WO2024165840A1 - Triboelectric coating - Google Patents

Abstract

A superhydrophobic nanocomposite coating (100a-e) for a wearable device (100), said coating comprising: a non-toxic hydrophobic polymer for providing a dielectric effect; a non- toxic nanoparticle filler for providing surface roughness and a further dielectric effect; and a non-toxic hydrophilic polymer for providing a matrix supporting the fluoropolymer and the nanoparticle filler, wherein the coating provides a triboelectric effect when sprayed onto a substrate.

Description

TRIBOELECTRIC COATING

Field of the Invention

The invention relates to a triboelectric coating, and in particular to a method of applying a triboelectric coating, and corresponding products comprising said coating.

Background

Sensorised gloves and fingertip mounted sensors have tremendous potential in a variety of healthcare engineering applications such as robotic interventions, prosthetics and humanmachine interfaces. Haptic sensors concentrating on mechanical stimuli such as pressure, force, and mechanical vibrations are typically used and offer benefits as a biosensing platform. A majority of these applications are served with the sensing hardware exploiting a battery or other portable power source for powering or in simple cable tethered mode.

Triboelectric coatings show promise in the use of sensors by enabling harvesting of otherwise wasted mechanical energy for self-powering. Recent work has introduced the idea of harvesting energy using electrochemical means to develop self-powered sensors that can be mounted on the fingertip. This potentially provides benefits as the thinness of these sensors avoids interfering with the sensor perception of the user.

However, integration of triboelectric coatings and sensorised gloves have a number of remaining hurdles. For example, an inherent ability to prevent contamination by dirt or impurities is desirable, the sensors and their components need to maintain sterility and sensor integration on a surgical glove is an important challenge and has multiple fabrication steps.

It is an object of the present invention to at least ameliorate the above defined issues with the existing prior art.

Summary of the Invention

In accordance with a first aspect of the present invention there is provided a superhydrophobic nanocomposite coating for a wearable device. The said coating comprises a non-toxic hydrophobic polymer for providing a dielectric effect, a non-toxic nanoparticle filler for providing surface roughness and a further dielectric effect and a non-toxic hydrophilic polymer for providing a matrix supporting the fluoropolymer and the nanoparticle filler. The coating provides a triboelectric effect when sprayed onto a substrate.

The superhydrophobic nanocomposite coating according to the first aspect, as a triboelectric force sensing layer which can be sprayed directly onto printed electrodes may offer numerous advantages. First, the nanocomposites may offer a wide choice of easily accessible, low-cost polymers that lends themselves to simple processing and can serve as a suitable matrix for the nanocoatings and sensors. Second, nanocomposites can be applied - just as coatings or films - on essentially any surface, as long as adhesion is managed. Finally, the introduction of nanoparticle fillers, which are used to control surface roughness, can enable tunable inclusion of additional properties, i.e., it can help add the multifunctionality outlined above. In fact, spray coated nanocomposite coatings readily impact two important properties; antibacterial activity and superhydrophobicity. These coatings may prevent biofouling, infection and moisture contamination which may be particularly important to avoid in interventional healthcare applications, since the vast majority of biofilm-associated infections are contracted in a hospital. In addition, bacterial infections and biofilms are also linked to antimicrobial resistance (AMR), because the bacteria evolve and are able to withstand the effect of medication designed to eradicate them. The initial contact with the bacteria can take place through the skin of either the healthcare worker or the patient, or through other means such as contaminated fluids. This raises the possibility of infection and biofilm formation. Biofilms on catheters or other medical devices can result in bacteraemia - clinicians handle these devices (with gloves), thus any sensors to be integrated on gloves must comply with the sterility requirements, and ideally prevent/mitigate sources of infections.

Superhydrophobic surfaces show promising anti-fouling properties, and can inhibit adherence and the initial attachment of bacteria. A selection of antibacterial nanoparticles as the filler may not only render the nanocomposite superhydrophobic, and thus antifouling, but also bactericidal. This implies that even if microbes manage to adhere despite the antifouling nature of the coating, they will be deactivated. In addition to featuring infection and adhesion control properties, the present coating also offers sensing and energy harvesting capabilities using the triboelectric effect.

Triboelectricity is a type of contact electrification which takes place when two dissimilar materials come into contact or are separated. This contact electrification, together with electrostatic induction, enables triboelectric sensors to produce a current when they come into contact or are separated/rubbed against an object. Triboelectric sensors can be self- powered due to their energy harvesting capabilities, i.e. they can serve as a triboelectric nanogenerator (TENG). TENGS have multiple operation modes: vertical contact-separation, lateral sliding, single-electrode and freestanding triboelectric-layer. Furthermore, most current TENGs are flexible (polymer-based), effectively portable and can be realised through inexpensive fabrication processes. TENG-based sensors are suitable for static and low- frequency dynamic pressures which enables them to detect different modes of contact, for example during tapping and rubbing. In addition, they can be integrated with piezoelectric nanogenerators (PENGs) to enable wide bandwidth energy harvesting. To this end, the nanocomposite coating of the present invention offers self-cleaning properties and ready integration directly on the substrate of choice, e.g. the flexible surgical gloves presented herein.

In some embodiments the non-toxic hydrophobic polymer may be a fluoropolymer, and optionally or preferably a Polyvinylidene Fluoride. The Poly(vinylidene fluoride) (PVDF) had an average Mw ~530,000, (pellets, Sigma Aldrich). From the PVDF’s safety data sheet following EU regulation 1907/2006 it is classified as non-toxic and biocompatible. Cytotoxicity tests have been carried out on PVDF for its use on medical devices, and it has shown no toxic effect.

In other embodiments the non-toxic hydrophobic polymer may be a thermoplastic elastomer, and optionally or preferably a thermoplastic polyurethane.

In an embodiment the non-toxic nanoparticle filler may comprise ZnO nanoparticles. Zinc oxide (ZnO) nanoparticles (nano powder, <100 nm particle size, Sigma Aldrich): The zinc oxide nanoparticles used for the coating were selected following the EU commission regulation standard 2016/621 , and as covered in the assessment of this standard, pose no risk of adverse effects on human skin. Zinc oxide nanoparticles’ toxicity has been extensively studied and it has been established that when applied on the skin surface they do not penetrate nor cause toxicity in the viable epidermis.

The concentration of zinc oxide nanoparticles in the coating formulation was selected such that the range of ZnO nanoparticle concentration which resulted in the highest contact angle (0A) measured was selected. Once it was established that this range of nanoparticle concentration was yielding superhydrophobic results following contact angle tests, a 10% ZnO nanoparticle concentration was chosen for the triboelectric coating since it resulted in the highest power density (106 pW/cm²).

The non-toxic hydrophilic polymer may comprise an acrylate polymer. In some embodiments the acrylate polymer may comprise Polymethyl methacrylate. Poly(methyl methacrylate) (PMMA) (average Mw ~120,000 by GPC, Sigma Aldrich): PMMA’s cytotoxicity has been studied and the results show that when in contact with human tissues it does not present a biological threat and is non-toxic and biocompatible. In another embodiment the acrylic polymer may comprise poly ethyl cyanoacrylate.

In an embodiment the coating comprises a thickness of between 1 pm and 100 pm. Optionally the thickness may be between 20 pm and 60 pm. Said thicknesses broadly correlate to the number of spray passes performed to apply the coating to a substrate. A thickness sufficient to yield the optimal/stable triboelectric sensor output is required. This is important to ensure that changes in the stiffness level/contact area of different tissues (for example to detect anal sphincter injury or sutures in a baby’s head) are detected in a repeatable manner. To detect changes in stiffness levels, the sensors need to be responsive to deformation and deformation rate changes, and as a result need to be flexible. Furthermore, thinness must also be ensured to allow perception to be maintained when the coating is applied to a wearable device that requires the user to also detect morphology changes through the surface. As an example, if applied to a surgical glove the coating needs to allow the glove to be flexible whilst still providing the triboelectric sensing effect.

According to a second aspect there is provided a triboelectric generator comprising a substrate having a nanocomposite coating according to any embodiment of the first aspect.

According to a third aspect there is provided a wearable device comprising the triboelectric generator of the second aspect. In an embodiment the substrate may further comprises an electrode beneath the coating.

According to a third aspect there is provided a method of applying the coating of any embodiment of the first aspect, wherein the method comprises the steps of: dissolving the non-toxic hydrophobic polymer and the non-toxic hydrophilic polymer in a solvent; combining the dissolved non-toxic hydrophobic polymer, the non-toxic hydrophilic polymer and the non- toxic nanoparticle filler in a sprayable dispersion ; and spraying the coating onto a substrate.

Spraying the coating may comprise a number of passes, each pass depositing a thickness. In some embodiments the number of passes may be between 2 and 1 5. Each pass may deposit approximately 1 .5 to 7 pm.

In an embodiment the number of passes is between 6 and 10 passes. In an embodiment each pass may deposit a thickness of approximately 2 to 4 pm per pass.

A minimum number of spray passes is desirable because a lower number of passes results in an erratic and variable sensor output current. This may be due to microscopically non-uniform coating thickness and the electrodes not being fully covered by the coating. Once a threshold number of passes and thickness of the coating was surpassed, a more stable operation range is obtained, where results are consistent and the sensor or triboelectric generator produced by the coating is able to accurately detect changes in stiffness levels. This is advantageous to detect slots of different materials, sutures in phantoms and the injuries, such as anal sphincter injury detection in ex vivo pig studies. However, beyond a certain threshold number of passes or thickness, such a high number of spray passes results in coatings that are stiff and flaky, which again resulted in lower and less stable current outputs. The reduced flexibility of the coating layer leads to a lower contact area and deformation which reduces the magnitude of the output peaks in current for the triboelectric sensor. Consequently, in an embodiment an optimal range of spray passes (6 - 10 spray passes, for which the triboelectric properties are most suitable may be used. It can be appreciated that the number of passes and/or thickness of the layers may differ as the non-toxic hydrophobic polymer, non-toxic nanoparticle filler or non-toxic hydrophilic polymer change.

In an embodiment method the non-toxic hydrophobic polymer and the non-toxic hydrophilic polymer are separately dissolved in a solvent. The non-toxic hydrophobic polymer may comprise a 5 to 20, 6 to 10, or preferably an 8 wt.% weight stock solution and the non-toxic hydrophilic polymer may comprise a 10 to 30, 15 to 25, or preferably a 20 wt.% weight stock solution. The solvent may be dimethylformamide, dimethyl sulfoxide, ethyl acetate, butyl acetate or any combination thereof. More generally, any suitable polar aprotic solvent (or combination thereof) may be used, such as (but not limited to): acetone, acetonitrile, dichloromethane, dimethylpropyleneurea (DMPU).

According to a fourth aspect there is provided a sensorised glove for biosensing, said glove comprising: one or more printed electrodes on fingertips of the glove; and a triboelectric coating according to any embodiment of the first aspect, deposited on the electrodes to form one or more triboelectric sensors.

The triboelectric single-electrode/coating-based sensors, such as employed with the aspects defined above may allow sensing of the tactile forces during normal manual contact/rubbing and demonstrate an ability to detect stiffness change. Unlike other coatings that rely on tapping signals, the present coating is able to detect stiffness changes from rubbing signals, such as using sensorised gloves. This may be more suitable for future uses in interventional applications where a clinician is often space constrained. This may be particularly useful for procedures where imaging/vision-based approaches are difficult to use and tactile feedback is an attractive alternative. The tactile force sensing and stiffness change detection may also be feasible with a second sterile glove on top. As a specific example of a clinical application, a sensorised glove such as that of the fourth aspect could be used for vaginal and rectal examination and the resulting tactile signals could aid diagnosis and assessment of maternal anal sphincter injuries after vaginal birth. This is particularly important as up to 12% of women can sustain such an injury, which can have long-lasting, devastating effects on quality of life. Effects include faecal incontinence, recurrent urinary tract infection, and fistula formation. Swift diagnosis and repair at the time of injury is required in order to restore normal anatomy. Nearly 40% of obstetric anal sphincter injuries are missed on primary assessment, and, if missed, can lead to worsening faecal incontinence.

The spray application of the present triboelectric nanocomposite coating lends itself to versatility of applications. For example, it can be applied/used as a simple strategy to posttreat commercial sensors and to develop a prototype to demonstrate enhanced energy generation. This may add features to the sensing capabilities of the commercial film, enabling the detection of two different modes of contact (tapping and rubbing).

The present invention describes an easy to fabricate, cost-effective triboelectric nanocomposite coating, that has been exploited, with both sensing and energy harvesting capabilities. In embodiments, integrating the nanocomposite into surgical glove fingertips or other typical clinical devices has strong potential for a wide range of clinical applications by enabling real-time force readings and detection in changes in stiffness of tissues. As an example, a surgical glove with fingertip based ultrathin sensors is likely to be favourable from the standpoint clinical translation; a glove identical to the one worn in usual procedures and surgeries could be used, with the prominent addition of embedded coatings with sensing capabilities.

Due to the superhydrophobic and antibacterial nature of the triboelectric nanocomposite, biofouling and biofilm formation can be avoided or at least mitigated.

Furthermore, the electrical outputs of embodiments having the proposed glove-based health monitors will not suffer from fluctuations upon changes in environmental conditions such as temperature and air humidity which may lead to maintained accurate sensing. Additionally, the energy harvested using the proposed films may be sufficient to drive many small electronics and opens the route for this tactile system to be fully self-powered in the future. All in all, this nanocomposite comprises a promising route in terms of multifunctionality rationally targeted towards healthcare applications.

Embodiments of the aspects may comprise any element of any embodiment of the other aspects of the invention.

Brief Description of the Figures

The invention is described in further detail below by way of example and with reference to the accompanying drawings, in which: figure 1. a) Photograph of the sensorised surgical glove prototype. Outputs of triboelectric sensors applied on substrates with different stiffness, b) glass, c) PDMS (mixing ratio 10:1 elastomer to curing agent) and d) ecoflex. Each fitting curves had R-value of > 0.94; figure 2. a) Schematic representation of the triboelectric basic freestanding rubbing test, b) Results obtained from the freestanding triboelectric-layer mode test; figure 3. a) SEM image of the cross-section of the triboelectric sensor, showing the thicknesses of the coating and the electrode, b) Picture of a water drop on the triboelectric coating, with 0A of 158° and A0 of 6°. c-f) Fluorescence microscopy images showing c) cultured control Human Dermal Fibroblast (HDF) cells (without sample) and d), e) and f) are cultured HDF cells treated with samples for 24, 48 and 72 h extracts respectively. Live cells are stained green and dead cells red. g) MTT assay results showing 75 ± 4 % HDF cell viability after 72 hours, h-l) Kirby-Bauer antibacterial test results showing no growth of neither S. aureus (h, i) nor E. coli (j, k) in the area where the sensor was placed. I) Kirby-Bauer antibacterial test results showing no growth of S. agalactiae (Group B streptococcus) in the area in which the sensor was placed (constituting the zone of inhibition); figure 4. Energy generation characteristics of the triboelectric sensors, a) Output power density at different loads, b) Rubbing test at different speeds, c) Open-circuit voltage, d) Photograph of LEDs powered by harvested energy. figure 5 a). Temporal maximum power generated by tapping on the commercial piezoelectric sensor, b) Temporal maximum power generated by tapping on the triboelectric- enhanced piezoelectric sensor, c) Schematic of the triboelectric-enhanced piezoelectric commercial film prototype, d) Commercial piezoelectric sensor’s response to controlled rubbing motion, e) Triboelectric-enhanced piezoelectric sensor’s response to controlled rubbing motion; figure 6. Change in stiffness tests using the 4 slots and controlled motorized stage setup, a) Tapping test performed on slots 1 and 3 (ecoflex) and 2/4 (PDMS) showing larger peak width for the harder material, b) Rubbing test performed through slots 1 to 2 (ecoflex to PDMS); figure 7. a) Schematic representation of the slots setup for the elastography tests (1 and 3 correspond to ecoflex and 2 and 4 to PDMS). b) Schematic illustration of the fingertipbased tactile sensor glove for surgical applications, c) Schematic illustration of the fingertipbased tactile sensor glove for surgical applications covered up by a second surgical glove on top. d) Comparison of peaks produced by rubbing against the 4 slots shown in Figure 4a using the sensorised glove (from left to right - blue signal, from right to left - red signal), e) Comparison of peaks produced by rubbing against the 4 slots shown in Figure 4a using the sensorised glove covered up by a second surgical glove (from left to right - blue signal, from right to left - red signal), f) 3 different tests using 3 different glove samples showing the peaks produced by rubbing against the 4 slots shown in Figure 4a using the sensorised glove from left to right, g) 3 different tests using 3 different glove samples showing the peaks produced by rubbing against the 4 slots shown in Figure 4a using the sensorised glove from right to left; figure 8. a) Obstetrician carrying out the test on the dissected ex vivo pig anal sphincter wearing the sensorised glove covered by a second surgical glove, b) Control test results on intact dissected anal sphincter, c) Test results showing accurate detection of the anal sphincter defect in the dissected sphincter, d) Obstetrician carrying out the test on an anal sphincter in a non-dissected whole pig cadaver wearing the sensorised glove covered by a second surgical glove, e) Control test results on intact non-dissected anal sphincter, f) Test results showing accurate detection of the defect in the non-dissected anal sphincter; figure 9. Median minimum current peak obtained throughout the anal sphincter detection tests on the intact sphincter vs. sphincter with a defect, a) Tests carried out on the dissected pig’s anal sphincter, b) Tests carried out on the pig cadaver’s anal sphincter; and figure 10a. Box plot showing the variability and magnitude of the current peaks recorded when carrying out a repeatable tapping test 30 times for each PVDF-PMMA-ZnO coating with varying number of spray passes. The area highlighted in green shows the optimal range of spray passes obtained; figure 10b. Box plot showing the variability and magnitude of the current peaks recorded when carrying out a repeatable tapping test 30 times on TPU-PMMA-ZnO nanocomposite coating with varying number of spray passes. The area highlighted in green shows the optimal range of spray passes obtained; figure 1 1 . Box plot showing the variability and magnitude of the measured roughness values for each PVDF-PMMA-ZnO coating with varying number of spray passes; figure 12. a) Optical microscope image of the PVDF-PMMA-ZnO coating (1 spray pass), b) Scanning electron microscope (SEM) of the PVDF-PMMA-ZnO coating (1 spray pass), c) Scanning electron microscope (SEM) of the PVDF-PMMA-ZnO coating (1 spray pass) at higher magnification; figure 13. a) Optical microscope image of the PVDF-PMMA-ZnO coating (2 spray passes), b) Scanning electron microscope (SEM) of the PVDF-PMMA-ZnO coating (2 spray passes), c) Scanning electron microscope (SEM) of the PVDF-PMMA-ZnO coating (2 spray passes) at higher magnification; figure 14. a) Optical microscope image of the PVDF-PMMA-ZnO coating (4 spray passes), b) Scanning electron microscope (SEM) of the PVDF-PMMA-ZnO coating (4 spray passes), c) Scanning electron microscope (SEM) of the PVDF-PMMA-ZnO coating (4 spray passes) at higher magnification; figure 15. a) Optical microscope image of the PVDF-PMMA-ZnO coating (6 spray passes), b) Scanning electron microscope (SEM) of the PVDF-PMMA-ZnO coating (6 spray passes), c) Scanning electron microscope (SEM) of the PVDF-PMMA-ZnO coating (6 spray passes) at higher magnification; figure 16. a) Optical microscope image of the PVDF-PMMA-ZnO coating (8 spray passes), b) Scanning electron microscope (SEM) of the PVDF-PMMA-ZnO coating (8 spray passes), c) Scanning electron microscope (SEM) of the PVDF-PMMA-ZnO coating (8 spray passes) at higher magnification; figure 17. a) Optical microscope image of the PVDF-PMMA-ZnO coating (10 spray passes), b) Scanning electron microscope (SEM) of the PVDF-PMMA-ZnO coating (10 spray passes), c) Scanning electron microscope (SEM) of the PVDF-PMMA-ZnO coating (10 spray passes) at higher magnification; figure 18. a) Optical microscope image of the PVDF-PMMA-ZnO coating (12 spray passes), b) Scanning electron microscope (SEM) of the PVDF-PMMA-ZnO coating (12 spray passes), c) Scanning electron microscope (SEM) of the PVDF-PMMA-ZnO coating (12 spray passes) at higher magnification; figure 19. a) Optical microscope image of the PVDF-PMMA-ZnO coating (14 spray passes), b) Scanning electron microscope (SEM) of the PVDF-PMMA-ZnO coating (14 spray passes), c) Scanning electron microscope (SEM) of the PVDF-PMMA-ZnO coating (14 spray passes) at higher magnification; figure 20. a) Optical microscope image of the PVDF-PMMA-ZnO coating (16 spray passes), b) Scanning electron microscope (SEM) of the PVDF-PMMA-ZnO coating (16 spray passes), c) Scanning electron microscope (SEM) of the PVDF-PMMA-ZnO coating (16 spray passes) at higher magnification; figure 21. a) Optical microscope image of the PVDF-PMMA-ZnO coating (18 spray passes), b) Scanning electron microscope (SEM) of the PVDF-PMMA-ZnO coating (18 spray passes), c) Scanning electron microscope (SEM) of the PVDF-PMMA-ZnO coating (18 spray passes) at higher magnification; figure 22. a) Optical microscope image of the PVDF-PMMA-ZnO coating (20 spray passes), b) Scanning electron microscope (SEM) of the PVDF-PMMA-ZnO coating (20 spray passes), c) Scanning electron microscope (SEM) of the PVDF-PMMA-ZnO coating (20 spray passes) at higher magnification; and figure 23 shows a plot of coating thickness vs spray passes.

It should be noted that the Figures are diagrammatic and not drawn to scale. Relative dimensions and proportions of parts of the Figures have been shown exaggerated or reduced in size, for the sake of clarity and convenience in the drawings. The same reference signs are generally used to refer to corresponding or similar feature in modified and different embodiments.

Detailed Description

Figure 1 a shows a photograph of a sensorised surgical glove 100 with screen printed silver ink electrode/interconnect 120 and sprayed nanocomposite triboelectric coatings 110a-110e at each of the five glove fingertips. The coatings are applied to electrodes 120 printed onto the surgical gloves. The electrodes are typically silver electrodes or interconnects that allow for electrical signals generated by the triboelectric coatings to be measured by an external detector. The electrodes and interconnects 120 were directly screen printed on the glove; the ink used for this kind of printing has substantially low curing temperature (minimum 60° C). The stencil designs were created using Eagle AutoCAD software. Following printing and curing of silver electrodes and interconnects, the triboelectric coating 110a-110e was sprayed on the fingertips using a custom stencil.

The triboelectric coatings 110a to 110e form tri bole lectric sensors on the fingertips of the glove 100 that detect deformation of a surface in response to applied force from the sensor, as will be described in more detail below. This is broadly shown in figures 1 b to 1 d, where the outputs of the tribolelectric sensors (in current nA) in response to applied force (in Newtons) is shown. The initial slope of the fitted curves (see Figure 1 b-d) increases as the substrate becomes softer (1.78 nA/N for ecoflex, 1.04 nA/N for PDMS and 0.93 nA/N for glass). This is understood to be due to the decrease in deformation (and thus the contact area) with increasing stiffness. As a material becomes softer, it deforms more for a given applied force, resulting in a larger slope. This realization enables detection of a stiffness change.

The inset of figure 1 a shows the hierarchically rough morphology created by the zinc oxide (ZnO) nanoparticles. The hierarchical roughness of our coatings not only enhances the antiwetting characteristics but also the contact area and thus triboelectrification. Furthermore, surface roughness can prevent full contact between the electrode and the coating, since some areas remain untouched, making it more sensitive to the applied force.

The triboelectric nanocomposite used for the coatings 110a-110e was formulated based on a fluoropolymer/acrylic blend to which nanoparticle fillers are added, resulting in a sprayable dispersion that can yield a superhydrophobic coating. The fluoropolymer used was PVDF due to its dielectric and known biocompatible nature. To overcome the poor adhesion of PVDF due to its hydrophobicity it was blended with PMMA (an acrylic). The dielectric nanoparticle fillers (ZnO) were added to increase the surface roughness. The resulting nanostructured morphology is important since surface roughness enhances the contact area and thus triboelectrification. The process of nanocomposite formation was adapted from our previous work, which used cyanoacrylate based acrylic instead. Briefly, PMMA and PVDF were dissolved separately in DMF, creating an 8 wt.% weight and 20 wt.% stock solutions, respectively. To dissolve the PVDF pellets, a temperature of 70 °C was needed. The sprayable dispersion comprised of 8.3% of the PVDF solution, 13.9% of the PMMA solution, 10% of ZnO nanoparticles and the balance being acetone (67.8%). The mixture was magnetically stirred for 1 h at room temperature, followed by bath sonication for 15 minutes and a second 15-minute magnetic stirring at 1000 rpm to obtain a well-dispersed composition before spraying directly on an electrode layer. An IWATA LPH-80 (1 mm nozzle) spray gun was used to apply the coatings. The coated electrodes were annealed for 2 h at 125 °C in open air.

To characterise the triboelectric properties, the coating was tested first as a standard freestanding triboelectric-layer. The triboelectric signals from the coating were recorded using a spring-loaded board as shown in figure 2a, with an electrode separation of 3 mm. This type of test comprises moving from side to side (in a sliding fashion) a freestanding dielectric layer 210 on two metal films 220, 220b on a polyimide layer 230, connected to an electronic load that serve both as the counter triboelectric material and as the electrodes, as can be seen in the series of images of figure 2a. This test effectively shows whether a film is triboelectric, since it creates electricity from charge transfer from friction as the layer 210 moves. The results are shown in figure 2b, which shows a series of current spikes generated by the triboelectric effect.

For the rest of the triboelectric tests, the sensor was fixed on a precision 3-axis motorised stage (Aerotech, high resolution: <2 nm, repeatability: 75 nm and accuracy: 250 nm) which was driven using G-code. The assembly ensured testing with reproducible application of contact forces on the sensor. The contact probe was a 3D printed polylactide (PLA) malletlike plunger. The plunger tip area was 3 cm x 3 cm to cover the surface of the sensors. The plunger was fixed to an optical table, while the motorized stage mounted sensor was moved up and down to simulate ‘tapping’ (normal contact) and then laterally to mimic the ‘rubbing’ (shear contact). The open-circuit voltage (VOC) was measured using an oscilloscope (MSO4032 Tektronix) with a 10 MQ load in parallel with the sensor and the short-circuit current (ISC) was measured by connecting either a digital multimeter (SDM3055) or a Keithley electrometer (6517B) in series. The Voc measurements were carried out in contactseparation mode by manually tapping on a setup comprising two small boards connected to each other by means of 4 parallel springs. The triboelectric coating (2 cm x 2cm) was placed on one of the boards while a layer of PES (also 2 cm x 2 cm in size) was placed on the other board to make contact with the coating. PES was chosen due to its opposite rank with respect to our coating in the triboelectric series. The energy harvesting performance was additionally investigated using different load resistances (103-109 Q). The effective power density was determined from: where A is the effective surface area, P the output power, V the voltage and RL the load resistance, respectively.

Output current from triboelectric sensors increases with the applied force, up to a threshold due to the limited deformability of the material. However, by spraying the triboelectric coating on different substrates with varying stiffness values, we were able to tune the measurable force range. The harder the substrate, the lower the measurable force range achieved by the sensor (Figure 1 b-d). Each sensor was calibrated individually by plotting the sensor output current against force values applied. For sensor calibration, the test setup was extended to include 25-N M4-5U digital force gauge. The force gauge was mounted on the contact probe, so that when the sensor touched it, the force created was recorded by the gauge. The force gauge module and the electrometer module worked simultaneously by means of a virtual interface developed using LabView. The resulting data was fitted using either power law or linear polynomial. For Figure 1 d, the linear fit yielded a high R-squared value of 0.94. The trends are consistent for similar triboelectric sensors in literature where linear and power function fit are typical. With an emphasis on interventional healthcare, a focus on the range of 0 to 20 N was used. The typical force applied across general surgery is 4.67 N (mean of average) and 11.4 N (mean of maximum), while for obstetrics specifically they are 8.69 N (mean of average) and 10.1 N (mean of maximum).

The nanocomposite coating on the triboelectric sensor 110a is shown in figure 3a. The coating 300 is formulated to be antibacterial, superhydrophobic, triboelectric and non- cytotoxic. The coating thickness was measured using scanning electron microscope (SEM) imaging (Figure 3a), which showed a thickness of 30 pm. Morphological characteristics of the coating were observed using a scanning electron microscope (SEM) (Zeiss, 1450XB). A small portion of the coating was directly placed on a carbon coated grid and then sputter coated with gold and observed at 20 kV. The thickness of the coating was also measured using the SEM by using a small portion of the single-electrode sensor and placing it on double 90 degree angled SEM aluminium stubs. All the constituent materials are biocompatible and non-toxic to human tissue following EU material regulations. As noted above, the superhydrophobicity of the coatings (Figure 3b) is obtained due to the low energy of the polymers together with the hierarchically rough morphology created by the zinc oxide (ZnO) nanoparticles. The hierarchical roughness of the coating not only enhances the anti-wetting characteristics but also the contact area and thus triboelectrification. The wettability was characterized by measuring dynamics contact angles, which showed an advancing contact angle 0A of 158° (Figure 3b) and contact angle hysteresis A0 of 6° which confirm superhydrophobicity and self-cleaning properties.

In order to assess biocompatibility, human dermal fibroblast (HDF) cell line was used as it is associated with tissue repairing and wound healing. Fluorescence imaging (Figure 3c-f) of HDF cells treated for different amounts of time, with extracts of the sensor material (coating) shows live cells (indicated by green fluorescence) with morphology characteristic of healthy HDF cells. The cells were cultured in Dulbecco's Modified Eagle Medium (DMEM medium) supplemented with 10% FBS, 1 % Antibiotic-Antimycotic and 2% GlutaMAX under standard cell culture conditions of 37 °C and 5% CO2. A LIVE/DEAD cell viability assay kit (Thermofisher, L3224) was used to test cell viability by fluorescence imaging. Initially, small pieces of the nanocomposite coating were immersed separately in 4 mL of phosphate buffered saline (PBS) for 24, 48 and 72 h to obtain sample extracts. Next, 50 pL of sample extract (from each day) was added to 400 pL of DMEM media containing 5x1 o⁴ cells/mL HDF cells in a 48-well plate. After three days of incubation, the cells were washed with PBS and 400 pL LIVE/DEAD assay solution (prepared by mixing 5 pL of calcein AM (Component A) and 20 pL of ethidium homodimer-1 (Component B) with 10 mL PBS) was added to each well and further incubated for 30 mins. Finally, the stained cells were observed directly using a fluorescence microscope (EVOS M5000). Biocompatibility of the sensor coating was further evaluated by MTT [3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyltetrazolium bromide] assay. The procedures and concentrations of seeding cells were the same as the fluorescence imaging described above. Treated cells were washed three times with PBS and the medium was replaced with 180 pL fresh DMEM media followed by addition of 20 pL of the MTT stock (5 mg/mL) solution and further incubated for 4 h. The medium was then replaced with 200 pL DMSO and was incubated again for 20 min to dissolve the formazan crystals. A microplate plate reader (TECAN, 200 Pro, Switzerland) was used to measure the absorbance at 570 nm. Cell viability percentage was calculated with respect to control (without any sample). All experiments were repeated five times. Compared to control (Figure 3c), HDF cells treated with extracts for 24 h (Figure 3d) and 48 h (Figure 3e) show a slightly smaller number of live cells, but the cell number significantly recovers with 72 h extract (Figure 3d). This may be due to the agglomeration of released ZnO particles in the 72 h extracts. Higher number of ZnO nanoparticles are released throughout the 72 h period compared to the 24 h or 48 h period, resulting in larger sized particle agglomerations. Being larger in size they produce less ROS and also are not able to penetrate cells, hence showing low cytotoxicity. HDF cell viability was further investigated by MTT assay (Figure 3g), and as expected the percentage of viable cells increased to 75 ± 4 % when treated with 72 h extract.

ZnO is well known for its biocidal activity, which does not promote microbial resistance. The antibacterial activity of ZnO has been linked to several simultaneously active mechanisms. Firstly, ROS generation leads to oxidative stress on the bacteria, secondly, cell walls are damaged due to ZnO-localized interaction and thirdly, there is membrane disruption and permeability. Finally, the toxicity is caused by the dissolved Zn ions in the bacterial cytoplasm. As a result, the ZnO nanoparticles render the coating antibacterial and have been shown to deactivate S. aureus (ATCC 43300 - gram positive), E. coli (ATCC 25922 - gram negative), S. agalactiae (gram positive - group B streptococcus which is commonly associated with maternal and neonatal sepsis. Briefly, overnight grown cultures of bacteria (10⁷ CFU/mL) were first plated on nutrient agar. Next, a small piece of the sample (15 mm x 15 mm - approximating a reasonable size of the fingertip area) coated with the nanocomposite (sensor material) was placed upside down at the centre of the agar plate so that the coating is in direct contact with the agar. The plates were then incubated at 37 °C for 24 hours before observing zone of inhibition. This is a derivation of the Kirby-Bauer test carried out to assess the antimicrobial properties of nanocomposites. As evident from Figure 3h-l, all the sensor materials show clear zones of inhibition for E. coli (Figure 3h and 3i), S. aureus (Figure 3j and 3k), and S. agalactiae (Figure 3I), which indicates no bacterial growth in the region where the coating is placed. It is also important to note that the zone of inhibition does not spread beyond the sample area. This is considered to be because ZnO nanoparticles are embedded in the coating and impart antibacterial activity through contact inhibition by ROS production. Thus, overall, the sensor material shows excellent antibacterial activity and good biocompatibility, which suggest the safe use of coatings (sensors) in healthcare application. As described below, during actual operation, for enhanced sterility, the sensorised glove 100 would be covered with a second surgical glove on top, which eliminates direct contact of sensor components with human tissue.

As shown in Figure 4a, a maximum output power density of 106 pW/cm² is obtained at a load resistance of 10 MQ. Figure 4b shows processed triboelectric signal from rubbing tests using the fingertip integrated sensors and the positioning set-up described under Methods section. The current peaks are negative and remain unaffected, at about -10 nA, by the rubbing speed; the signal peak width corresponded to the time the rubbing. Next, manual contactseparation mode measurements were performed and yielded an open circuit voltage of up to 150 V, as shown in Figure 4c. Varying electrical load were connected to the TENG sensors to measure maximum power out. The output of the triboelectric nanocomposite coating could be used to light up to 50 commercial blue LEDs 410 (Figure 4d). This shows the potential of creating a fully self-powered tactile system in the future.

In addition, the coatings 300 can be readily sprayed on piezoelectric films to enhance their power output and bandwidth (while also rendering them antimicrobial and superhydrophobic). This attests to the scalability of the sprayed nanocomposite, which is suitable for use nearly any flexible electronics platform by simply spray coating on a conductive surface. For proof of concept demonstration, the triboelectric coating 502 was sprayed directly on commercial piezoelectric sensors electrodes 504 (Figure 5a). An LDT1 -028K Piezo film was used to demonstrate the proof of concept. The coated film was annealed at 60 °C for 2 h. The leads 510a, 510b on the commercial piezo film were used for circuit connections. After sensor fabrication, the test bench and procedure covered in the previous sections was again set up to measure open-circuit voltage (Voc) and short-circuit current (Isc). Both off-the-shelf piezoelectric sensor and the triboelectric-enhanced piezoelectric sensor prototypes were subjected to controlled tapping and rubbing contact to measure outputs experimentally. A digital signal-processing algorithm was implemented to filter noise and a full-wave bridge rectifier (shown in the supporting information) was set up to compare the temporal power density from the two sensors. A synergistic effect was found in both tapping and rubbing tests when the triboelectric layer was coated on the piezoelectric film, with clear enhancements in sensitivity. The triboelectric-enhanced prototype (Figure 5b) shows peak powers that are ca. twice that from the commercial piezoelectric sensors (Figure 5c). Furthermore, it enables the detection of both normal (tapping) and shear (rubbing) contacts, as opposed to commercial piezoelectric films that only detect tapping. Commercial piezoelectric sensors can detect ‘higher frequency’ tapping contact; the triboelectric nanocomposite coating should extend its ability and enables capturing ‘lower frequency’ contacts such as the rubbing (Figure 5d-e). Figure 5d shows five negative peaks, corresponding to the first contact of the test probe with commercial sensor during rubbing and approximately zero current for the rest of the test periods. Figure 5e also shows the five negative peaks, which remain the same, and the added response of the triboelectric-enhanced sensor featuring 10 nA constant current output. The current is induced by electron transfers from the coated surface to the top electrode of the commercial film.

Following the stiffness-dependent calibration a number of tests were carried out to assess the ability of the nanocomposite sensors to distinguish changes in stiffnesses in different materials, performed both manually and using the controlled motorized positioning setup. The purpose of these tests was to use the nanocomposite sensors to distinguish changes in stiffnesses in different materials. In order to make sure that the measurements taken and the signal interpretations are not influenced by the relative position of the materials in contact with the sensor in the triboelectric series, the first tests were carried out with the sensor covered up by a surgical glove. The reasoning behind this was to not only decouple the influence of the materials position in the triboelectric series (since there would always be an isolating layer - the surgical glove - between the sensor and the materials it comes into contact with) from the influence of the materials stiffness, but also to establish the idea of feasible stiffness change detection even when the sensor is covered by a surgical glove. The controlled ‘tapping’ was carried out on 2 different materials (PDMS with a mixing ratio 10 - parts base elastomer and 1 - part curing agent and ecoflex) of different stiffnesses, PDMS being the stiffest and ecoflex the softest. Next, rubbing mode was used. The covered sensors were rubbed against 2 slots of different materials of the same height. Slots 1 and 3 are made out of ecoflex and slots 2 and 4 are made out of PDMS, with no change in height from one slot to the other. This was done by pouring and curing them on a 3D-printed mould.

Both uncovered and covered sensors on the index fingertip of the glove were rubbed against either 2, 3 or 4 slots of different materials of the same height. Slots 1 and 3 are made out of ecoflex and slots 2 and 4 are made out of PDMS, with no change in height from one slot to the other. This was done by pouring and curing them on a 3D-printed mould. These tests were performed manually by wearing the surgical sensorised glove.

These tests were performed using the controlled set up and two different modes of contact (“tapping” and “rubbing”). For a second set of stiffness detection test in S5, we used 3 slots of 3 different materials (Teflon, PDMS and ecoflex) and manually rubbed across them with a covered sensorised glove. The interconnect coming from the index fingertip sensor was connected to the digital multimeter (SDM3055) which was in turn connected to the laptop. Signals were recorded using LabView. Figure 4 shows results to this end. The sensor signals were recorded by rubbing across 4 slots filled with materials of different stiffness (Figure 4a) directly with a sensorised glove (Figure 4b) and a sensorised glove covered by a surgical glove on top (Figure 4c). The thickness of the materials in each of the four slots was kept the same to ensure repeatability. Figure 4a depicts the 4 slots used, 1 and 3 correspond to ecoflex (“softer”) and 2 and 4 corresponding to PDMS (“harder”). The resulting signal recordings are plotted in Figure 4d-e and show the distinctive peaks that form when going from one material to another, irrespective of the sensors being exposed or covered by an extra surgical glove. In these tests, negative peaks form when the sensor makes contact with a given material, and similarly positive peaks are formed when contact is released from the material (Video S2). The fact that a current peak will be created when separating from the material with which the sensor has exchanged charges through previous contact is well established in literature. In turn, current peaks created when contacting a material for the first time have also been reported and seem to be due to the contribution of flexoelectricity. Mechanical deformations caused by the contact or rubbing motion can lead to flexoelectric coupling which is reflected by the first current peak when contacting slot 1 (or slot 4 in the reversed direction movement).

The peaks are dependent on the either increased or decreased contact area and the creation of strain gradients associated with the coating’s deformation which takes place when variations in material stiffness are encountered, that is, when going from slot to slot (PDMS to ecoflex, or ecoflex to PDMS). The higher or lower deformation depending on the stiffness of a material results yields different sensor output.

When going from ecoflex to PDMS (such as from 1 to 2 in the blue signals or from 3 to 2 in red signals in Figure 7a-b), there is a decrease in contact area due to decreased deformation fields in PDMS relative to ecoflex, thus resulting in an effective “released contact”. When going from PDMS to ecoflex (such as from 2 to 3 in blue signals or from 2 to 1 in red signals in Figure 7a-b), there is an increase in contact area due to increased deformation fields in ecoflex relative to PDMS, thus resulting in “new contact”. Due to the surface roughness of the coating, when crossing from a harder material to a softer one, the softer material is able to conform to the coating filling up all the microscale “air gaps”. TENGs which rely on changes in contact area to produce alternating current (AC) current outputs have similarly been reported. As a result of mixed triboelectric and flexoelectric charge transfer at the interface between the rough coating and the slots, we can observe that charge transfer is reversed with material strain, when both effects are in opposite directions.

Finally, when contact is released, a current peak is always created in the positive y-axis direction. This is due to contact-separation with a material which is triboelectrica lly more negative than the material of the sensor.

Figure 7f-g show 3 repetitions of the aforementioned tests, in which we can once again see the trend described. The use of tactile sensors for detection of changes in stiffness such as the ones seen in Figure 7 could be invaluable for a number of clinical applications. An example application can be using the sensorised surgical glove to facilitate accurate determination of fetal position. As a further example application, ex-vivo detection of anal sphincter defects was investigated. For these tests, the sensorised gloves were covered by a second sterile surgical glove which is routinely used in surgical practice. An obstetric Senior House Officer (SHO) tested the sensorised glove firstly ex vivo on an intact and cut pig anal sphincter (Figure 8a) and then in a whole pig cadaver (Figure 8b). This was performed by rubbing across the sphincter from side to side to demonstrate detection of the sphincter defect.

As shown in Figure 8b-c, the sensors on the index fingertip produce repeatable and distinguishable peaks when passing through the defect (the area which was cut) in the dissected sphincter, and when coming back in contact with the intact sphincter after the defect is traversed, due to the difference in stiffness and contact. For the sphincter tests on the pig cadaver, the results once again show a distinguishable peak when crossing the defect (Figure 8e-f). For the latter tests, the obstetrician always kept the sensorised glove in contact with the sphincter due to limited space for maneuverability, which is why the peaks corresponding to “contact with sphincter” and “detachment” are not observed in the results. These results suggest the same trend and behaviour of the sensors as that shown in figure 7, which served as the foundation to pursue this clinical application-based case study. In these tests, positive peaks form when the sensor makes contact with the sphincter (as opposed to negative peaks such as the ones seen in figure 7, which is due to the relative positions of the material in contact with the sensor and the sensor itself in the triboelectric series), and similarly negative peaks are formed when contact is released from the material. When rubbing from the intact sphincter to the defect (where the sensor comes in contact with the softer anal mucosa at the base of the defect) the peak formed is in the positive y-direction, due to the increased contact, deformation and strain gradients associated with it. When going back from the defect to the intact sphincter, a negative peak is formed, due to decreased deformation fields in the intact sphincter relative to the softer anal mucosa beneath the defect. Notably, in the tests carried out on the non-dissected sphincter in the whole pig cadaver, the peak produced when encountering the defect is always in the negative direction. This may seem counter-intuitive at first, based on our previous rationale since the anal mucosa beneath the sphincter defect is softer than the sphincter muscle. Our interpretation is that the aforementioned lack of space and maneuverability, together with the finer cut created in the sphincter to produce the defect, resulted in the creation of a “slot” in the sphincter and a small air gap between the sensor (together with the obstetrician’s fingertip) and the skin. Stiffness is no longer the governing factor affecting the induced current, since when the defect is crossed, the sensor is unable to fully make contact with the softer anal mucosa underneath and instead is subject to contact separation due to the small gap (“slot”) created instantaneously when rubbing through the sphincter.

The obstetrician repeated these tests 30 times for each scenario (dissected intact sphincter, dissected sphincter with a defect, non-dissected intact sphincter and non-dissected sphincter with a defect). As can be seen in the box plots in S6, the sensor achieves 100% sensitivity and 100% specificity in detecting the sphincter defect for the non-dissected pig cadaver tests. A simple user-friendly interface set up using LabView is able to detect the defect by simple specifying a threshold current level (-1.5 nA) below which the peak created corresponds to the defect (clearly visible in the boxplot in figure 9). Ultimately, the sensors produce repeatable and distinguishable peaks when going through the defect in the sphincter, leading to successful detection of the injury. Additionally, due to the thinness achieved by integrating the coating directly on to the surgical glove, clinicians’ sensory perception remained unaffected, which is critical towards its effective integration in the clinical workflow.

Based on this ex vivo sphincter case study, it is anticipated that this technology has the potential to improve detection of not only obstetric anal sphincter injury, but also a range of surgical applications in which contact/force/stiffness feedback is relevant and required. From a human-computer interface (HCI) perspective, different representations of the force/stiffness data need to be carefully evaluated and combined with appropriate/user-friendly interfaces. This may lead to a complete system comprising the sensorised glove and the interface to communicate real-time force/stiffness information to the surgeons constructively, with the aim of improving the safety and training of microsurgical tasks. Force/stiffness perception is currently highly subjective and this system could lead to objective metrics that could serve as a foundation for safe and optimized procedures and training.

A systematic investigation of the number of spray passes used to make the sensor which yields the optimal/stable triboelectric sensor output is shown in Figure 10a and 10b. This is important to ensure that changes in the stiffness level/contact area of different tissues (for example to detect anal sphincter injury or sutures in a baby’s head) are detected in a repeatable manner. To detect changes in stiffness levels, the sensors need to be responsive to deformation and deformation rate changes, and as a result need to be flexible. Furthermore, thinness must also be ensured to not alter the perception of the clinicians when wearing the sensorised glove. With the latter in mind, the coating would ideally be as thin as possible. Figure 10a shows the triboelectric current output for sensors prepared with different number of spray passes (which correspond roughly to increase in thickness).

These measurements were obtained by tapping the sensors using precision stage (see Figure 5) which ensured repeatable speed, duration and force of contact. Each box plot in Figure 10a corresponds to 30 measurements. It is clear that a minimum number of spray passes must be surpasses as the low number of passes resulted in an erratic and variable sensor output current (resulting in larger boxes). This may be due to microscopically non-uniform coating thickness and the electrodes not being fully covered by the coating (see Figure 10a). Once this threshold was surpassed, a more stable operation range is obtained, where results are consistent and the sensor is able to accurately detect changes in stiffness levels. This is important to detect slots of different materials (Figure 7), sutures in the phantoms we have developed in the lab and the anal sphincter injury detection in ex vivo pig studies (Figure 8). However, beyond a certain threshold, increase in the number of spray passes resulted in coatings that were stiff and flaky, which in turn resulted in decreasing (as is apparent from the evolution of the sample means in Figure 10a) and less stable current outputs. The reduced flexibility leads to lower contact area and deformation which seems to reduce the magnitude of the output peaks. Consequently, we have found an optimal range of spray passes (6 - 10 spray passes, the highlighted in green zone in Figure 10a) for which the triboelectric properties are most suitable for our application.

Surface roughness enhances the contact area and thus triboelectrification. As a result, parallel tests to quantify the roughness of all the samples within this range are shown. The demonstrated generality of approach, manufacturing procedure and tests were also carried out by replacing PVDF in the nanocomposite with another the hydrophobic polymer, thermoplastic polyurethane (TPU). The synthesis was carried out following analogous steps to those in the PVDF-PMMA-ZnO coating. Figure 10b covers the results obtained from the repeatable tapping test on TPU-PMMA-ZnO coatings of varying number of spray passes. Once again, an optimal threshold was found (10 - 14 spray passes) and is highlighted in green in Figure 10b. Following the same trend as the PVDF-PMMA-ZnO coating, a lower number of spray passes resulted in a more variable current output, while a larger number of spray passes away from the optimal range lead to stiffer/more flaky coatings which also generated less stable and lower current outputs. The results in Figure 10b confirm that our approach is general and the three set of functionalities, namely self-cleaning, anti-bacterial properties and optimal triboelectric properties needed to detect stiffness change relies on an optimal range of spray passes.

A minimum number of spray passes is desirable to keep thickness as low as possible. However, lower number of passes results in an erratic and variable sensor output current. This is believed to be due to microscopically non-uniform coating thickness and presence of any defects leading to a high possibility of crack formation, which when present can quickly reach the electrode. The morphology images (figures 12-16) show the presence of microclusters (see the arrow optical microscope image, figure 12a) and non-uniformity in coating thickness at low passes, even though all parts of the substrates were in fact coated (corresponding SEM images showed no holes). The exposed electrodes may touch the contacting object during a tapping test and can thereby change the sensor output in an unstable manner. This explains the larger spread in sensor output currents at low spray passes.

Once a threshold number of passes and thickness of the coating is surpassed, the current output and roughness of the coating (as shown in figure 11) stabilize and a more stable operation range is obtained. Such stable triboelectric output is important for reliable detection of changes in the stiffness levels and also to detect differences between slots of different materials, sutures in phantoms and the injuries, such as the anal sphincter injury described above (in ex vivo pig studies).

However, beyond a threshold, increasing number of spray passes results in coatings that are stiff and flaky, which again reduces the current output and makes the output unstable. The reduced flexibility of the coating (with greater thickness) together with the decrease in roughness (as shown in figure 11) leads to a lower contact area and deformations, which in turn reduces the magnitude of the current output peaks. This can also be seen in figures 12- 22, where the filling up of the “gaps” between the microclusters is clear, once 12 spray passes are surpassed. At higher number of passes, cracks appear in the coating microstructure, as shown in figures 19-22, which increases the variability of the output current signals again. The crack formation is increasingly evident as the number of spray passes increase, perhaps due to uneven drying of the thicker coating which leads to significant amounts of solvent trapped below the top layer. This in turn results in the build up of stress during a drying/annealing stage, which are known to be accompanied by substantial cracking. There is an optimal window for number of spray passes - 6 to 10 passes - for which the triboelectric output is most reliable for usage.

The estimated thickness that this optimal window corresponds to is shown in figure 23, which plots the number of spray passes against the coating thickness as measured using a 3D optical microscope for the samples shown in figures 12 to 22. In particular, to make the measurements, a metal stub was coating (with the corresponding number of spray passes) and a scraping made to determine thickness. It can be seen that there is a relationship between spray passes and coating thickness whereby each spray pass provides on average between 8 and 2 pm, with fewer passes typically resulting in a higher deposition thickness per pass.

Materials'. Acetone (ACS reagent, >99.5%, Sigma Aldrich), Advanced Dulbecco’s Modified Eagle Medium (Advanced DMEM) (Thermofisher Scientific), Antibiotic-Antimycotic (100X, Thermofisher Scientific), N,N-Dimethylformamide (DMF) (anhydrous, 99.8%, Sigma Aldrich), DMSO (sterile filtered, TOCRIS), Fetal Bovine Serum (FBS), GlutaMAXTM-1 (200 mM 100X, Thermofisher Scientific), LB Agar (Invitrogen, powder (Lennox L agar), LB broth (Invitrogen, powder (Lennox L agar, Thermo Fisher Scientific), LIVE/DEADTM Viability/Cytotoxicity Kit (L3224, Thermofisher Scientific), MTT assay reagents (CyQUANTUM MTT Cell Viability Assay kit, Thermofisher Scientific), Polyimide (PI) sheets (50 pm thickness, Sigma Aldrich), Poly(methyl methacrylate) (PMMA) (average Mw ~120,000 by GPC, Sigma Aldrich), (Poly(vinylidene fluoride) (PVDF) (average Mw ~530,000, pellets, Sigma Aldrich), Screen printable silver ink (DM-SIP-2005, Dycotec), Trypan Blue Solution (0.4%, Thermofisher Scientific), Trypsin-EDTA (0.25%, Thermofisher Scientific), Zinc oxide (nanopowder, <100 nm particle size, Sigma Aldrich) were all used as received.

It can be appreciated that the various embodiments described above contain complimentary features that may be combined depending upon the need of the user.

Other embodiments are intentionally within the scope of the invention as defined by the appended claims.

Claims

1. A superhydrophobic nanocomposite coating for a wearable device, said coating comprising: a non-toxic hydrophobic polymer for providing a dielectric effect; a non-toxic nanoparticle filler for providing surface roughness and a further dielectric effect; and a non-toxic hydrophilic polymer for providing a matrix supporting the fluoropolymer and the nanoparticle filler, wherein the coating provides a triboelectric effect when sprayed onto a substrate.

2. The coating of claim 1 , wherein the non-toxic hydrophobic polymer is a fluoropolymer, and optionally or preferably wherein the fluoropolymer is a Polyvinylidene Fluoride; or wherein the non-toxic hydrophobic polymer is a thermoplastic elastomer, and optionally or preferably wherein the thermoplastic elastomer is a thermoplastic polyurethane.

3. The coating of claim 1 or claim 2, wherein the non-toxic nanoparticle filler comprises ZnO nanoparticles.

4. The coating of any preceding claim, wherein the non-toxic hydrophilic polymer comprises an acrylate polymer, and optionally or preferably wherein the acrylate polymer comprises Polymethyl methacrylate, or poly ethyl cyanoacrylate.

5. The coating of any preceding claim, wherein the coating comprises a thickness of between 1 pm and 100 pm.

6. The coating of claim 5, wherein the thickness is between 20 pm and 60 pm.

7. A triboelectric generator comprising: a substrate having a nanocomposite coating according to any preceding claim.

8. A wearable device comprising the substrate of claim 7.

9. The device of claim 8, wherein the substrate further comprises an electrode beneath the coating.

10. A method of applying the coating of any one of claims 1 to 6, wherein the method comprises the steps of: dissolving the non-toxic hydrophobic polymer and the non-toxic hydrophilic polymer in a solvent; combining the dissolved non-toxic hydrophobic polymer, the non-toxic hydrophilic polymer and the non-toxic nanoparticle filler in a sprayable dispersion; and spraying the coating onto a substrate with a number of passes, each pass depositing a thickness of approximately 2 to 7 pm per pass, and wherein the number of passes is between 2 and 15.

11. The method according to claim 10, wherein the non-toxic hydrophobic polymer and the non-toxic hydrophilic polymer are separately dissolved in a solvent.

12. The method according to claim 10 or claim 11 , wherein the non-toxic hydrophobic polymer comprises an 8 wt.% weight stock solution.

13. The method according to any one of claims 10 to 12, wherein the non-toxic hydrophilic polymer comprises a 20 wt.% weight stock solution.

14. The method according to any one of claims 10 to 13, wherein the solvent is a polar aprotic solvent such as Dimethylformamide.

15. A sensorised glove for healthcare sensing, said glove comprising: one or more printed electrodes on fingertips of the glove; and a triboelectric coating according to any one of claims 1 to 6 deposited on the electrodes to form one or more triboelectric sensors.