FIELD OF THE INVENTION

The present invention relates to the general field of piezoresistive sensors and in particular such sensors suitable for detecting bodily motion of a subject.

BACKGROUND OF THE INVENTION

There is an ever-increasing demand in therapeutic and diagnostic applications to detect and/or measure bodily motion of a subject (human or animal).

Conventional techniques for monitoring bodily motions such as heartbeat, muscle function or brain activity include electrocardiogram (ECG), electromyogram (EMG) and electroencephalographic (EEG). Mechanical counterparts to such techniques also include mechanomyogram (MMG).

While conventional techniques enable the detection of most bodily motions, they typically require bulky equipment that is complicated to operate and detection often needs to be conducted under well-controlled conditions. That in turn limits their ability for continuously mapping the full range of bodily motions with high fidelity.

Due to at least their convenience of use so-called "wearable electronics", that seek to interface rigid electronics with the relatively soft form of a subject's body, are attracting increasing attention for monitoring bodily motion.

Given that skin can convey a diverse range of biological vital signs and it is inherently attached to a skeleton via fasciae, muscles and tendons, a variety of bodily motions can be mechanically translated through and monitored from skin deformations.

While a number of wearable electronic devices have been developed with sensors for detecting different types of bodily motions, including pulse, limb/digit movement and internal organ activity, they still remain limited in practical application.

For example, bodily motions include those that can be detected at low frequency (<5 Hz), such as pulse or simple limb/digit movement, and also those detected at relatively high frequency (e.g. 5-40 Hz), such as fast limb/digit movement or the oscillation of muscle and motor systems.

High fidelity detection of high frequency bodily motions require a sensor that not only exhibits a similar range of stretchability (e.g. >60% strain) and modulus (e.g. 25-225 kPa) to that of skin, but also the ability to accurately and simultaneously respond to stretches at high frequencies (e.g. up to about 40 Hz or greater).

Piezoresistive sensors have found considerable application in monitoring bodily motions. Conventional wearable piezoresistive sensors have to date been limited in their ability for high fidelity detection of both low and high frequency bodily motions.

Piezoresistive sensors are favoured due to their relatively high sensitivity, fast response, ease of fabrication and low power consumption. Such sensors typically make use of a piezoresistive material made from polymer comprising an electrically conducting material, for example graphene-based material, metal fibres or carbon nanotubes.

In more recent times, some research effort has been directed toward developing piezoresistive sensors based on elastic substrate/thin film piezoresistive material laminates. While demonstrating some benefits, the thin film piezoresistive materials typically have a much higher elastic modulus than the elastic substrate on which they are laminated resulting in de lamination (or even tearing) of the thin film upon the elastic substrate being subjected to tensile deformation. Such delamination or tearing of the thin film results in loss of sensitivity or even failure of the sensor.

Some research effort has also been directed at using ultralight piezoresistive cellular material in piezoresistive sensors.

Those skilled in the art will appreciate an ultralight material is a material having a density of less than 10 mg/cm ³ .

While ultralight material can exhibit advantageous piezoresistive properties, they are well known in the art for their inherent fragility.

In an attempt to take advantage of their excellent piezoresistive properties and avoid their inherent fragility, ultralight piezoresistive cellular materials have been infdtrated with elastomers.

Infdtrating such ultralight cellular materials with elastomer provides a simple and effective way for improving the mechanical properties and processability of the cellular material. By fdling pores of the cellular material with the elastomer, the cellular networks become strongly bonded to and within the polymer matrix. The resulting composite material can exhibit significantly improved mechanical strength and flexibility, especially stretchability, bendability and twistability. Other techniques have been developed for promoting strong bonds between a polymer matrix and ultralight piezoresistive materials to improve their durability and processability.

While such composite polymer/ultralight material structures have a practical effect of improving the durability of the ultralight material, numerous other unique features of the ultralight material are unfortunately sacrificed at the same time. For example, such hybridisation of the ultralight material and polymer significantly increases the weight and correspondingly the modulus of the resulting composite structure. Furthermore, due to the intrinsic features of the infiltrating polymer, such as electric insulation and frequency/time - dependent viscoelastic performance, fixing the ultralight material to/within the polymer matrix significantly changes the electromechanical performance of the ultralight material per se. Typical problems associated with such polymer infiltrated ultralight materials include their changes in resistance under cycling stretching/release and signal decay. Sensors based on such technology are therefore generally not very reliable.

Without wishing to be limited by theory, it is believed such problems stem from at least viscoelastic deformation of the infiltrating polymer, the effect of which inherently adversely affects the piezoresistive performance of the ultralight material bonded to or in the polymer. An opportunity therefore remains to develop a piezoresistive sensor that more effectively captures the merits of both an ultralight piezoresistive material (e.g. low modulus, broad and reliable piezoresistive sensing range, etc) and elastic material (e.g. good stretchability, bendability and flexibility, etc). SUMMARY OF THE INVENTION

The present invention provides a piezoresistive sensor comprising an ultralight piezoresistive foam material (UPFM) and an elastic substrate, wherein the UPFM is in releasable contact with the elastic substrate.

It has now surprisingly been found possible to produce a robust and durable piezoresistive sensor that has a dynamic sensing capability over a broad frequency range in high fidelity.

In contrast with conventional piezoresistive sensors comprising ultralight material, the sensor in accordance with the present invention does not include ultralight foam material that is strongly bonded to or interlocked by the elastic substrate. Instead, the UPFM is placed in releasable contact with the elastic substrate.

Surprisingly, placing the UPFM in releasable contact with the elastic substrate not only improves the durability, especially stretchability, bendability and flexibility, of the UPFM, but the UPFM substantially retains its piezoresistive properties, especially its dynamic sensing capability over a broad range of frequencies.

Unlike elastic substrate/thin film piezoresistive material laminates that have a single continuous surface area interaction point between the substrate and the thin film that is prone to delamination, the UPFM as used in accordance with the invention has a discontinuous interface (by virtue of its foam structure) that presents a high number of releasable interaction points at the interface between the substrate and the UPFM. Without wishing to be limited by theory, that high number of releasable interaction points is believed to surprisingly provide a pathway for at least the elastic substrate to impart improved physical durability to the UMFM without significantly adversely affecting the piezoresistive properties of the UPFM.

The piezoresistive sensor in accordance with the present invention therefore advantageously is able to derive the merits of both the UPFM and the elastic substrate to provide for a piezoresistive sensing system that exhibits high fidelity over a broad frequency range.

In one embodiment, the UPFM is located in between two layers of the elastic substrate.

In another embodiment, the UPFM is encapsulated by the elastic substrate.

In a further embodiment, the elastic substrate is an elastomer.

In a further embodiment, the elastomer is selected from polysiloxanes (e.g. poly(dimethyl siloxane)), natural rubbers, styrene -butadiene block copolymers, polyisoprene, polybutadiene, ethylene propylene rubber, ethylene propylene diene rubber, fluoro- elastomers, polyurethane elastomers and nitrile rubbers.

In another embodiment, the UPFM is a graphene-based aerogel. In yet a further embodiment, the UPFM has a bulk tensile modulus of less than about 2 kPa. In another embodiment, the piezoresistive sensor has a dynamic sensing capability over frequencies ranging from 0 to about 180 Hz, or from greater than 0 to about 180 Hz, or from about 0.02 to about 180 Hz.

The present invention also provides for use of a piezoresistive sensor according to the present invention in measuring or detecting bodily movement of subject.

Examples of bodily movements include, but are not limited to, pulse, limb or digit movement and general muscle activity.

The present invention also provides a method of manufacturing a piezoresistive sensor, the method comprising positioning an ultralight piezoresistive foam material (UPFM) in releasable contact with an elastic substrate.

In one embodiment, the UPFM is placed in releasable contact with the elastic substrate by locating the UPFM in between two layers of the elastic substrate.

In yet a further embodiment, the UPFM is placed in releasable contact with the elastic substrate by the UPFM being formed on a surface of the elastic substrate.

The present invention further provides a wearable article for measuring or detecting bodily motion of a subject, the article comprising a piezoresistive sensor according to the present invention.

Examples of wearable articles include, but are not limited to, glove, wrist band, arm band, ankle band, a leg band, a knee band (e.g. a knee compression band, or a knee support band), a collar band, a head band, and a chest band.

Further aspects and embodiments of the invention are described in more detail below. BRIEF DESCRIPTION OF THE DRAWINGS

The invention will hereinafter be described with reference to the following non-limiting drawings in which:

Figure 1 shows Scanning Electron Microscope (SEM) images of ultralight graphene aerogel (UGA) that can be used as UPFM in accordance with embodiments of the invention, showing the UGA in (a) honeycomb and (b) bonded fibre-like configuration. Suitable UPFM configuarations also include lamellar and micro-lattice configuration structures;

Figure 2 shows an embodiment procedure for the fabrication of an embodiment sensor;

Figure 3 shows (a) design parameters of an embodiment elastic substrate, (b) finite element analysis (FEA) computed stress concentration factors (SCF) against the loading tensile strain of the isolated UGA (top data points) and the UGA sandwiched between PDMS layers (bottom data points), respectively, and (c) Young’s modulus and the FEA computed mean interfacial shear stress required as a function of the UGA density for conformably deforming the sensor to a given 20% tensile strain;

Figure 4 shows (a) photos of an embodiment UGA/PDMS sensor subjected to tensile tests, and (b) corresponding stress-strain curve showing up to 100% elongation and soft modulus around 250 kPa;

Figure 5 shows cross-sectional optical microscopic images of an embodiment UGA/PDMS sensor undergoing strains from (a) 0%to (b) 30% (Scale bar: 100 pm), and (c) a schematic model of an embodiment UGA/PDMS sensor under stretching showing reversible microcracks generation with strain deformation;

Figure 6 shows relative change in electrical resistance of an embodiment UGA PDMS sensor in which the UGA was sandwiched between two PDMS layers having thickness of 500 um (a, b), 200 (c, d) um, and 100 um (e, f) in response to different applied tensile strains at 10% (a, c, e) and 100% (b, d, f), applied at a strain rate of 0.5 mm/s;

Figure 7 shows the electromechanical performance of an embodiment UGA/PDMS sensor in which the UGA was sandwiched between two 400 pm thick PDMS layers, each having in-built grooves with different thicknesses of 200 um (a, b), and 100 um (c, d) in response to different applied tensile strains at 10% (a, c) and 50% (b, d) respectively, applied at a strain rate of 0.5 mm/s;

Figure 8 shows the electromechanical performance of an embodiment UGA/PDMS sensor assembled by squeezing a (a) 12 mm thick and (b) 8 mm thick UGAs into a 400 pm deep groove in response to a 10% strain, applied at a strain rate of 0.5 mm/s;

Figure 9 shows the electromechanical performance of an embodiment UGA/PDMS sensor in response to stepped tensile strain changes (solid line curve), at one-step change of strain from 0 - 50% at 3 %/s rate (Inset: close-up of signal overshoot).

Figure 10 shows the electromechanical performance of an embodiment UGA/PDMS sensor in response to different periodic loading cycles under strains at a) 10%, b) 25%, c) 50% and d) 100% applied at a 5 mm/s rate;

Figure 11 shows hysteresis and durability characterization of an embodiment UGA/PDMS sensor (a) being stretched by 10%, 25%, 50%, 70%, and 100%, and (b) being stretched for 10, 100, and 1000 cycles at different strains varied from 10 % up to 50%;

Figure 12 shows frequency-dependent dynamic electromechanical performance of an embodiment UGA/PDMS sensor at (a) 0.28 Hz, (b) 5 Hz, (c, d) 25 Hz, (e) 15.5 Hz, and (f) 180 Hz;

Figure 13 shows the real-time electromechanical performance of an embodiment UGA/PDMS sensor applied to a user’s finger; Figure 14 shows the real-time electromechanical performance of an embodiment UGA/PDMS sensor applied to a user’s biceps; and Figure 15 shows (a) the real-time electromechanical performance of an embodiment UGA/PDMS sensor applied to a user’s biceps (top plot) compared to a rectified and integrated sEMG signal (bottom plot), and (b) the relative change in resistance of the sensor in response to different forces produced by isometric muscle contractions, which are defined as the different percentages of the maximum volume contraction (MVC), during voluntary muscle contraction;

Figure 16 shows (a) a comparison of the recorded relative resistance changes of a UGA/PDMS sensor (top curve) and voltage outputs of the sEMG (bottom curve) in response to stimulated biceps muscle contractions at different frequencies of 1 Hz (top), 3 Hz (middle) and 5 Hz (bottom), (b) the relative resistance changes of the sensor in response to the stimulated biceps muscle contractions at stimulation intensity of 50 mA (top), 20 mA (middle) and 10 mA (bottom), with the stimulation frequency of 3 Hz, and (c) a comparison of peak-to-trough values of the relative resistance changes of the sensor (square data points, left y-axis) with the peak-to-peak range changes of the recorded muscle action potential from the sEMG (circle data points, rigth y-axis) in response to the stimulated biceps muscle contractions at different stimulation intensities from 6 mA to 50 mA;

Figure 17 shows the resistance change recording from a UGA/PDMS sensor (top curve) and the sEMG muscle action potential recording (bottom curve) from a target biceps muscle group following a direct stimulation of the target peripheral nerve at Erb point;

Figure 18 shows average latencies between two output signals of a UGA/PDMS sensor and the sEMG when synchronously subjected to monitor stimulated biceps muscle contractions, in which (a) refers to the recorded signal average latencies (average of 6 trails) at different stimulation frequencies of 1 Hz, 3 Hz, and 5 Hz, using a stimulation intensity of 50 mA, and (b) the recorded average signal latencies (average of 6 trials) at different varied stimulation intensities of 10 mA, 20 mA, 30 mA, 40 mA and 50 mA, at a stimulation frequency of 3 Hz; and

Figure 19 shows (a) the relative change in resistance of a UGA/PDMS sensor used for monitoring stimulated wrist flexor muscle contractions at varied stimulation intensities of 25 mA, 20 mA, 12 mA and 10 mA using a stimulation frequency of 5 Hz, with the simultaneous sEMG recording at a stimulation intensity of 25 mA is shown in the bottom as a reference, and (b) a comparison of the absolute relative change in resistance of the UGA/PDMS sensor (square data points) and the absolute voltage change of the sEMG (circle data points) in response to the monitored stimulated wrist flexor muscle contractions at different given stimulation intensities.

DETAILED DESCRIPTION OF THE INVENTION

The present invention provides a piezoresistive sensor. Those skilled in the art will appreciate a sensor is a device that detects and responds to some type of input from the physical environment to which the sensor is exposed. A piezoresistive sensor comprises a piezoresistive material that can exhibit a change in electrical resistivity upon being subjected to mechanical strain. The piezoresistive sensor therefore practically detects an applied mechanical strain through measurement of a resistance change in the piezoresistive material in response to that material being subjected to the applied mechanical strain.

The piezoresistive sensors can be used to measure or detect bodily movement in a subject. By "subject" is meant an animal or human subject.

Piezoresistive materials are known in the art.

The present invention makes use of an ultralight piezoresistive foam material (UPFM). As known in the art, an "ultralight" material is one having a density less than 10 mg/cm ³ . Accordingly, the UPFM used in accordance with the invention is one having a density of less than 10mg/cm ³ . UPFMs are known in the art.

In one embodiment, the UPFM used in accordance with the invention has a density of less than 8 mg/cm ³ , or less than 6 mg/cm ³ , or less than 4 mg/cm ³ , or less than 3 mg/cm ³ , or less than 2.5 mg/cm ³ , or less than 2 mg/cm ³ , or less than 1.5 mg/cm ³ , or less than 1 mg/cm ³ .

The use of UPFM's with higher density can sometimes give rise to a non-linear sensor response between its relative resistance changes and an externally applied strain. Without wishing to be limited by theory, that effect is believed to be caused by slippage at the UPFM/elastic substrate interface. Surprisingly, the selection of a lower density UPFM in an equivalent system can reduce that phenomenon.

Lower density of the UPFM can assist with improved stretchability sensors in accordance with the invention due to its ultralow modulus which provides an extremely low mechanical resistance to the deformation. Although the interfacial bonding strength of the releasable interaction is weak, the conformable stretching of UPFM in the sensor can be effectively achieved. In another embodiment, the UPFM used in accordance with the invention has a density of less than 4 mg/cm ³ , or less than 3 mg/cm ³ , or less than 2.5 mg/cm ³ , or less than 2 mg/cm ³ , or less than 1.5 mg/cm ³ , or less than 1 mg/cm ³ .

There is no particular limitation on the physical shape the UPFM used in accordance with the invention may take.

As used in herein, the term "foam" in the context of the UPFM is intended to mean a material having a cellular or porous-like matrix or lattice. Examples of UPFM's suitable for use in accordance with the invention include, but are not limited to, silica aerogels, carbon nanotube aerogels, graphene -based aerogels, metallic foams, polymer foams, metallic microlattices.

The physical shape and size of the UPFM used in accordance with the invention will vary depending upon the configuration of the piezoresistive sensor, which in turn will typically be dictated by the intended application of the sensor.

There is no particular limitation on the shape and size of the UPFM used, apart from being configured to work in the sensor.

UPFM's used in accordance with the invention can advantageously be prepared using techniques known in the art.

In one embodiment, the UPFM is a graphene-based aerogel. By being "graphene-based" is intended to mean the aerogel may comprise one or more of graphene, graphene oxide, reduced graphene oxide and partially reduced graphene oxide.

An aerogel is in effect an ultralight foam that is derived from a gel.

Techniques for preparing graphene-based aerogels are known in the art. For example, such aerogel structures may be prepared in accordance with the teaching outlined in WO 2014/028978, the contents of which is incorporated herein by cross-reference. In that disclosure, the graphene-based aerogels are prepared by providing a dispersion of graphene sheets, partially reduced graphene oxide sheets, reduced graphene oxide sheets or a combination thereof in a freeze-castable medium and subjecting the so-formed dispersion to freeze-casting. The resulting graphene -based aerogel is then typically annealed.

By way of example, Figure 1 shows Scanning Electron Microscopy (SEM) images of embodiment graphene-based UPFMs having (a) honeycomb and (b) bonded fibre-like configurations. The UPFM may also have lamellar or micro-lattice configurations.

In one embodiment, the UPFM used in accordance with the invention has a bulk tensile modulus lower than about 2 kPa.

In one embodiment, the UPFM used in accordance with the invention is provided in the form of a layer having a thickness ranging from about 2mm to about 12mm.

The piezoresistive sensor in accordance with the invention also comprises an elastic substrate. By the expression "elastic substrate" it is meant a solid substrate that, upon application of a force, can be stretched in length from its relaxed, unstretched length, and which upon release of the stretching force will recover at least about 95 %, or 98 % or 99 % or 100 % of its elongation.

The type of elastic substrate used will primarily depend upon the intended application of the piezoresistive sensor. The elastic substrate will generally be selected on the basis of the degree of movement the sensor is intended to detect.

The elastic substrate may be in the form of a polymer.

In one embodiment, the elastic substrate is a polymer. Elastic polymers are known in the art as elastomers.

There is no particular limitation on the type of elastomer that may be used as the elastic substrate in accordance with the invention. In one embodiment, the elastic substrate is an elastomer selected from polysiloxanes (e.g. poly(dimethyl siloxane)), natural rubbers, styrene -butadiene block copolymers, polyisoprene, polybutadiene, ethylene propylene rubber, ethylene propylene diene rubber, fluoro-elastomers, polyurethane elastomers and nitrile rubbers. The size and shape of the elastic substrate used will generally be dictated by the intended application of the piezoresistive sensor. In one embodiment, the elastic substrate is provided in the form of a layer, for example in the form of a sheet or fdm. The thickness of such a layer (e.g. sheet or fdm) may range from about 100 pm to about 600 pm .

In one embodiment, the UPFM is located in between two layers of the elastic substrate.

In another embodiment, the UPFM is encapsulated by the elastic substrate. To enhance the releasable contact of the elastic substrate with the UPFM it may be desirable to treat the surface of the elastic substrate that comes into contact with the UPFM. Suitable surface treatment techniques include, but are not limited to, plasma surface treatment (e.g. argon or oxygen plasma surface treatment), functionalising the surface of the elastic substrate with charged (or chargeable) functional groups.

As used herein a "chargeable" functional groups is intended to mean a functional group that exists in a neutral state, but can be readily converted in to a charged state, for example through addition or removal of a hydrogen atom. Suitable chargeable functional groups are well known in the art and include amines (addition of a hydrogen atom to become positively charged) and organic acids (removal of a hydrogen atom to become negatively charged).

In one embodiment, the surface of the elastic substrate that comes into contact with the UPFM is a plasma treated surface. In another embodiment, the surface of the elastic substrate that comes into contact with the UPFM comprises charged or chargeable functional groups.

Such charged functional groups can be positive or negative charged functional groups. For example, positive charged functional groups can be provided by modifying the surface of the elastic substrate with amine functional groups and negative charged functional groups can be provided by modifying the surface of the elastic substrate with organic acid (e.g. - COOH) functional groups.

In one embodiment, the surface of the elastic substrate that comes into contact with the UPFM is modified with 3-aminopropyltriethoxysilane to provide positive charged functional groups.

In a further embodiment, the plasma treated surface is an oxygen or argon plasma treated surface. Adoption of an oxygen or argon-plasma treated surface can also improve the correlation between the applied strains and the resistivity changes of the sensor.

The elastic substrate may be provided with a channel or groove which assists with accommodating the UPFM within the sensor. For example, the elastic substrate may be in the form of a layer (e.g. sheet of film) having a channel or groove (i.e. an indentation) in a surface thereof, wherein the UPFM is located in the channel of groove.

In one embodiment, the elastic substrate is provided in the form of two layers and the UPFM is located in a channel or groove defined in one or both of the two layers.

In a further embodiment, the elastic substrate is in the form of a sheet or film that has a channel or groove therein, wherein the UPFM is located within the channel or groove.

A skilled person will be capable to envisage suitable sensor architectures in which the UPFM is located in between two layers of the elastic substrate. For example, the sensor may comprise the UPFM sandwiched between two opposing planar sheets made of the elastic substrate. This configuration may be obtained by placing the UPFM on the surface of one of the sheets, and using the other sheet to sandwich the UPFM between the two sheets such that the UPFM remains encapsulated between the sheets. This configuration is particularly suitable for UPFMs of relatively flat shape.

In some embodiments, at least one of the two sheets comprises a groove that receives the UPFM. An example of a planar sheet made of the elastic substrate and having a groove is shown in Figure 3. As seen from above (top image), the sheet may be dog -bone shaped with larger ends of width W i, and a middle section of smaller width W2. The sheet has a channel shaped groove located along the axis of the major dimension of the sheet. The groove has a width W3 and a depth t2.

Accordingly, in some embodiments the elastic substrate is provided in the form of two opposing planar sheets, and the UPFM is located in a channel or groove defined in one of the two opposing sheets. In those instances, the UPFM would be located within a cavity defined by the groove and a surface of the opposing sheet.

In some embodiments, the elastic substrate is provided in the form of two opposing planar sheets, and the UPFM is located in a channel or groove defined in both sheets. In those instances, the sheets would be positioned relative to one another such that the channels or grooves are aligned to define a cavity which depth is the sum of the depth of each channel or groove. An example of one such arrangement is shown in Figure 2. In the case of Figure 2, the depth of the channel defined by opposing grooves upon layering of the sheets is double the depth of each groove.

When located between two layers of the elastic substrate, the UPFM may be in a partially compressed state relative to the UPFM in its native isolated state (i.e. absent the substrate). This may be achieved, for example, by having one or both opposing sheet(s) defining a channel or groove which depth is smaller than the UPFM, resulting in the UPFM sitting within the channel or groove in a partially compressed state. In those instances, the UPFM located between the two layers of elastic material may have a volume that is at least 0.1% smaller than the volume of the UPFM in its native isolated state (i.e. absent the substrate). In some embodiments, the UPFM between the two layers of elastic material has a volume that is at least 1%, at least 10%, at least 25%, at least 50%, or at least 75 % smaller than the volume of the UPFM in its native isolated state. In those configurations, the UPFM fits snuggly in the void and makes good contact with the substrate. That in turn can translate into a particularly efficient transmission of mechanical vibrations from the substrate to the UPFM, providing the sensor with improved sensitivity and accuracy.

An important feature of the present invention relates to the UPFM being in releasable contact with the elastic substrate. By the UPFM being in "releasable contact" with the elastic substrate, is meant that while the UPFM is in physical contact with the elastic substrate, the UPFM is itself not strongly bonded or rigidly fixed to the elastic substrate. Accordingly, the UPFM has no primary bonds with or is not mechanically interlocked by the elastic substrate.

In other words, by the UPFM being in releasable contact with the elastic substrate, it is intended to mean there is no primary bond between the UPFM and the elastic substrate.

By "primary bond", it is meant a covalent, ionic or metallic bond.

By the UPFM being in releasable contact with the elastic substrate, it will be appreciated such contact is intended to exclude a situation where the UPFM is infiltrated by the elastic substrate, for example where the UPFM's cellular matrix is infiltrated by elastomer. Such an arrangement would of course provide for at least mechanically interlocking between the elastic substrate and the UPFM.

Without wishing to be limited by theory, it is believed the UPFM is in releasable contact with the elastic substrate through only so-called secondary bonds. As known to those skilled in the art, secondary bonds are intended to include weak attractive forces such as van der Waals forces, electrostatic forces and/or hydrogen bonds.

In the context of such secondary bonds, it will be appreciated any electrostatic forces at play are not intended to embrace ionic bonds.

As outlined above, in an attempt to address the fragility of UPFM's, composite structures have been developed whereby the UPFM is strongly bonded to an elastomer, for example by infiltrating an elastomer within the cellular matrix of the UPFM. While such composite structures are quite durable, the improvement in durability comes at the expense of a loss in electromechanical performance of the UPFM, such as its dynamic piezoresistive sensing behaviour.

Surprisingly, it has now been found that providing a UPFM in releasable contact with an elastic substrate not only enables the production of a piezoresistive sensor having excellent durability, stretchability, bendability and flexibility etc, but the resulting sensor has excellent sensitivity and accuracy, low response time, wide frequency sensing range and high repeatability over a wide range of frequencies.

Piezoresistive sensors in accordance with the invention can therefore advantageously exhibit both the physical and mechanical properties of the elastic substrate in combination with the high electromechanical performance of the UPFM, such as its dynamic piezoresistive sensing behaviour.

Without wishing to be bound by theory, it is believed the absence of strong bonds between the UPFM and the elastic substrate enables the UPFM to more efficiently retain and translate its piezoresistive properties upon the elastic substrate being subjected to strain. By the UPFM being in releasable contact with the elastic substrate, it is believed the UPFM on a molecular level can continuously release and grip the substrate as it is subjected to a tensile force thereby providing a means to prevent the entire force acting directly on the UPFM. That in turn enables the unadulterated UPFM to still sense the force and yet derive the practical durable elastic properties of the substrate. The sensor system according to the present invention therefore enables the UPFM to retain advantageous piezoresistive properties unlike, for example, conventional polymer infiltrated UPFM based systems.

As a result of the UPFM being in releasable contact with the elastic substrate, the UPFM can sustain significantly more strain without showing significant structural failure relative to an isolated for of the UPFM. Without wanting to be limited by theory, this could be possibly attributed to the uniform stretching environment created by the elastic substrate upon the UPFM. The interaction points generated at the interfacing areas of the UPFM and the substrate can serve as support pivots. These may allow effective stress transfer from the substrate to the UPFM, where the stress can effectively dissipate within the UPFM structure. Moreover, such uniform stretching environment may help distribute the tensile loading more uniformly within elements forming the structure of the UPFM, which in turn enhances the structural stretchability of the overall UPFM/substrate structure.

In some embodiments, the UPFM is encapsulated by the elastic substrate, the elastic substrate is moulded around the UPFM, or the UPFM is located between two layers of the elastic substrate.

While the UPFM may be contained by the elastic substrate in such ways, it is to be understood the UPFM will still be in releasable contact with the elastic substrate as described herein. In other words, while the UPFM may be physically contained by the elastic substrate, it will nevertheless remain in releasable contact with the substrate as described herein.

The piezoresistive sensor in accordance with the invention may be configured in any shape or size, with that shape or size typically being dictated by the intended application of the sensor.

The piezoresistive sensor in accordance with the invention of course also includes suitable electrodes that make electrical contact with the UPFM to enable a change in resistivity of the piezoresistive material to be measured. Those electrodes will typically comprise or be made of silver, gold, copper, platinum, carbon black or carbon fibers.

The present invention also provides a method of manufacturing a piezoresistive sensor. The method comprises positioning an UPFM in releasable contact with an elastic substrate. The UPFM and elastic substrate in that context are as herein described.

The method may comprise subjecting the surface of the elastic substrate that is to make contact with the UPFM to surface treatment before the UPFM is positioned to make releasable contact with the substrate. Such surface treatment may include plasma treatment and functionalising the surface of the elastic substrate with charged (or chargeable) functional groups as herein described.

In one embodiment, the method comprises subjecting the surface of the elastic substrate that is to make contact with the UPFM to surface treatment before the UPFM is positioned to make releasable contact with the substrate.

In another embodiment, the surface treatment is selected from plasma treatment and functionalising the surface of the elastic substrate with charged or chargeable functional groups.

Positioning of the UPFM in releasable contact with the elastic substrate may involve placing a pre-formed UPFM in contact with the elastic substrate or forming the UPFM directly on a surface of the elastic substrate.

In one embodiment, the method comprises positioning a pre-formed UPFM on a surface of the elastic substrate to be in releasable contact therewith.

In a further embodiment, the method comprises forming the UPFM on a surface of the elastic substrate such that the so-formed UPFM is in releasable contact with the substrate.

In another embodiment, the method comprises positioning the UPFM between two layers of the elastic substrate so as to be in releasable contact therewith.

In a further embodiment, one or both of the two layers of elastic substrate have a channel or groove in a surface thereof which upon coming together define a space within which the UPFM is positioned so as to be in releasable contact with the layers.

In a further embodiment, the method comprises positioning the UPFM is a space provided in a moulded form of the elastic substrate so at to be in releasable contact therewith. A piezoresistive sensor in accordance with the invention can be particularly sensitive and accurate for the detection of both low and high amplitude motions over a wide range of frequencies. Accordingly, the invention also relates to the use of a piezoresistive sensor of the kind described herein in measuring or detecting bodily movement of a subject.

In use, the sensor would be placed in direct physical contact with a target body that can emit mechanical skin vibrations/deformations, which would constitute the input signal of the sensor. The elastic substrate will then ensure efficient mechanical transmission of the input signal to the UPFM, which will react by changing its electric resistance. The change in electric resistance can be picked up and measured by suitable electrodes in electrical contact with the UPFM. The electrical signal measured by the electrodes would constitute the output signal of the sensor. The output signal may be in the form of any measurable electric signal that would be known to the skilled person. For example, the output signal may be in the form of electric resistance, electric voltage, electric current, or a combination thereof.

The output signal would be characterised by an amplitude which is proportional to that of the input signal. In some instances, such proportionality is linear. Accordingly, in some embodiments, the output signal of the sensor changes linearly with the input signal. That is, the sensor can provide linear electric response to the externally applied strains such that the electric resistance increases/decreases linearly with loading/unloading strains applied to the sensor. Such linear relationship is particularly advantageous in that it assists with the predictability and accuracy of the output signal. This can greatly simplify the overall signal pathway for real-life sensor applications, with significant benefits in terms of construction simplicity and sensor reliability. Also, high linearity between input strain and output signal of the sensor makes it highly reliable for the precise monitoring of bodily motions (e.g. the bending status of the human fingers) even at relatively high motion speed (e.g. > 5 Hz).

Typically, when the sensor is used to continuously monitor motions of a body, the profile of the output signal (i.e. the time-resolved amplitude of the output signal) repeats that of the input signal. For instance, when the amplitude of the input signal changes with a certain frequency, the amplitude of the output signal will change with substantially the same frequency. The particularly effective transmission of mechanical vibrations from the elastic substrate to the UPFM ensures that the sensor can detect motion accurately over a wide range of frequencies. In that regard, the sensor of the invention may be characterised by a frequency sensing range that encompasses frequencies typically associated with motion of any part of a subject body. For example, the sensor may have a frequency sensing range up to about 180 Hz.

The sensor of the invention is also advantageously accurate. That is, the sensor is able to clearly detect and discriminate small changes of the input signal. In that regard, the sensor of the invention can reliably detect changes of the input signal of less than 0.01%. For example, when monitoring mechanical strain, the sensor can accurately detect strain variations (e.g. amplitude or frequency changes) of less than 0.01%.

Further, the stable arrangement of the UPFM relative to the elastic substrate ensures the sensor is characterised by high signal repeatability. By “signal repeatability” is meant the ability of the sensor to consistently return a certain value of the output signal in correspondence to the same value of the input signal. As a result, the sensor of the invention is particularly reliable for the accurate monitoring of cyclical input signals, showing little to negligible signal hysteresis. This is believed to be linked to the negligible effect of the viscoelastic deformation of the substrate on the mechanical integrity of the UPFM. In some embodiments, the sensor provides an output signal which remains accurate within 0.1% upon application of an input signal (e.g. strain) that changes up to 10%, 25%, 50%, 75%, 100% of its initial value in a cyclical manner for at least 10, at least 100, or at least 1,000 cycles.

In addition, the specific arrangement of the UPFM in relation to the elastic substrate advantageously provides the sensor with fast response time . By “response time” is meant the time between a change in the input signal (e.g. frequency change, amplitude change, etc.) and a detectable change in the output signal. In some embodiments, the sensor provides a response time of less than 200 ms, less than 150 ms, less than 100 ms, less than 75 ms, less than 50 ms, or less than 25 ms. In those embodiments were the UPFM is sandwiched between two opposing layers (e.g. planar sheets) made of the elastic substrate, the thickness of the layers can influence the response time and modulate the linearity character of the sensor response over a wide range of input frequencies. In those instances, it is therefore possible to maximise sensor responsiveness and ensure input/output linearity of the sensor response over a wide range of input frequencies by having each layer, independently, with a thickness of at least about 150 pm. For example, each layer of the elastic substrate may independently have a thickness of at least about 200 pm, at least about 300 pm, at least about 400 pm, or at least about 500 pm.

In one embodiment, the elastic substrate is in the form of a layer having a thickness of at least about 200 pm, at least about 300 pm, at least about 400 pm, or at least about 500 pm.

Accordingly, a piezoresistive sensor in accordance with the invention can be particularly sensitive and accurate for the detection of both low and high amplitude motions over a wide range of frequencies. The specific arrangement of the UPFM in relation to the elastic substrate can also ensure that the sensor responds to external stimuli with extremely low response time. Those characteristics are combined with high repeatability of the output signal (i.e. the sensor’s ability to consistently provide the same output signal in response to the same input over time), making the sensor particularly advantageous for the real-time monitoring of complex bodily motions over prolonged activity time. Also, the combined high sensing performance in terms of sensitivity, accuracy, response time, and repeatability indicate that the sensor the invention is adequate for instantaneous time- and frequency- independent electromechanical responsive performance.

The sensor may find application in a number of devices. In that regard, the invention also provides a wearable article for measuring or detecting bodily motion of a subject, the article comprising a piezoresistive sensor of the kind described herein.

The wearable article may be one that can be worn directly on the body portion of interest. For example, the article may be in the form of a band that can be worn around the body portion of interest. Accordingly, in some embodiments the article is selected from a wrist band, an ankle band, a leg band, a knee band (e.g. a knee compression band, or aknee support band), a collar band, a head band, and a chest band. As a skilled person would understand, the article would be designed to ensure that once the article is worn, input signals associated with the movement of interest can be readily transmitted to the sensor. This may be achieved, for example, by locating the sensor in the portion of the article that will be in direct contact with the body part of interest.

The wearable article may also be provided in the form of a wearable patch. In those instances, the article may be applied directly on the body part of interest, for example by being adhered to that body part. In some embodiments, the wearable article is in the form of a stick-on patch. As a skilled person would understand, the patch would be designed to ensure that once the patch is adhered to the body part of interest, input signals associated with the movement of interest can be readily transmitted to the sensor. This may be achieved, for example, by ensuring that the elastic substrate is in direct contact with the body part of interest.

The wearable article may also be one that is worn as a conventional article of clothing for the continuous monitoring of motions of specific parts of a subject’s body. For example, the article may be selected from a glove, a sock, a shoe, a shirt, a top, pants, etc. As a skilled person would understand, such article of clothing would be one onto which the sensor can be integrated to ensure input signals associated with the movement of interest can be readily transmitted to the sensor. This may be achieved, for example, by positioning the sensor in the portion of the article corresponding to the body portion of interest. The sensor may be integrated to the article of clothing by any means known to the skilled person, including known fabrication techniques such as embroidery, sewing, weaving, non-woven, and knitting. EXAMPLES

EXAMPLE 1 Fabrication of ultralight graphene aerogels (UGA)

UGA was produced as an embodiment UPFM. To that effect, graphene oxide (GO) dispersions were fabricated using a modified Hummers method and further dispersed using a bath sonicator (Branson, B2500R) for 30 mins. The UGAs were subsequently fabricated via a modified one-step freeze-casting process. Specifically, a GO dispersion (12 ml; 1 mg/ml) was first mixed with ascorbic acid (AA) according to a weight ratio at 4:1 in a cylindrical glass tube. The mixture was then placed in a boiling water bath for 30 min to get a partially reduced graphene hydrogel. The glass tube was subsequently placed in a dry ice bath for 45 min to fully freeze the mixture, followed by a freeze-drying process. The obtained UGAs were annealed at 200 °C in air to further reduce the graphene oxide and bum off the AA residues. The obtained UGAs were than be laser-cut into a cuboid with a varied dimension of 10x 10x24 mm ³ , 8x8x24 mm ³ , 6x6x24 mm ³ for further use.

EXAMPLE 2

Assembly of UGA/PDMS strain sensors A number of UGA/PDMS sensors in the form of UGA sandwiched between two layers of PDMS were fabricated using a procedure of the king schematised in Figure 2.

3D-printed moulds with designed dog-bone shape geometries were firstly fabricated for producing each PDMS layer. Although Figure 2 shows the fabrication of PDMS layer with a groove, moulds for the production of PDMS layers without a groove were also fabricated (not shown).

An example of design parameters of a PDMS layer with a groove is shown in Figure 3. The width (Wi) and length (Li) of PDMS substrate were set at 35 mm and 70 mm respectively. The width of the dog-bone necking part (W2) and groove were set at 14 mm and 10 mm respectively. The thickness of the PDMS layer (ti) and groove (t2) were varied with different samples. The thickness of the in-built 3D-printed mould was modulated to obtain PDMS layers with thickness of 100 pm, 200 pm and 400 pm . The 3D-printed moulds were designed to provide each PDMS layer with a groove, which was intended to define an elongated receptacle for positioning the UGA during sensor assembly. For those samples with a groove, the thickness of the groove was set at 100, 200 and 300 ums, respectively.

The PDMS layers were prepared by pouring the de-bubbled mixture of the silicone elastomer base (Sylgard* 184) and the silicone elastomer curing agent (Sylgard* 184) (weight ratio at 10: 1) into the mould for curing at an ambient condition at 50 °C for 12 h. The mixture was evacuated in a vacuum chamber for 5 min before the polymerization process to remove the bubbles from the liquid PDMS.

To assemble UGAs into a strain sensor, two layers of Au electrodes (30 nm thick, see square elements in Figure 2) with T1O2 as an adhesive layer (2 nm) were first deposited onto the shaped PDMS films by electron-beam (E-beam) evaporator technique. Two conductive wires were sticked onto the two Au electrodes separately by Ag paste (one wire per electrode). Subsequently, laser-cut UGAs of the kind obtained by the procedure described in Example 1 were sandwiched in between the PDMS layers or encapsulated into the grooves formed by PDMS films with the two ends connected to the Au electrode areas separately. The oxygen plasma treatment at the PDMS surface for 2 mins.

EXAMPLE 3

Morphological and Mechanical Characterization

The dimensions and weights of UGAs were determined with a Vernier (Stamvick) with an accuracy of 0.01 mm and a balance (AND GH-252) with an accuracy of 0.01 mg. The SEM images of as-assembled UGAs were characterized by using a Nova (model) SEM operated at 5kV beam voltage. The internal microstructure morphologies of the assembled strain sensors were performed by using a laser microscope (VK-9700, Keyency). A homemade clamping device was fixed onto the stage of the microscope with two in-build strain applying screws to allow the in-situ microstructure examination under deformation.

Mechanical tensile tests were performed using a Mini Instron (Micro Tester, 5845, Instron) using a 100 N load cell with the strain control mode for static and low-frequency (< 1 Hz) dynamic testing. A small pre-strain was applied onto each sample for making sure no bending of the sample can occur during the cyclic deformations. The strain rate was set at 0.5mm/s. The applied strains were varied from 10% to 100%. For high-frequency (> 1 Hz) dynamic testing, the tensile tests were performed using an electromagnetic shaker (Bruel & Kjaer V200), a device that could provide accurate high-frequency vibrations, to provide high-frequency uniaxial deformations on the strain sensors with tuneable amplitudes. A function generator with a voltage amplifier was used to allow the shaker to generate vibrations with frequencies up to 180 Hz. The applied strain deformations onto the strain sensor were detected by a laser (LK-G32, Keyence) which detected the up-down movements of the shaker. Owning to the limited sample rate of the laser, the cyclic deformation of the shaker at a frequency higher than 25 Hz cannot be detected.

As can be seen in Figure 4(a), the UGAs with a density of 1 mg/cm ³ was assembled in between two PDMS layers, the resulting composite was stretched up from 0% to 30% strain without showing significant cracks or fractures along the UGAs network. The data confirms the large stretchability of UGA that can be reached using the proposed UGA/PDMS structures.

Figure 4(b) shows typical stress-strain curves obtained on UGAs/PDMS samples in a tensile test mode, with a strain rate at 0.5 mm/s. The structure displays excellent stretchability up to 100% elongations with extremely soft modululs of around 250 kPa. As can be seen in Figure 4(b), the strain sensor illustrated the high elasticity (> 100 %) with limited mechanical hysteresis performances and high linearity. Moreover, the UGAs/PDMS sample presented the ultrasoft tensile modulus of PDMS at around 250 kPa as no conductive fillers with higher stiffness have been embedded within the polymer matrix. EXAMPLE 4

Microstructural Characterization of UGA/PDMS samples

Stretchability examinations of UGAs/PDMS structures were carried out by monitoring in real time the microstructural changes of UGAs within the sandwich structure under tensile deformations using an optical microscope. Figure 5(a) shows atypical cross-sectional optical microscopic image of a sample UGA/PDMS structure. The porosity of the UGA is clearly detectable (1 mg/cm ³ density). The image can be compared with a corresponding SEM image of the isolated UGA shown in Figure 1(c). It can be noticed that in the assembled UGA/PDMS sample the porous structure of UGAs was slightly distorted and more aligned in the direction perpendicular to the applied pressure applied from the PDMS layers in the structure.

Increased strain steps at 0%, 15%, 45% and then releasing back after ten times of the loading/unloading cycles were performed. When the laminate structure was subjected to a uniaxial strain, both the UGA/PDMS interfacing areas and the porous internal structures of UGAs were stretched uniformly, showing the similar elongation value to that of the applied external strain, with the PDMS layer along the strain applied direction. When the external strain was released after ten cycles, the UGA/PDMS interfacing areas can mostly recover to their original states with negligible de-attachment occurred between graphene and the polymer layer, showing around 91.3% (which was calculated used the software ‘Image G) of the contacting areas of UGAs and PDMS still remained in contact.

Figure 5 illustrates the sequentially captured optical images of the low-density- UGAs/PDMS structures under tensile deformations from cross-sectional view. When subjected to a uniaxial tensile strain, the assembled UGAs within the structure demonstrated continued structural extension along the strain direction with the PDMS layers at both interfacial areas and intermediate sections uniformly. Figure 5 shows the cross-sectional optical microscopic images of the fabricated structure sequentially deformed from 0 to 13% and then released. Two typical areas (I and II) are also highlighted during the cross-sectional optical characterization of the sample. As can be observed from the images, UGAs was deformed and oriented with the PDMS layers in the stretching direction mainly via two mechanism: firstly, porous structure realigned with the applied strain (I) and then crack generated for enabling the branches extensions (II) to further conformal to the applied strain under large deformations.

The characterisation shows that in the tested arrangements of UGA and PDMS layers, UGA can be stretched more than 10% strain without showing significant structural failure along the networks.

When subject to external applied tensile deformations, cracks may generated inside the UGAs networks without affecting the provision of a reliable electrical signal response to the externally applied strains. As schematised in Figure 5(c), the highly conductive UGAs initially provide continuous conductive pathway between the PDMA layers (top). When the external tensile strain is applied, microcracks generated within the UGAs, attributing to the separation of graphene sheets (middle), can lead to an increase of the resistance of the sample due to the decrease of contacting area between conductive materials. However, when the applied strain is released (bottom) the UGAs can recover to its original shape with the help of PDMS, leading to the reconnection of the breaking branches and the recovery of the bulk resistance of the UGAs/PDMS initial structure.

EXAMPLE 5

Electromechanical characterisation of UGA/PDMS strain sensors - no groove

The piezoresistive electromechanical sensing performance characterization of UGAs/PDMS sensors were mainly carried out by monitoring their resistance changes via a potentiostate synchronously while subjecting the sensors into different tensile tests. Electric characterisation of the strain sensors was performed by a potentiostat with a model of 446. Alternatively, a digital storage oscilloscope (Tektronix TPS 2024) was also used to record the resistance changes of the strain sensors under deformations. The sampling rate was set at 5000 s ¹ . During the cycling of the stretching of the synthesized strain sensor, a Chronoamperometry was used to provide a stable potential at IV and to measure the obtained current changes between the two conductive wires of the device. For dynamic strain sensing tests, a gate resistor, possessed the similar resistance value (i.e. 1 kO, 2 kO, etc.) to the tested strain sensor, was connected in series to the strain sensor to convert the resistance changes into voltage changes to allow the signal recorder recording the voltage changes of the strain sensor when a fixed voltage was input. The voltage divider circuit used for the testing contains a surface mount gate resistor and the tested strain sensor, allowing the convert of the resistance changes of the strain sensor under dynamic tensile tests to a voltage variation. The voltage output from the strain sensor can be derived as: V strain sensor =

K sir am sensor

- V input R Resistor J where V input refers to the resistance of the tested strain sensor and the applied gate resistor, respectively.

Figure 6 shows typical resistance change measured on UGAs/PDMS sensors, having PDMS layers with different thicknesses of 500 pm (a, b), 200 pm (c, d) and 100 pm (e, f) in response to different applied tensile strains - 10% (a, c, e) and 100% (b, d, f). The applied strain rate was set at 0.5 mm/s. The assembled UGAs possessed density of 1 mg/cm ³ . The interactions between UGAs and PDMS were generated via a 2 min oxygen plasma surface treatment of PDMS layers.

Data in Figure 6 demonstrates the electromechanical performance of UGAs/PDMS sensors when being assembled from PDMS layer (no groove) with different thickness of 100 pm, 200 pm and 500 pm . The UGAs/PDMS sensors are shown to provide linear electric response to the externally applied strains. The resistance was measured to increase and decrease linearly with the applied loading and unloading strains. However, sensing capability was shown to be adversely affected by decreasing thickness of the PDMS substrates.

As can be seen in Figures 6a, c and e, all the sensors provided reliable and reversible strain- induced resistance changes in response to 10% applied strain, showing full recovery of the resistance under the stretching/releasing cycles. However, as the PDMS layer thickness approaches lower values (i.e. 100 pm), the applied strain and the output electric signal were observed to deviate. As can be seen in Figure 6e, the output electric signal of the sensor with 100 pm -thick PDMS-layers shows a protruding resistance changes behaviour. This indicates that the electrical resistance of the sensor increases faster than the applied strain when compared to the others (Figures. 6a and c), indicating the non-uniform crack formation inside the layers.

The relatively non-uniform stress and crack controlling mechanism of the thin-layer-PDMS assembled sensor was further proved by subjecting the sensor into a larger tensile strain deformation (i.e. 100% strain). UGAs/PDMS sensors assembled from thicker (200 pm and 500 pm) PDMS layers show reversible deformation of the UGAs with reliable sensing response (Figures 6b and d). However, Figure 6f demonstrates poor electric response of the thin-layer-PDMS (100 pm) sensor when being stretched up to 100%.

EXAMPLE 6

Electromechanical characterisation ofUGA/PDMS strain sensors with groove

The electromechanical characteristics of UGA/PDMS sensors made with a groove on each PDMS layer (as described in Example 2) were tested according to the same procedure described in Example 4.

Figure 7 illustrates the electromechanical performance of UGAs/PDMS sensors assembled with the PDMS layers (with thickness ti of 400 pm - see Figure 3) each having in-built grooves with different thickness ti of 100 pm and 200 pm (with reference to thickness Ϊ2 in Figure 3). As such, the UGA sandwitched between the layers had a thickness of 200 pm and 400 pm, respectively (i.e. double the depth of each groove). The data relates to sensors with grooves with different thicknesses of 200 um (a, b), and 100 um (c, d) in response to different applied tensile strains at 10% (a, c) and 50% (b, d) respectively. The applied strain rate was set at 0.5 mm/s. The assembled UGAs possessed density of 1 mg/cm ³ . The thickness of the PDMS layers was 400 um. The interactions between UGAs and PDMS were generated via a 2 min oxygen plasma surface treatment of PDMS layers.

The relative change in resistance of UGAs being pressed into 200 pm and 400 pm total groove thickness was around 0.23 and 0.20 respectively (Figure 7a and c), which are similar to those of the 100 mih-PDMS-layer assembled without groove (~ 2) (Figure 6e), when being stretched by 10%. This indicates that an expansion of the UGAs in between the thin soft PDMS substrates. Moreover, by comparing the gauge factor (GF), which is the ratio of the relative change in resistance to the applied strain, of 200 pm-grooved-400 pm-PDMS- substrate assembled sensor (Figure 7a) with the 200 pm-PDMS-substrate assembled sensor (Figure 7c), a significant incrase from 1.75 to 2.7 can be noticed. This sensitivity enhancement of the sensor under lower strain range deformations is in consistent with the 100 pm-PDMS-layer assembled sensor that the less pre -compression upon the UGAs networks within the sensor structure can induce more effective crack-induced resistance changes of the conductive networks.

Effect of groove depth / UGA compression

The amount of pre-compression upon the UGA within the sensor can tuned by modulating the groove depth, and/or by confining larges samples of UGA within a given groove. Figure 8 demonstrates how would the amount of pre-compression upon the UGA affect the electromechanical performance of the sensors. Figure 8 shows typical reletive change in resistance of UGAs/PDMS sensors assembled with UGAs with different densities of a) 12 mm and b) 8 mm in response to the external applied 10% strain. The applied strain rate was set at 0.5 mm/s. The assembled UGAs possessed density of 1 mg/cm3. The thickness of the PDMS layers was 400 um having an in-built groove with thickness of 200 um. The interactions between UGAs and PDMS were generated via a 2 min oxygen plasma surface treatment of PDMS layers.

It is seen in Figure 8a, when a very thick piece of UGA (i.e. 12 mm) is squeezed into the groove, although cyclic resistance changes are observed, the linear relation of the input/output signal is degraded. In contrast, when a 8 mm-thick-UGAs is squeezed into the groove, the resulting sensor displays highly linear response between the input and output signals (Figure 8b). This issue can be mainly ascribed to the less effective stress transfer from the PDMS layer to the UGAs networks when the UGAs was too thick that the limited surface interaction could not afford the efficient shear stress for achieving the uniform stretching of the enclosed UGAs. Besides, the reduction of the resistances during the cyclic loading/unloading cycles also proved the sliding occurred at the interface that the assembled UGAs was not deformed with PDMS layer when being stretched.

EXAMPLE 7

Static Electromechanical Characterization of the Optimised UGAs/PDMS Sensors

To detailed demonstrate how would the strain and the resistance interplay during the tensile deformation of the sensor, the electric resistance of the UGAs/PDMS sensors was measured while a constant strain or stepped strain was applied to the samples (Figure 9). Figure 9 relates to relative change in resistance (dot curve) of UGAs/PDMS sensor in response to stepped tensile strain changes (solid line curve), at one-step change of strain from 0 - 50% at the speed at 1 mm/s (Inset: close-up of signal overshoot) during tensile. Inset: close-up of the relaxation-tensile transition point. The sensor resistance changed immediately to the externally applied tensile strain, and remained at the new output value as the strain was maintained. The insets of Figure 9 illustrate that signal overshooting is negligible. In particular, the UGAs/PDMS sensor exhibited negligible overshooting behaviour (< 1 %) with fast recovery time within 5 s, indicating that the viscoelastic behaviour of PDMS caused less effect onto the electromechanical performance of the composite.

As shown in Figure 9, the sensor only took around 5 s to reach a electromechanical static state within the sensors and then provide unchanged electric performances over a long time period. The resistance of the UGAs/PDMS sensor remained almost unchanged until the strain was further changed.

EXAMPLE 8

Dynamic Electromechanical Characterization of the Optimised UGAs/PDMS Sensors

To further confirm the dynamic electromechanical performance of the oxygen plasma treated UGAs/PDMS sensors, cyclic loading/unloading deformations of the as-prepared sensors with varied using strain levels from 10% up to 100% for up to 1000 multiple cycles at relativle lower frequency.

Figure 10 shows electromechanical properties of the oxygen plasma treated UGAs/PDMS sensor based strain sensors, with plots a, b, c, and d showing relative changes in resistances of the strain sensor in response to different periodic loading cycles under strains at a) 10%, b) 25%, c) 50% and d) 100%, respectively. All strain rates were set at 5 mm/s. As can be seen in Figures 10 a, b, c and d, the electrical responses of the strain sensor as a function of the cyclic tensile deformations were detected, where four different strains (10%, 25%, 50% and 100%) were applied. Proportion relations between the resistance changes and the applied strains were obtained with high repeatability and negligible signal delaying. Additionally, it was notable that no significant resistance drops occurred during repeated cyclic deformations, indicating the good structural stretchability achievement of the conductive network and negligible deattachment occurred between the UGAs and PDMS layers.

Figure 11 shows hystersis and durability characterization of oxygen plasma treated UGAs/PDMS sensors. Figure 1 la refers to relative change in resistance of the strain sensor being stretched by 10%, 25%, 50%, 70%, and 100%. The plot shows that all strains provoke the same relative change in resistance (i.e. the plots overlap). The original resistances of the sensor were fully recovered after releasing from the applied strains. This electric signal continuity and the reversible resistance response of UGAs/PDMS sensors to strain up to 100% demonstrated the good consistent with previous microstructural monitoring of the sensor that only reversible microcracks inside the network were generated rather than structural fractures occurred during deformations. Additionally, the hysteresis curve also demonstrated the negligible signal variable during the loading and unloading process, which could be mainly attributed to the proper design of the sensor construction that allowing the negligible viscoelastic behaviour effects caused from the polymer matrix to the conductive networks. Moreover, the high linear resistance changes also indicate the relatively uniform micro-crack generation inside the assembled structures. Figure 1 lb refers to relative change in resistance versus strain for multiple-cycles: 10 cycles, 100 cycles and 1000 cycles at different strains varied from 10 % up to 50%. The plot shows that all strains provoke the same relative change in resistance (i.e. the plots overlap). The cycling tests suggest that the UGAs/PDMS sensor provide nearly unchanged strain-response for over 1000 cycles at 50% strain, indicating that the sensors can fully recover after repetitive loading/releasing cycles.

To further examine the capability of the designed construction for effectively diminishing the viscoelastic effect brought in by the polymeric elastomer, we then subjected the oxygen plasma treated UGAs/PDMS sensors into a high-frequency (> 5 Hz) dynamic tensile testing. A periodical tensile test (with frequencies ranging from 0.028 Hz to 180 Hz) was carried out synchronized with the electrical conductivity monitoring. The input tensile strains were applied by an electromagnetic shaker. The output electrical signal was recorded simultaneously by a laser detector for displacement measurement and a voltage divider circuit system for resistance changes recording.

Figure 12 testify to the excellent high-frequency dynamic behaviors of the oxygen plasma treated UGAs/PDMS strain sensor for all tested conditions. For all frequencies tested (from 0.028 to 180 Hz), the UGAs/PDMS sensors showed nearly instantaneous responses to the applied tensile strains with high stability and excellent reproducibility. Excellent correlations between the input and output with very small delay ratios at 0.28Hz, 5 Hz and 25 Hz were demonstrated (Figures 12a, b, and c).

As shown in Figures 12d and e, when gradually changed the amplitude of the applied deformations at relatively high deformation frequency of 15 Hz or 25 Hz, good electric response of the sensors to the varied strains were also be able to be maintained within this structure. In that regard, the plot of Figure 12d shows the resistance changes of the strain sensor induced by a gradually increasing strain (75% over the test duration) during a small strain-scale deformation at 25 Hz. Also, Figure 12e shows reliable and instantaneous signal response of the UGE/PDMS composite based strain sensor to a fast-changed cyclic strain at around 15.5 Hz when a 19% pretension was applied. Furthermore, even at an extremely high frequency up to 180 Hz, the output signal recorded still followed the applied distinguishable signal waveforms well (Figure 12f).

EXAMPLE 9

Use of sensor to monitor finger motions

Bending and straightening of human fingers were sensed and monitored by a strain sensor in accordance to the invention. When applied on a user’s finger, the sensor returned upwards and downwards curves of relative resistance changes, as shown in Figure 13. The measured resistance changes followed quite well the different bending angles and bending speed of the user’s finger. As shown in Figure 13b, at a relatively fast finger bending mode with frequency up to 5 Hz, the maximum relative resistance change values of the strain sensors increased significantly from 0.4 up to 1 with the increase of the finger’s bending angles from 45° to 135°. The high linearity between the input stain and output signal of the UGAs/PDMS sensors enabled the possibile uses of such devides for preciesly monitoring and quantifying the bending status of the human fingers even at relatively high motion speed (> 5 Hz).

EXAMPLE 10

Experimental and equipment setup for monitoring skeletal muscle activities

All the tests were performed with a target temperature of the upper limb of > 32 C°. If required, the limbs were warmed with a heat pack. sEMG was used as the reference to validate the sensing performance of the sensor for skeletal muscle activity detection.

Sensors in accordance to the invention and sEMG electrodes were placed on the muscle group being studied, after thoroughly cleaning the skin. The sensor was placed in between the reference and active electrodes of the sEMG over the midportion of the muscle under study. The reference electrode was placed on the non-adjacent tissue near the target muscle group. The active electrode was placed as near as possible to the motor point. The electrical signal of sensor signal and the raw sEMG were recorded synchronously by a computerised acquisition system with the assistance of a PowerLab 2/26 and a digital potentiostat 466 System with a sampling rate set at 2000 s ¹ .

Two types of sEMG, including the Bio Amp FE231 and a MyowareTM Sensor AT-04-00, were used here for different purposes: the three-lead sEMG (Bio Amp FE231, Signal channel sEMG with a PowerLab 2/26) was used as the primary sensor for detecting electrically evoked muscle activities; an alternative sEMG (MyowareTM Sensor AT-04-001) with an in-built rectified-integrated (AKA the EMG’s envelope) was used for voluntary cyclic muscle contraction detections for easier signal analysations and comparison.

EXAMPLE 11

Use of sensor to monitor voluntary muscular contractions

Voluntary biceps muscle activities were monitored by banding a sensor in accordance to the invention around the arm right onto the belly area. The subject wore the sensors and sat on a chair with a 90° anteflexion between the arm and the forearm. A weight was hung on the wrist via a soft belt. Cyclic muscle contraction and relaxation were performed. During the muscle contractions, the subject was asked to hold the weight and keep the forearm/arm angle at 90°, with the palm turned towards the shoulder. There was to be no bending at the wrist. This body position can largely minimise the error in measurement caused by the contraction of the adjacent muscle groups (i.e., triceps muscle group) which can also alter the arm circumference changes during contractions. The subject then produced a sustained contraction of the biceps muscle at different forces at 20% MVC, 40 % MVC, 60 % MVC and 80 % MVC for several seconds, respectively. After each test, 1-minute of muscle relaxation was allowed.

During the cyclic voluntary muscle contraction detection, the MyowareTM Sensor (AT-04- 001) was used and to detect the muscle activities synchronously with the sensor. The sEMG is designed to be used directly with a microcontroller which can provide either the primary raw output EMG signal and an alternative amplified, rectified, and integrated EMG signal output (namely EMG linear envelope). The processed EMG signal was commonly used to interpret the muscle mechanical activity from the EMG signal for easier comparisons to MMG signals. However, the complex data processing procedure could cause delay and show varied signal latency between sEMG and sensors, measured as 15 ms.

Figure 14 shows resistive monitoring of muscular activities induced by biceps contractions using a UGAs/PDMS strain sensor in accordance to the invention. Figure 14a shows a photograph of the placement of strain sensor onto the subject bicep muscle of the subject. Figure 14b show multiple biceps relaxing/contracting cycles as picked up by the sensor. The top curve in Figures 14b relate to the signal output of the strain sensor, while the bottom curves relate to the response of the same movements as picked up by a commercial sEMG. Both strain sensor and sEMG were attached onto the same biceps muscle group. Figure 14c shows the signal output of the strain sensor under muscle relaxation (bottom) and muscle contraction (top) during muscle contraction.

Figures 14b demonstrate the cyclic muscle contractions performed by the subject monitored by the UGAs/PDMS strain sensor. sEMG, which is the one of the most common and reliable techniques for detecting muscle activates, has been used as a reference, and measured the muscle electrical activity that occurs during muscle contractions for characterizing the muscle activities. To achieve the high consistency of the comparative tests and demo the practicality of the strain sensor, sEMG and UGAs/PDMS strain sensor were placed onto a same bicep muscle with two data recording systems synchronized together. As can be seen in Figures 14b and c, during multiple biceps contracting/relaxation cycles both sensors provided instant response to the muscle contractions with different contracting periods, frequencies and amplitudes.

When zoomed in the time period for muscle contraction into 1 seconds, significant differences obtained from the recorded data of the strain sensor during muscle relaxation and contraction could be observed (Figures 14b and c). As can be seen in Figure 14b, the strain sensor presented nearly unchanged relative resistance changes at around 0.03 during the muscle relaxing status, indicating the strain sensor was tightened up on the bicep muscle groups with a pre-strain applied onto the strain sensor. When the bicep muscle was contracted the baseline of the relative resistance changes of the strain sensor was raised up to around 0.8 due to the whole muscle shape changes with some further resistance variations from 0.01 up to 0.15 of the absolute values for each small peak. This result illustrated an interesting applications of strain sensors for monitoring the skeleton muscle activities at relative high frequency ranges.

To further demonstrate the capability of the strain sensor, a detailed comparison between the signals recorded from the synchronised strain sensor and sEMG was analysed. Figure 15a shows the typical signal responses of the strain sensor (top) and sEMG (bottom) during the bicep muscle contraction. Basically, EMG measures the electric potential that generated by the muscle cells and induces the muscle fibre contractions. The value of each peak varied with the contracting forces and the number of the fibre being detected. By rectifying and integrating the sEMG signal, a sum-up the electric potentials of the detected fibres can be presented. Although it is still not well understood, the area underneath the rectified and integrated sEMG curves were reported to be referred as force indicators, which demonstrates a sum-up behaviour of the muscle contractions. As our strain sensors also monitored the sum-up lateral thickness changes of the muscle groups, in this case, we compared the strain sensor signals with the rectified & integrated sEMG signals.

Figure 15b shows the relative change in resistance of the UGAs/PDMS strain sensor in response to different forces produced by isometric muscle contractions which are defined as the different percentages of the maximum volume contraction (MVC) during the voluntary muscle contraction.

Apart from detecting the mechanical deformation of muscle groups during contractions, we find that the UGAs/PDMS strain sensor can monitor the muscle activities by correlating it resistance changes to the force exertion, which is of particular interests to applications such as muscular fatigue detection or muscle injuries prevention. It is worth noting that the primary function of skeletal muscle contraction is to produce force. By banding the UGAs/PDMS strain sensor around the muscle group, its resistive changes decrease or increase accordingly with the change of the contraction forces. In that regard, Figure 15b shows the relative change in resistance of the UGAs/PDMS strain sensor in response to different forces produced by isometric muscle contractions which are defined as the different percentages of the maximum volume contraction (MVC) during the voluntary muscle contraction. The numbers and amplitudes of the sub-peaks of the UGAs/PDMS strain sensor signal during the muscle contractions are all increased with the contraction forces, which may be possibly caused by the increasing number of muscle fibre contraction and twitching for producing larger forces. Besides, signal latencies have also been observed between the onset of the UGAs/PDMS strain sensor and the force and the latency spans from the cessation of muscle activation to muscle return toward their pre contraction status which are attributed mainly to the complex muscle contraction and relaxation mechanisms.

EXAMPLE 12 Use of sensor to monitor stimulated muscular contractions

To provide predictable muscle contraction behaviours, the biceps muscle was evoked by an electrical stimulator (Dantec Keypoint G4 Workstation, EMG machine) with standard electrodes as used for routine nerve conduction studies. Controlled electrical stimuli produced a peripheral nerve axon action potential that triggers the target muscle’s contractions. The UGA PDMS strain sensor was fixed with two non-elastic soft bands at the two ends with the adhesive rubber pastes and banded onto the arm. The sensor was placed over the belly of the biceps muscle group as near as possible to the motor point, to optimise the signals that could be recorded from the skin. The sensor was tightened up on the arm with a small pre-strain (~ 1%) applied. The sensor can measure the circumferential changes of the arm. The study was also repeated, recording over the wrist flexor muscle group.

The electrical stimulation pulses were set at 1 Hz, 3 Hz, and 5 Hz, for a period of 2 seconds. For each test, a 5-minute rest was given to allow muscle relaxation. Higher stimulation frequencies (> 5 Hz) were not performed as they may cause strong pain to the subjects. Thus, evaluation of the sensor activity at higher frequencies was not attempted. Additionally, even under the lower frequency range, the experiment was to be ceased if the subject felt uncomfortable.

The supramaximal electrical stimulation intensity (defined as the stimulus intensity required to produce a maximal compound muscle action potential) was firstly determined by a qualified neurologist using the electrical stimulator. To adjust different applied stimulation intensities, the input stimulation current was changed. If required, the duration of the square wave stimulus was increased from 0.1 ms to 0.2 ms. The sensor in accordance to the invention and sEMG (Bio Amp FE231) signals were recorded synchronously during the tests.

Figure 16a shows a comparison of the recorded relative resistance changes of the UGA/PDMS strain sensor (top curve in each plot) and voltage outputs of the sEMG (bottom darker curve in each plot), which refers to the muscle action potential, in response to the evoked biceps muscle contractions at different frequencies of 1 Hz (top plot), 3 Hz (middle plot) and 5 Hz (bottom plot).

Figure 16b shows the relative resistance changes of the UGA/PDMS strain sensor in response to the evoked biceps muscle contractions when the stimulation intensity was adjusted at 50 mA (top), 20 mA (middle) and 10 mA (bottom). The stimulation frequency was set at 3 Hz.

Figure 16c shows a comparison of peak-to-trough values of the relative resistance changes of the UGA/PDMS strain sensor (square data points, left y-axis) with the peak-to-peak range changes of the recorded muscle action potential from the sEMG (circle data points, right y- axis) in response to the evoked biceps muscle contractions at different stimulation intensities from 6 mA to 50 mA.

At all the given stimulation frequencies of 1, 3 and 5 Hz, the UGA/PDMS strain sensor responds in resistance in the corresponding frequency ranges (Figure 16a). The response is nearly linear to the evoked muscle contraction within the stimulation intensity ranges examined (Figure 16b and 16c).

Surface electromyographic (sEMG) technique was also used to monitor the electrical aspect of the skeletal muscle activities simultaneously. The resistive signal pulses recorded from the UGA/PDMS strain sensor appears to in good agreement with the sEMG recordings from both the frequency and relative intensity aspects, except that a slight signal latency of around 1 - 7 ms between the UGA/PDMS strain sensor and the sEMG recordings is observed (Figures 17-18).

Figure 17 shows a typical signal recordings from the UGA/PDMS strain sensor and a sEMG for monitoring the evoked bicep muslce contraction.

In particular, Figure 17a shows the resistance change recording (top curve) and the muscle action potential recording (bottom curve) from the target biceps muscle group following a direct stimulation of the target peripheral nerve at Erb point. The supramaximal stimulation intensity was set at 50 mA and 0.2 ms duration with a stimulation frequency set at 3 Hz. The latency between the two signals were calculated from the two highlighed onset points (pointed by the arrows). The onset of stimulation is also illustrated by an arrow before the onset of the electrical response detected by the sEMG. This is due to the conduction along the terminal nerve from the point of stimulation to the synapse and then the synaptic transmission onto the target muscle. Such identified stimulation-sEMG delay observed during our experiment is around 4.5 ms.

Figure 17b shows the signal recordings from Figure 17a at a higher amplification. The UGA/PDMS strain sensor shows a negative ‘initial wave (IW)’ prior to the demonstration of the onset of the strong recording resistance changes of the strain sensor. The figure indicates the IW detected by the UGA/PDMS strain sensor. The detection of such subtle muscle fibre movements indicates the high sensitivity of the UGA/PDMS strain sensor for monitoring human sketetal muscle activities. The EMG only reflects the electrical rather than the mechanical events of the muscle contraction. During the electro-mechanical coupling effect that induce the muscle contraction, EMG only represents the propagation of motor unit action potentials along muscle fibres, which serving as the initiator of the muscle fibre shortening. Such time span between the electrical events and the actual muscle contraction is defined as the electromechanical latency (EML). Details of the measured EML by using the UGA/PDMS strain sensors are illustrated in Figure 18.

Figure 18 shows average latencies between the two output signals of the UGA/PDMS strain sensor and the sEMG when synchronously subjected to monitor the evoked biceps muscle contractions.

In particular, Figure 18a shows the recorded signal average latencies (average of 6 trails) between the two sensors when monitoring the evoked bicep muscle contractions at different stimulation frequencies of 1 Hz, 3 Hz, and 5 Hz. The stimulation intensity was set at 50 mA. Figure 18b shows the recorded average signal latencies (average of 6 trials) between the two sensors when monitoring the evoked bicep muscle contractions at different varied stimulation intensities of 10 mA, 20 mA, 30 mA, 40 mA and 50 mA. The stimulation frequency was given at 3 Hz. The figure indicates the frequency-dependence of the EMF recorded from the resistive changes of the UGA/PDMS strain sensor during the evoked muscle contractions. It is shown that the EMF decreased with the increase of the stimulation frequency. This phenomenon is likely contributed by muscle fibres types, including type I (reach the peak force at a slow speed), type Ila (reach the peak force quicker) and type lib (produce the highest amount of force within a shortest time period), may be triggered differently by the given stimulation with different frequencies and result in different respond speed to the applied stimulus.

Figure 18b illustrates the intensity-dependence of the EMF recorded from the resistive changes of the UGA/PDMS strain sensor during the evoked muscle contractions. It is shown that the EMF increased with the increase of stimulation intensity. The measured EMF from the evoked bicep muscle contraction detected by the UGA/PDMS strain sensor ranged from ~1 to 7.6 ms.

EXAMPLE 13

Wrist flexor muscle contractions

The UGA/PDMS strain sensor of the invention is also found to be responsive to the other skeletal muscle groups, such as for the wrist flexor muscle contractions. The comparative test results also resonate well with the sEMG signal (Figure 19).

Figure 19a shows the relative change in resistance of the strain sensor for monitoring stimulated wrist flexor muscle contractions at varied stimulation intensities of 25 mA, 20 mA, 12 mA and 10 mA. The stimulation frequency was given at 5 Hz. The simultaneous sEMG recording at the stimulation intensity of 25 mA is shown in the bottom as a reference.

Figure 19b shows a comparison of the absolute relative change in resistance of the UGA/PDMS strain sensor (square data points) and the absolute voltage change of the sEMG (dots data points) in response to the monitored stimulated wrist flexor muscle contractions at different given stimulation intensities. The variation in these two signal outputs, indicated as error bars, were calculated from 9 multiple tests recorded by 3 different GP-laminates with similar gauge factors.

The reference in this specification to any prior publication (or information derived from it), or to any matter which is known, is not, and should not be taken as an acknowledgment or admission or any form of suggestion that that prior publication (or information derived from it) or known matter forms part of the common general knowledge in the field of endeavour to which this specification relates.

Throughout this specification and the claims which follow, unless the context requires otherwise, the word "comprise", and variations such as "comprises" and "comprising", will be understood to imply the inclusion of a stated integer or step or group of integers or steps but not the exclusion of any other integer or step or group of integers or steps.