shows Figure 1A the structures on different material layers of a microneedle array electrode and a negative mold used for fabricating an MNE in the process shown in

[0032] Figure 2C shows Figure 2A the dimensions of the microneedle array electrode structure of

[0033] Figure 3A is a diagram of the electrode-skin interface of a wet Ag / AgCl electrode.

[0034] Figure 3B is an equivalent circuit model of the electrode-skin interface of a wet Ag / AgCl electrode.

[0035] Figure 3C It is a diagram of the electrode-skin interface of a planar electrode.

[0036] Figure 3D It is an equivalent circuit model of the electrode-skin interface of a planar electrode.

[0037] Figure 3E It is a diagram of the electrode-skin interface of MNE.

[0038] Figure 3F It is an equivalent circuit model of the electrode-skin interface of MNE.

[0039] Figure 4 It is an equivalent circuit for electrode-skin interface impedance (EII) testing.

[0040] Figure 5A It is a diagram of a planar electrode.

[0041] Figure 5B It is a diagram of a wet Ag / AgCl electrode.

[0042] Figure 5C It is a diagram of MNE.

[0043] Figure 5D It is a diagram of a patient wearing a pair of MNEs on the arm for EII testing.

[0044] Figure 6A It is a chart showing the EII spectra of different electrodes, covering frequencies from 1 Hz to 10 5 Hz.

[0045] Figure 6B It is a bar chart showing the extracted EII of different electrodes at current frequencies of 1000 Hz, 100 Hz, 10 Hz, and 1 Hz.

[0046] Figure 6C It is a chart showing the comparison of the EII spectra of MNE and wet Ag / AgCl electrodes immediately after application and 24 hours later.

[0047] Figure 6D It is a photograph and chart showing the normalized resistance change of the electrode at bending radii from 3 to 10 mm.

[0048] Figure 6E It is a photograph of the electrode wearing position after the patient immediately removes the MNE after continuous wearing for 2 hours.

[0049] Figure 6F It is a photograph of the electrode wearing position after the patient immediately removes the wet Ag / AgCl electrode after continuous wearing for 2 hours.

[0050] Figure 7AIs a photograph showing the characterization of the skin reaction at the wearing position after removing the wet Ag / AgCl electrode and MNE after continuous wearing for 2 hours.

[0051] Figure 7B Is a photograph showing the characterization of the skin reaction at the wearing position captured 0.5 hour after removing the wet Ag / AgCl electrode and MNE after continuous wearing for 2 hours.

[0052] Figure 8 Is a typical ECG waveform and its characteristic pattern.

[0053] Figure 9 Is a diagram of the ECG detection in Lead II configuration.

[0054] Figure 10A Is a graph showing a series of ECG signals recorded by planar electrodes.

[0055] Figure 10B Is an enlarged view showing a single characteristic signal recorded by planar electrodes.

[0056] Figure 10C Is a graph showing a series of ECG signals recorded by wet Ag / AgCl electrodes.

[0057] Figure 10D Is an enlarged view showing a single characteristic signal recorded by wet Ag / AgCl electrodes.

[0058] Figure 10E Is a graph showing a series of ECG signals recorded by MNE.

[0059] Figure 10F Is an enlarged view showing a single characteristic signal recorded by MNE.

[0060] Figure 11A Is a graph showing simultaneous ECG signals recorded by planar electrodes (top), wet Ag / AgCl electrodes (middle), and MNE (bottom) during static standing.

[0061] Figure 11B Is a graph showing simultaneous ECG signals recorded by planar electrodes (top), wet Ag / AgCl electrodes (middle), and MNE (bottom) during mild exercise.

[0062] Figure 11C Is a graph showing simultaneous ECG signals recorded by planar electrodes (top), wet Ag / AgCl electrodes (middle), and MNE (bottom) during vigorous exercise.

[0063] Figure 11DIt is a graph showing the simultaneous ECG signals recorded by a planar electrode (top), a wet Ag / AgCl electrode (middle), and an MNE (bottom) during deep breathing.

[0064] Figure 11E It shows the comparison results of the difference (MAX - MIN) between the maximum and minimum values of the PQRS characteristic values of the electrocardiogram signal monitored by the electrode under different motion states, so as to quantitatively evaluate the anti - motion artifact performance of the electrode.

[0065] Figure 11F It is a graph showing the standard deviation (SD) of all ECG data points.

[0066] Figure 12A It is a graph showing the EMG signals captured by an MNE under 5 kg, 10 kg, and 15 kg grip forces.

[0067] Figure 12B It is a bar graph showing the relationship between the change in the amplitude of the EMG signal and different grip forces.

[0068] Figure 12C It is a bar graph showing the RMS signal, RMS noise, and SNR values calculated from the EMG signal during wrist elevation.

[0069] Figure 12D It is a photo showing a robotic hand (open hand) controlled using the EMG signal measured by an MNE.

[0070] Figure 12E It is a photo showing a robotic hand (closed hand) controlled using the EMG signal measured by an MNE.

[0071] Figure 12F It is a graph showing a photo of a hand and the EMG signals generated by the flexion and extension of different fingers of the hand.

[0072] Figure 13A It is a graph showing the baseline noise captured by a planar electrode.

[0073] Figure 13B It is a graph showing the overall EMG signal captured by a planar electrode under a 5 - kg grip force.

[0074] Figure 13C It is a graph showing the EMG signal under a 5 - kg grip force captured by a planar electrode during muscle activation.

[0075] Figure 13D It is a graph showing the baseline noise captured by a wet Ag / AgCl electrode.

[0076] Figure 13E It is a graph showing the overall EMG signal captured by a wet Ag / AgCl electrode under a 5 - kg grip force.

[0077] Figure 13F It is a graph showing the EMG signals at a 5 kg grip force captured by wet Ag / AgCl electrodes during muscle activation.

[0078] Figure 13G It is a graph showing the baseline noise captured by MNE.

[0079] Figure 13H It is a graph showing the overall EMG signals captured by MNE at a 5 kg grip force.

[0080] Figure 13I It is a graph showing the EMG signals at a 5 kg grip force captured by MNE during muscle activation.

[0081] Figure 14A Schematic diagram showing the positions where the electrodes detect the electromyogram signals of the left forearm flexor muscles when the left hand pulls the grip dynamometer.

[0082] Figure 14B Schematic diagram showing the positions where the electrodes detect the electromyogram signals of the left forearm extensor digitorum muscles when the left hand wrist is lifted from parallel to the left arm to 90° vertically.

[0083] Figure 14C Schematic diagram showing the positions where the electrodes detect the electromyogram signals of the left forearm extensor digitorum muscles when the left hand opens and closes.

[0084] Figure 14D Schematic diagram showing the positions where the electrodes detect the electromyogram signals of the right posterior arm biceps brachii muscles when the right forearm is bent upward at the elbow by 120°.

[0085] Figure 15A It is a graph showing the SNR of the left forearm flexor muscles when monitoring the left hand pulling the grip dynamometer at 5, 10, and 15 kg using 3 different types of electrodes.

[0086] Figure 15B It is a graph showing the signals, RMS amplitudes of noise, and SNR of the left forearm extensor muscles (left hand wrist lifted) monitored using 3 different types of electrodes.

[0087] Figure 15C It is a graph showing the signals, RMS amplitudes of noise, and SNR of the left forearm extensor muscles (left hand opening and closing) monitored using 3 different types of electrodes.

[0088] Figure 15D It is a graph showing the signals, RMS amplitudes of noise, and SNR of the right posterior arm biceps brachii muscles (elbow bent to 120°) monitored using 3 different types of electrodes. Detailed implementation

[0089] Through their experiments and trials, inventors have learned that Ag / AgCl gel electrodes operated with conductive gels can obtain surface bioelectrical potentials in clinical settings. Despite their widespread use, they face several drawbacks, including the need for skin pretreatment, discomfort, and potential skin irritation. Additionally, they are prone to signal degradation during continuous monitoring due to the evaporation of the liquid in the gel electrolyte.

[0090] Another type of electrode is the dry electrode, which does not require the use of an electrolyte gel, can eliminate problems associated with wet electrodes, and thus may be an alternative for long-term bioelectrical potential detection.

[0091] Dry electrodes can be further divided into capacitive (non-contact) electrodes and dry contact electrodes. Although capacitive electrodes exhibit better comfort during long-term wear without direct contact with the skin, some electrodes may encounter significant challenges due to ultra-high electrode-skin interface impedance (EII) and motion artifacts.

[0092] For dry electrodes, performance can be improved by including a poly(3,4-ethylenedioxythiophene):poly(styrenesulfonic acid) (PEDOT:PSS) layer, as it is easy to process and has excellent electrical properties, transparency, and biocompatibility. Although attempts have been made to mix it with waterborne polyurethane (WPU) and D-sorbitol, or combine it with two-dimensional materials such as graphene to improve its performance, the composite materials may result in reduced conductivity (and thus performance), and the manufacturing process is complex.

[0093] Preferably, noble metal biocompatible materials such as gold (Au), titanium (Ti), and platinum (Pt) may be beneficial for dry electrodes because they have high conductivity and good biocompatibility. However, their applications may be limited by low adhesion, low flexibility, and high cost. More preferably, microstructured electrodes can overcome these limitations. Among various structures, electrodes with microneedles may exhibit lower EII due to the advantages of the microneedle structure, as the surface area is increased and they are able to penetrate the stratum corneum, further reducing EII. Additionally, the microstructure can also improve flexibility and closer contact with the skin, thus further enhancing the signal quality.

[0094] In some example embodiments, the fabrication of electrodes with microneedle structures may heavily rely on expensive processing techniques such as lithography, etching, and metal deposition. Alternative methods may rely on nanosecond laser micromachining and 3D printing, but the fabricated electrodes may have a rough surface and blunt tips. Additionally, due to the visible microneedle shape and wearing discomfort, as well as the residual marks caused by skin penetration, patients may change their natural behavior patterns, which may affect the accuracy of electrophysiological quality assessment and disease diagnosis. In summary, these above factors may affect the performance or experience of using percutaneous microneedle electrodes.

[0095] In an example microfabrication manufacturing process for fabricating a microneedle array electrode, it involves multiple techniques, including standard machining for creating a microneedle template, barrel plating to prevent rusting, physical and chemical reactions to form a microneedle substrate, magnetron sputtering to coat the substrate metal layer, and constant current electrodeposition for surface modification. Among these techniques, machining, physical and chemical reactions, and magnetron sputtering are mainly used for producing the microneedle template and the metal layer. The fabricated microneedle array consists of a PEDOT:PSS, a metal (Ti / Au) conductive layer, and a polyimide substrate (PI).

[0096] In another example, 3D printing technology can be utilized to fabricate the microneedle template. Then a chemical reaction, called the silver mirror reaction, is carried out, and electroplating is performed to construct a conductive layer composed of silver (Ag) and nickel (Ni). Subsequently, the 3D printing material undergoes a chemical etching process involving strong base, resulting in its corrosion.

[0097] Another manufacturing process can be used, which involves multiple steps, including thermal oxidation of a silicon wafer to form a SiO2 layer, dicing with a wheel grinder to create periodic grooves in a two-dimensional square column array, isotropic chemical etching of the silicon wafer to produce a microneedle array, etching the SiO2 layer using hydrofluoric acid, and magnetron sputtering deposition of a metal (Au or Ti) to complete the fabrication of the conductive layer.

[0098] In an example armband device, the armband includes an electrophysiological detection sensor that employs unstructured dry metal electrodes machined directly on a metal block.

[0099] Another electrode belonging to the category of wet electrodes is a wet Ag / AgCl (silver / silver chloride) electrode based on a silver / chloride silver substrate, covered with a conductive adhesive gel on top. This polymeric material improves the adhesion to the skin and subsequently forms an equivalent circuit to collect physiological electrical signals. However, compared with dry electrodes, wet electrodes have some disadvantages, which will be further explained in this disclosure.

[0100] The present invention relates to the fabrication of highly conductive, flexible, and ultrathin microneedle electrodes (MNE) using a templated 3D electrodeposition method. This technique is cost-effective and suitable for practical applications as it avoids the need for expensive equipment and allows for the reuse of the electrodeposition template.

[0101] Reference Figure 1A , shows a method 100 for constructing a flexible micro-needle electrode (MNE) for biopotential monitoring according to an embodiment of the present invention, including the following steps: providing a negative mold 104 having a plurality of micro-needle structures; depositing at least one layer of conductive material on the negative mold 104; and peeling off the at least one layer of conductive material from the negative mold 104 to obtain a flexible micro-needle electrode 102 including a plurality of micro-needle structures.

[0102] In this example, the flexible microneedle electrode 102 can be used as an electrode for electrically connecting a biopotential monitoring device, such as a monitoring device for measuring electrocardiogram (ECG), electroencephalogram (EEG), and electromyogram (EMG). The electrode can act as a sensor to convert the ionic current of the body into an electronic current that can be measured by the device.

[0103] Reference Figures 1B to 1D , the fabricated microneedle electrode 102 (i.e., the MNE in this disclosure) is substantially planar, and a plurality of microneedle structures 106 are formed on one side of the planar electrode. The plurality of microneedle structures 106 are arranged at regular intervals or spacings, and the microneedle structures 106 are identical to each other. The microneedle structures 106 have sharp pyramidal tips, and the bases are attached to the planar electrode surface, so that the electrode 102 can be fixed on the skin surface (not shown), which may have different conditions, that is, the skin surface has sweat or oil, or has a rough morphology of the stratum corneum and / or the living epidermal layer in contact with the electrode, or even becomes rough due to scars or hair.

[0104] Alternatively, the microneedle structures 106 can adopt other shapes, such as conical or cylindrical, protruding from the planar electrode surface by a certain length to overcome the roughness of the skin surface. In addition, the electrode 102 is preferably flexible, so that the electrode 102 can be flexibly applied to the skin surface of a patient with curvature. Reference Figure 1B , the microneedle electrode 102 can be bent into a curve using tweezers.

[0105] The microneedle electrode 102 preferably includes two layers of conductive metal - gold (Au) 108 and nickel (Ni) 110. When the microneedle electrode 102 is applied to the skin surface, the skin surface contacts the gold layer 108. The thickness of the gold layer 108 can be 500 nanometers to ensure that the sheet resistance of the electrode is within an acceptable range in various applications, and the relatively thin gold layer 108 is supported by another thinner nickel layer 110, preferably about 5 microns, so that the microneedle electrode 102 has sufficient mechanical strength to meet the actual application of the microneedle electrode 102 while maintaining its flexibility. Those skilled in the art should understand that the thicknesses of the gold layer 108 and the nickel layer 110 can vary in other designs, which may require different electrical / mechanical properties, and the conductive material can also be composed of other conductive materials, that is, metals or compound-based conductive materials, or a combination of multiple layers of conductive materials. Any suitable manufacturing process can be used to form the necessary microneedle structures 106 on a substantially planar conductive material sheet.

[0106] Figure 1B The photo in [[ ]] highlights the ultra-high flexibility of the thin MNE 102, with a thickness of about 5.5 microns, providing a seamless combination with the skin to achieve high-fidelity signal detection. Figure 1C and Figure 1DShows the tilted view and top view of MNE 102 characterized by an optical microscope. The two-layer metal layer (Au / Ni) conductive film is uniformly covered by electrodeposition, with a smooth surface, suitable for use as a wearable electrode.

[0107] Reference Figure 1A , shows a diagram of method 100 according to a preferred embodiment of the present invention. In this embodiment, the manufacturing process generally includes preparing a mold (stamp or mold), such as using a photocurable material called OrmoStamp, manufactured by nanoimprinting, and the corresponding microneedle electrode 102.

[0108] Preferably, the mold is a negative mold 104 for producing a "positive structure", that is, the voids 112 (reverse image of the protruding structure) on the mold become protrusions (microneedle structure 106) after the deposited layer is peeled off. In the preferred embodiment further described, the negative mold 104 is conductive, and at least one layer of conductive material is deposited on the negative mold 104 by electrodeposition.

[0109] Preferably, to facilitate electrodeposition using the negative mold 104, the negative mold 104 is conductive, and at least one layer of conductive material is deposited on the negative mold 104 by electrodeposition. The gold layer 108 and the nickel layer 110 are sequentially electrodeposited on the negative mold 104 by a step current source. The preparation process of the negative mold is further described below.

[0110] Preferably, the negative mold 104 includes an indium tin oxide (ITO) layer 104A covering the mold substrate 104B, where the mold substrate 104B can be formed using an ultraviolet curable molding material (such as OrmoStamp). The OrmoStamp mold or the mold substrate 104B can be patterned using a positive mold (such as a PDMS mold 114) or a hard mold manufactured by photolithography and etching techniques.

[0111] The process starts with preparing a conductive mold, namely the negative mold 104, including the OrmoStamp substrate layer 104B and the ITO coating 104A, and these layers are further structured with pyramid-shaped microneedles (i.e., voids 112). Reference Figure 1A , the preparation work includes replicating the microneedle structure 106 from the PDMS mold 114 using an ultraviolet curable molding material (OrmoStamp). This step is crucial for creating the substrate, allowing for the effective deposition of conductive materials in subsequent stages. Next, an ITO layer approximately 250 nanometers thick is sputtered on the OrmoStamp surface by magnetron sputtering or other known deposition processes. This ITO layer makes the microneedle mold conductive and thus suitable for the subsequent electrodeposition process.

[0112] The next stage involves electro-depositing gold (Au) 108 and nickel (Ni) 110 layers with thicknesses of approximately 500 nanometers and approximately 5 micrometers, respectively, on the microneedle ITO substrate, in sequence, via a stepper current source and strategy. Au is selected because of its superior biocompatibility, excellent physical and chemical stability, and high electrical conductivity, making it an ideal candidate material for direct skin contact. Meanwhile, the Ni layer 110 is electro-deposited on the substrate or negative mold 104 to enhance the hardness and rigidity of the metal film and reduce the manufacturing cost. The stepper current source is applied because an excessively high deposition current would result in too fast an electro-deposition rate, leading to the formation of a rough and non-dense metal film on the ITO surface, which is prone to defects such as cracks or peeling.

[0113] After electro-deposition is completed, the MNE 102 on the negative substrate 104 is cleaned with deionized (DI) water to remove any residual electrolyte and finally dried with nitrogen. Finally, an independent Au / Ni metal layer with a microneedle structure is obtained by carefully peeling the electro-deposited metal layer from the ITO substrate. This step produces the final product, namely the flexible and highly conductive MNE 102. The entire process successfully addresses the resolution limitations often encountered in current manufacturing methods and provides a cost-effective solution by fabricating microneedle electrodes with micron-scale features.

[0114] Different from other manufacturing methods, the method of the present invention eliminates the need to use expensive equipment, and the ITO substrate can be reused for multiple electro-depositions, thus achieving cost-effectiveness and a simple procedure suitable for practical applications.

[0115] Other metals that can be electro-deposited and have similar biocompatibility and mechanical properties can also be used. The electro-deposition is carried out using an electrochemical workstation (CHI 660E), allowing precise control of the deposition process through a constant current. The electroplated samples are then thoroughly cleaned with deionized water to eliminate any residual electrolyte. Subsequently, the samples are dried using nitrogen in preparation for the final step.

[0116] Reference Figures 2A to 2C , the microneedle structure 106 of the MNE 102 fabricated by the method of Figure 1A has various key dimensions. Referring to the microscopic images shown in Figures 1B to 1D , it can be observed that the metal layer uniformly covers and has a smooth surface, being suitable for use as a wearable electrode. The base length or diameter, pitch, and height of the microneedles are approximately 50, approximately 70, and 34 micrometers, respectively, which conform to the dimensions of the microstructured PDMS mold 114. Microneedles with other geometries (such as conical, cylindrical, etc.) and dimensions (such as height ranging from 20 to 200 micrometers, pitch ranging from 50 to 500 micrometers, length ranging from 20 to 300 micrometers, etc.) can also be used. Additionally, referring to Figure 2A and 2B, the dimensions of the microneedle structures 106 on each individual layer, i.e., the nickel layer 110, the gold layer 108, and the negative mold 104, are the same and replicate the Figure 1A microneedle structure on the original PDMS mold 114 shown.

[0117] Now referring to Figures 3A to 3F , it is worth noting that hair 302, sebum 304, and skin irregularities 306 may impede optimal skin contact during biosignal acquisition, potentially affecting detection accuracy. As Figures 3E to 3F shown, the MNE 102 of the present invention is applied to the skin surface 312, and the microneedles 106 can effectively bypass sweat and sebum and maintain good contact with the uneven skin texture, providing better conductivity than the planar electrode 308. In addition, due to its increased total surface area, the MNE 102 has a larger skin contact area, and thus compared with the Figure 3C and 3D shown planar electrode models, the equivalent contact resistance of the MNE is reduced. This helps to reduce the electrode-skin interface impedance (EII) and motion artifacts during biosignal measurement. Compared with the equivalent circuit of the wet Ag / AgCl electrode 310 ( Figure 3A and 3B ), the EII of the wet Ag / AgCl electrode is higher than that of the MNE electrode due to the additional resistance / impedance introduced by the conductive gel layer 314. In addition, due to the improved manufacturing resolution of the microneedle structure 106, a painless and more comfortable wearing experience is provided for the user without perceiving the presence of the microneedles on the skin.

[0118] EII is crucial for the acquisition of high-quality electrophysiological signals. A lower EII indicates better signal quality, with a higher signal-to-noise ratio and reduced baseline drift. Referring to Figures 4 to 5D , the inventors conducted an EII analysis between the electrode and the skin. A pair of electrodes were placed on the inner side of the human forearm using medical tape with a spacing of 5 cm. The analysis was based on the Figure 4 shown equivalent circuit model. Figure 5A is a diagram of a planar electrode (0.9 cm 2 ), Figure 5B is a diagram of a wet Ag / AgCl electrode (2.01 cm 2 ), Figure 5C is a diagram of an MNE (0.64 cm 2 ), Figure 5D is a diagram of a patient wearing a pair of electrodes on the arm for EII testing. In this example, the patch electrode for biopotential monitoring may include the flexible microneedle electrode (MNE) described above; and a conductor configured to electrically connect the flexible microneedle electrode to a biopotential monitoring device for subsequent analysis, such as ECG, EMG, EEG, or EOG.

[0119] In this example experiment, EII analysis was performed on three types of electrodes, including wet gel Ag / AgCl electrodes (Cathay Manufacturing Corp, CH55RB), MNE, and electroplated planar electrodes. To minimize the influence of the connection between different electrodes and the instrument, the electrodes were fixed on standard snap connectors, and the Ag / AgCl plate was exposed after removing the gel. Considering the contact area between different types of electrodes and the skin, the EII values were normalized to more accurately represent the electrode performance as follows: |EII| normalized = |EII| measured ×A where, |EII| measured is the measured EII (kΩ), A is the contact area (cm 2 ), and |EII| normalized is the normalized EII (kΩ*cm 2 ).

[0120] Figure 6A Figure shows the normalized EII spectra of the three types of electrodes, with the current frequency range from 1 to 10 5 Hz, and the |EII| normalized values at specific frequencies are as Figure 6B shown. It is worth noting that MNE exhibits a lower |EII| normalized , which is better than wet electrodes and planar electrodes. This is particularly evident in the low-frequency bandwidth (<1 kHz), which covers the frequency ranges of EMG and ECG signals. Specifically, the |EII| normalized of MNE at 100 Hz is 4.7×10 4 kΩ*cm 2 , which is nearly 5 times and 3 times lower than that of planar electrodes and wet electrodes, respectively. Compared with planar electrodes, the reduction in EII of MNE strongly supports the significant contribution of the microneedle structure in enhancing signal quality by increasing the effective skin contact area. Specifically, due to the pyramid geometry of the microneedles protruding from the planar surface, their four-sided structure has a larger surface area than the unstructured planar electrodes.

[0121] In addition, the protruding profile of the pyramid enables it to traverse the uneven surface of the skin, including wrinkles and surface impurities such as sebum, sweat, and hair, thus promoting direct skin contact. In contrast, planar electrodes may reduce skin contact due to gaps caused by surface unevenness and impurities. By increasing the effective skin contact area, the microstructure of MNE reduces the interference of electrophysiological signals and the influence of surface impurities.

[0122] The inventors also further evaluated the long-term usability of MNE by measuring the |EII| normalized, it was observed that the MNE electrodes could be worn for at least twenty-four hours without losing performance. Figure 6C showed that the EII of the MNE remained stable after 24 hours, with a slight increase. In contrast, the |EII| of the wet Ag / AgCl after 24 hours normalized showed a significant increase, almost three times that of the EII measured immediately. The degradation of the wet Ag / AgCl can be explained by the drying of the gel over time, i.e., the observed phenomenon can be attributed to the degradation and drying of the conductive gel in the wet Ag / AgCl electrodes. Over time, these changes lead to an increase in the contact resistance of the wet Ag / AgCl electrodes. In contrast, the design and structure of the MNE avoid these problems, making the MNE a more favorable choice for long-term physiological signal acquisition.

[0123] Next, the electromechanical stability of the MNE was examined. The MNE was fixed on a polyethylene terephthalate (PET) substrate and then bent at different angles. The four-probe resistance change was monitored. It was observed that in the bending radius range of 3 to 10 mm, the resistance change of the electrode was negligible, as Figure 6D shown, indicating its ability to work as a flexible electrode. The enhanced flexibility of the MNE may be related to the microneedles, which act as stress reservoirs to accommodate the strain energy applied by external forces.

[0124] Another requirement for wearable electrodes is good comfort and high biocompatibility. Benefiting from the small features of the microneedle array on the MNE, painless and minimally invasive electrophysiological detection was achieved, suitable for long-term monitoring. Referring to Figure 6E and 6F , the skin conditions after wearing the MNE and wet Ag / AgCl electrodes for 2 hours were compared. The area within the circle represents the area where the two types of electrodes contact the skin, while the surrounding area represents the coverage of the same medical tape. It was clearly shown that the MNE could significantly reduce the red marks compared to the wet Ag / AgCl electrodes without using a conductive gel. No skin irritation was reported within the sensing area, while more severe redness and swelling were observed in the tape-covered area.

[0125] Referring to Figure 7A and 7B , it was observed that within 0.5 hours after removing the MNE, the marks of the MNE could completely disappear. In contrast, the residual redness and swelling of the gel electrode were still visible after the same time. In addition, compared with uncomfortable invasive microneedles, non-invasive microneedles do not leave micro-perforations and surface material residues on the skin surface during insertion and removal. This demonstrates the excellent biocompatibility and biosafety of the non-invasive MNE of the present invention during long-term wear.

[0126] The MNE of the present invention can be used as a wearable dry electrode to detect epidermal bioelectric potential. In the human body, each periodic contraction of the heart generates periodic changes in electric signals in various body regions, and this phenomenon is called ECG, as Figure 8 shown. For example, referring to Figure 9 , in order to record the ECG signal, two pairs of MNEs are placed respectively under the right collarbone and under the left collarbone, while the wet Ag / AgCl electrode as a reference is placed under the left rib.

[0127] The wet Ag / AgCl and electroplated planar electrodes are also used as working electrodes to measure the ECG signal for comparison. Figure 10A 、 10C and 10E show the ECG signals recorded by three types of electrodes at the same test point. In the static state, the enlarged views of the single characteristic waves of all electrodes are as shown in Figure 10B 、 10D and 10F. It is worth noting that compared with the wet electrode (1.28 mV) and the planar electrode (1.85 mV), the signal obtained from the MNE shows a larger peak-to-peak voltage (2.28 mV) and has more obvious ECG characteristic waves. The larger amplitude and distinguishable ECG characteristic waves captured by the MNE are easy to identify, showing the potential for clinical diagnosis of various cardiac signal abnormalities.

[0128] In addition, in the example experiment, the inventors also captured the ECG signal during various body movements to verify the anti-artifact ability of the MNE. The ECG signal was also collected simultaneously by three types of electrodes including the MNE, the planar electrode, and the wet Ag / AgCl electrode. The volunteer (24-year-old male) performed four consecutive activity states, including rest (standing still), mild exercise (walking slowly), strenuous exercise (running), and deep breathing. The signals of all three electrodes were recorded for 60 seconds, and 4.5 seconds were selected simultaneously in each activity state, as shown in Figures 11A to 11D shown.

[0129] Referring to Figure 11A , the ECG signals recorded by the three types of electrodes showed comparable performance in the standing still state, with distinguishable PQRST waveforms. However, when the volunteer walked slowly, it was observed that the baseline of the ECG signal of the planar electrode drifted significantly, as shown in Figure 11B shown. During more strenuous exercise, such as running, the baseline fluctuation intensified, as shown in Figure 11C shown. Due to the adhesion of the ionic gel, the wet Ag / AgCl electrode showed better anti-motion artifact ability than the planar electrode during body movement.

[0130] Advantageously, the MNE showed the best anti-artifact performance among the three types of electrodes and could maintain a stable baseline even during strenuous exercise. Finally, when transitioning from strenuous exercise to deep breathing, as shown in Figure 11DAs shown, the ECG signals recorded by all three types of electrodes can be clearly distinguished without any baseline drift. This observation indicates that the differences in ECG signals during mild and strenuous exercise, as described previously, are due to the artifact rejection ability of the electrodes, thus excluding other interfering factors. In addition, the maximum - minimum (MAX - MIN) of the PQRS complex of the ECG signal and the standard deviation (SD) of the QRS complex were calculated to evaluate the artifact level in the ECG signal, referring to Figure 11E and 11F . MNE showed the highest MAX - MIN value of the QRS complex and the lowest SD value in the active state. In summary, the measurement results indicate that MNE performs excellently in suppressing motion artifacts, superior to planar electrodes and wet electrodes. This enhanced ability stems from the prominent pyramidal microneedles, which not only provide a larger effective contact area but also provide lateral tangential forces to counteract motion.

[0131] In addition, the MNE manufactured according to the embodiments of the present invention can also be used as a dry electrode for detecting action potentials generated by muscle contractions. EMG signals play a crucial role in various applications such as medical diagnosis, sports training, human - machine interaction, etc. Referring to Figures 12A to 12F , the ability of MNE in EMG signal detection was evaluated. First, surface EMG recordings were made by placing a pair of MNEs parallel and transversely on the flexor digitorum muscle on the inner side of the volunteer's forearm, as Figure 14A shown, and a reference electrode was placed at the ankle.

[0132] In the experiment, the volunteer held a dynamometer and measured the grip strength with the same hand. To compare and evaluate the detection ability of different electrodes for EMG signals, the participants were required to repeatedly apply the dynamometer to 5 kg without prior knowledge of the electrode type. Between each application of force, sufficient rest time for the muscle was ensured, and the consistency of electrode placement was maintained. Figure 13B 、 13E and 13H show the EMG signals generated by the contraction of the flexor muscle captured by three types of electrodes. They all show comparable peak - valleys corresponding to motion activation and relaxation. To further evaluate the quality of the detected EMG signals, 500 - millisecond data was extracted from the center of the high - amplitude EMG signals generated by muscle activation, referring to Figure 13C 、 13F and 13I, while the noise segments were obtained by selecting 500 - millisecond data from the center of the relaxation part between every two EMG signals, referring to Figure 13A 、 13D and 13G. As an EMG detection electrode, MNE shows lower baseline noise and more distinguishable signals compared to the other two types of electrodes, thus helping to interpret muscle activity.

[0133] To further evaluate the capabilities of the MNE, EMG signals were measured using the MNE at different grip strengths, as Figure 12A shown. The peak-to-peak amplitude and signal intensity were consistent with the grip strength, as Figure 12B shown, and the peak of the EMG signal increased with the increase in the left hand grip strength. For wearable applications, it is highly desirable for dry electrodes to maintain reliable performance during body movement. Therefore, EMG detection during body movement was performed with the electrodes attached to a significantly curved skin surface under different scenarios, as Figures 14A to 14D shown.

[0134] As Figures 15A - 15D shown, the MNE exhibited the lowest baseline noise RMS value among various movements. In addition, experiments were designed to detect EMG signals in multiple different types of body movement states, including: contraction of the left forearm flexor at different grip strengths when pulling the left hand grip dynamometer, contraction of the left forearm extensor digitorum when the left hand wrist is lifted 90° from parallel to the left arm to vertical, relaxation and contraction of the left forearm extensor digitorum when the left hand opens and closes, and contraction of the right posterior arm biceps when the right forearm is bent 120° upward at the elbow. Among the three electrodes evaluated in the test, compared with wet Ag / AgCl electrodes and planar electrodes, the MNE exhibited significantly lower noise levels and superior SNR. Although planar electrodes obtained higher RMS during each muscle contraction, their noise was relatively high, resulting in an SNR lower than that of the MNE. These findings confirm that the MNE provides superior signal stability and accuracy during body movement, which is beneficial for electrophysiological monitoring in epidermal electronic devices.

[0135] The accurate and highly sensitive electrophysiological detection capabilities of the MNE demonstrate great potential for high-precision control of robotic hands. The EMG signals generated by the opening and closing actions of the hand, as Figure 14C shown, captured by the MNE, can be used as a human-machine interface to coordinate the real-time opening and closing actions of the robotic hand, as Figure 12D and 12E shown. In addition, the MNE can also identify low-amplitude EMG signals generated by finger flexion and extension. As Figure 12F shown, the distinguishable EMG signal intensity measured by the MNE demonstrates the potential of the MNE in human-machine interface applications. Generally speaking, it can be foreseen that the MNE of the present invention has high conductivity, adaptability, and mechanical flexibility, and is suitable for more complex applications involving human-machine interaction.

[0136] The MNE manufacturing materials of the present invention include a polydimethylsiloxane (PDMS) microneedle mold with a periodicity of about 70 microns, a substrate length of about 50×50 microns, and a height of about 34 microns. It was obtained from Nanjing University. OrmoPrime08 and OrmoStamp materials were purchased from Micro Resist Technology GmbH. The electroplating solutions, including Au (Plug N’Plate Gold Solution) and Ni (Plug N’Plate Nickel Solution), were procured from CASWELL, USA. The wet Ag / AgCl electrodes (CH55RB) were bought from Cathay Manufacturing Corp.

[0137] The fabrication of the OrmoStamp mold starts with a layer of OrmoPrime (with a film thickness of about 130 nm), which is first spin-coated on a clean 2 cm×2 cm glass at 4000 rpm for 60 seconds and then baked on a hot plate at 150 °C for 5 minutes. After the glass is cooled to room temperature in air, the ultraviolet curable molding material (OrmoStamp) is cast and cured by exposing it to 365 nm ultraviolet light using a lithography machine (URE 2000 / 35, Chinese Academy of Sciences, China) for 300 seconds, subsequently solidifying to accurately replicate the microstructure of the PDMS microneedle mold. Subsequently, an ITO layer with a thickness of about 250 nm is deposited on the OrmoStamp surface by magnetron sputtering to make the microneedle mold conductive for electroplating.

[0138] The electroplating of Ni and Au is achieved by a stepped current source using an electrochemical workstation (CHI660E). The ITO mold is successively immersed in the Au and Ni electroplating solutions for the electroplating of Au and Ni while being connected to the cathode of the current source. Correspondingly, the anode of the current source is connected to platinum (Pt) and nickel (Ni) rods immersed in the solution. After electroplating is completed, the sample is thoroughly rinsed with deionized (DI) water to remove any residual electrolyte and finally dried with nitrogen. Then, an independent Au / Ni metal layer with a microneedle structure is obtained by carefully peeling it off the ITO substrate.

[0139] To evaluate the ability of the MNE to monitor epidermal bioelectrical potentials, the EII was measured using a two - electrode system of an electrochemical workstation (CHI660E). Before the start of the test, the skin area where the electrodes were to be placed was thoroughly cleaned using a 75% alcohol solution. To minimize the influence of electrode - device connection in the study, both the MNE and the planar electrode were adhered to the Ag / AgCl plate after removing the gel using a conductive adhesive. Subsequently, they were connected to the instrument using standard button connectors. The center distance between the two electrodes was precisely set to 5 cm.

[0140] For ECG measurement, the volunteer (a 24-year-old male) was asked to wear two test electrodes, one placed under the right collarbone and the other under the left collarbone. In addition, a wet Ag / AgCl electrode was used as a reference and placed under the left rib, as Figure 9 shown. In the static state (upright sitting position), the ECG signal was first recorded and analyzed using a data acquisition system (PowerLab26T, ADInstruments Pty Ltd), with the band-pass filter set between 0.5 - 50 Hz. The connection points were encapsulated using waterproof and stretchable medical tape (3M Tegaderm) to ensure a stable recording system. Subsequently, the subject performed five different activity states in sequence, namely the static state (upright sitting position), standing still, mild exercise (slow walking), vigorous exercise (running), and deep breathing. A custom-made 32-channel biosignal recording system from the Neuroengineering and Clinical Electrophysiology Laboratory of the University of Hong Kong was used to simultaneously record the signals of three different types of electrodes during these activities. During the measurement, the filtering parameters were kept consistent.

[0141] EMG signals were also measured and analyzed using the data acquisition system (PowerLab 26T, ADInstruments Pty Ltd) with the same electrode connection procedure as the ECG test. The surface EMG signals generated by the extensor digitorum, flexor digitorum, and biceps brachii were tested, and their positions were as Figures 14A to 14D shown. The center distance between electrode pairs was precisely configured to 50 mm, and the reference electrode was specifically placed at the ankle. The signal-to-noise ratio (SNR) analysis of the EMG signal had to go through the following steps: the monitored raw EMG data was filtered through a two-part filter (including a 10 Hz high-pass finite impulse response (FIR) and a 50 Hz low-pass digital filter) to obtain a baseline without oscillations. At the same time, the power frequency interference was eliminated by using a 50 Hz harmonic infinite impulse response (IIR) comb digital filter.

[0142] Therefore, the inventors tested the basic functions and performance of MNE and demonstrated that they can clearly monitor and transmit bio-signals such as ECG and EMG. These MNE have micron-sized microneedle dimensions, addressing the drawbacks of commonly used wet Ag / AgCl electrodes and planar electrodes. MNE are fabricated through a low-cost large-scale manufacturing process of magnetron sputtering ITO on a microneedle mold and electrodepositing a microneedle metal layer. The produced MNE show superior advancements in electrical and mechanical stability, biocompatibility, and comfort. The introduction of microstructures enhances the precise detection ability for capturing electrophysiological signals such as EMG and ECG, and has minimal motion artifacts, outperforming wet electrodes and planar electrodes. In terms of comfort, traditional gel electrodes require skin pretreatment, which may cause discomfort and potential skin irritation. In contrast, due to its microneedle array structure of only a few tens of microns, MNE can comfortably adhere to the skin without causing pain or irritation.

[0143] The inventors designed dry electrodes to ensure high-quality signals and minimal motion interference for long-term biopotential monitoring. These dry electrodes can adhere very well to the skin while maintaining strong electromechanical stability and detection accuracy. Advantageously, the present invention presents a novel, cost-effective, and scalable surface biopotential monitoring solution by introducing highly conductive, flexible, and ultra-thin microneedle electrodes (MNE) fabricated using a 3D electrodeposition method. Different from traditional Ag / AgCl gel electrodes, these MNE eliminate the problems of skin pretreatment, discomfort, potential skin irritation, and degradation over time.

[0144] Compared with existing dry electrodes, the MNE fabricated according to the embodiments of the present invention overcomes the problems of high material cost, low adhesion, and low flexibility. This innovative design ensures that the MNE has excellent conductivity, can adhere comfortably to the skin without causing pain, and maintains a high signal-to-noise ratio, marking a significant advancement in the field of wearable electrophysiological monitoring.

[0145] Advantageously, the invention of MNE utilizes a cost-effective and scalable manufacturing process based on 3D electrodeposition. The invented MNE and manufacturing process address the long-standing challenges of existing electrodes for surface biopotential monitoring. The MNE of the present invention are dry electrodes that overcome the signal degradation problem caused by gel evaporation in commonly used gel electrodes. More importantly, due to the microneedle structure of MNE providing superior stability and the ability to penetrate skin surface contaminants, the signal-to-noise ratio of signal monitoring can be significantly improved, while reducing skin irritation and enhancing user comfort. In addition, the scalable and cost-effective manufacturing process based on 3D electrodeposition of the present invention enables the production of MNE with consistent performance and significantly reduced costs.

[0146] The MNE of the present invention has good electrical conductivity and excellent electromechanical stability and is used for wearable electrophysiological monitoring, which is confirmed by electromechanical bending tests. With microneedles having a height of about 34 microns, the MNE can adhere to the skin without causing any pain, ensuring good wearability and electrode-skin contact. In addition, the MNE exhibits excellent ability in detecting electrophysiological signals, with a signal-to-noise ratio 1.35 times higher than that of planar electrodes and 1.33 times higher than that of wet electrodes.

[0147] In addition, the MNE of the present invention also has a lower EII than electroplated planar electrodes and wet Ag / AgCl electrodes, which is attributed to the increased overall electrode surface area, conformal contact with the skin, and the ability to bypass the barriers of sweat and grease on the skin. In addition, the ability of the MNE to accurately and reliably detect electrophysiological signals (including EMG and ECG) has been demonstrated, with a signal-to-noise ratio higher than that of planar electrodes and wet electrodes. In addition, the MNE is less sensitive to body movement, which is crucial for electrophysiological monitoring. To fully utilize the potential applications of the MNE, human-machine interaction based on the EMG signals measured by the MNE was performed. In summary, the MNE becomes an ideal dry electrode to replace wet Ag / AgCl electrodes for long-term health monitoring and human-machine interaction.

[0148] In addition, the MNE fabricated according to the embodiments of the present invention exhibits excellent flexibility with almost no change in resistance when the bending radius is from 3 to 10 millimeters. In addition, traditional electrode manufacturing methods usually involve expensive equipment and procedures. The present invention overcomes these problems by using 3D electrodeposition of conductive metals on a structured ITO substrate. This manufacturing method of the MNE is scalable and cost-effective, including reusing the ITO substrate mold for multiple electrodepositions and avoiding the use of expensive processes.

[0149] The above are only specific embodiments of the present invention and are not intended to limit the protection scope of the present invention. Any modification or substitution that is obvious to those skilled in the art should be within the protection scope of the present invention. Therefore, the protection scope of the present invention shall be subject to the protection scope of the claims.

Claims

1. A method for constructing a flexible microneedle electrode for biopotential monitoring, comprising the following steps: providing a negative mold having a plurality of microneedle structures; depositing at least one layer of conductive material on the negative mold; as well as The at least one layer of conductive material is peeled off from the negative mold to obtain a flexible microneedle electrode comprising the plurality of microneedle structures. 2 . The method of claim 1 , wherein the negative mold comprises a mold substrate fabricated using nanoimprint lithography.

3. The method of claim 2, wherein the negative mold defines a plurality of microneedle structures having a pyramidal, conical, or cylindrical shape.

4. The method according to claim 2, wherein the plurality of microneedle structures have a height ranging from 20 to 200 microns, a pitch ranging from 50 to 500 microns, and a length ranging from 20 to 300 microns.

5. The method according to claim 2, wherein the step of providing a negative mold having a plurality of microneedle structures comprises the following steps: providing a molding material to replicate the plurality of microneedle structures from a positive mold; The molding material is cured to provide the mold substrate. The method of claim 5 , wherein the positive mold comprises a PDMS positive mold defining a plurality of microneedle structures. The method of claim 5 , wherein the molding material comprises a polymer molding material.

8. The method of claim 5, wherein the molding material is UV-curable.

9. The method of claim 2, wherein the negative mold is conductive and the at least one layer of conductive material is deposited on the negative mold by electrodeposition.

10. The method of claim 9, wherein the negative mold comprises an ITO layer covering the mold substrate.

11. The method according to claim 10, wherein the step of providing a negative mold having a plurality of microneedle structures further comprises the following steps: The ITO layer is deposited on the mold substrate to provide the negative mold.

12. The method of claim 11, wherein the ITO layer is deposited on the mold substrate by sputtering.

13. The method of claim 10, wherein the ITO layer has a thickness of about 250 nanometers.

14. The method of claim 9, wherein the at least one layer of conductive material comprises gold and nickel.

15. The method of claim 14, wherein the gold layer and the nickel layer are sequentially electrodeposited on the negative mold by a stepping current source.

16. The method of claim 14, wherein the thickness of the gold layer and the nickel layer are 500 nanometers and 5 micrometers, respectively.

17. A flexible microneedle electrode for biopotential monitoring, comprising at least one layer of conductive material made by the method of claim 11 and having a plurality of microneedle structures.

18. A patch electrode for biopotential monitoring, comprising the flexible microneedle electrode for biopotential monitoring according to claim 17; and an electrical conductor configured to electrically connect the flexible microneedle electrode to a biopotential monitoring device.

19. The patch electrode according to claim 18, wherein the patch electrode is a dry electrode suitable for being worn by a patient.

20. The patch electrode of claim 19, wherein the flexible microneedle electrode is suitable for being worn for at least twenty-four hours without losing performance.

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