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. 2022 Oct 10;7(42):37674–37682. doi: 10.1021/acsomega.2c04563

Epidermal Inspired Flexible Sensor with Buckypaper/PDMS Interfaces for Multimodal and Human Motion Monitoring Applications

Sharon J Paul †,‡, Indu Elizabeth §,*, Shubhda Srivastava ‡, Jai S Tawale ∥, Prakash Chandra †, Harish C Barshilia §, Bipin K Gupta ‡,*
PMCID: PMC9608422  PMID: 36312412

Abstract

graphic file with name ao2c04563_0009.jpg

The advancements in the areas of wearable devices and flexible electronic skin have led to the synthesis of scalable, ultrasensitive sensors to detect and differentiate multimodal stimuli and dynamic human movements. Herein, we reveal a novel architecture of an epidermal sensor fabricated by sandwiching the buckypaper between the layers of poly(dimethylsiloxane) (PDMS). This mechanically robust sensor can be conformally adhered on skin and has the perception capability to detect real-time transient human motions and the multimodal mechanical stimuli of stretching, bending, tapping, and twisting. The sensor has feasibility for real-time health monitoring as it can distinguish a wide range of human physiological activities like breathing, gulping, phonation, pulse monitoring, and finger and wrist bending. This multimodal wearable epidermal sensor possesses an ultrahigh gauge factor (GF) of 9178 with a large stretchability of 56%, significant durability for 5000 stretching–releasing cycles, and a fast response/recovery time of 59/88 ms. We anticipate that this novel, simple, and scalable design of a sensor with outstanding features will pave a new way to consummate the requirements of wearable electronics, flexible touch sensors, and electronic skin.

Introduction

E-skin sensors are a prominent research area for novel stimuli sensing and monitoring applications.1−7 The essential requisite for an e-skin sensor is that it should be a versatile strain sensor that is highly sensitive to detect subtle pressure signals and large-scale deformations. These pressure signals are quantified through the corresponding electrical signals. Various studies have been dedicated to fabrication of reliable, multifunctional sensors based on carbon nanomaterials.8−13 Carbon nanotubes (CNTs) have attracted great interest as effective nanoscale reinforcements in the polymer matrix for high-end applications as flexible strain sensors owing to their excellent mechanical properties, and high electrical and thermal conductivities.9,14−17 CNTs when combined with thermoplastics, thermosets, and elastomers as polymer composites serve as a vital group of relatively inexpensive materials with peculiar engineered architecture for prospects as strain sensors. The CNTs when incorporated in the polymer matrix construct a network that allows the transition of insulating material to semiconductive or conductive material. This phenomenon is referred as the electrical percolation threshold, where the conductive pathways form at a critical filler concentration within the insulating polymer matrix.

The successful employment of CNTs in wearable pressure sensors depends on their homogeneous dispersion in the polymer matrix. CNTs have a tendency to agglomerate in the course of synthesis due to the van der Waals attraction between them. Also, the high aspect ratio (>1000) substantially increases the viscosity of the polymer, thereby making the processability of CNT-based composites a tedious task and affecting their dispersion process. The low dispersion of entangled nanotubes hinders the full potential of CNTs as reinforcements due to the poor interfacial interaction between the polymer matrix and the nanofillers.18−21 Also, it becomes difficult to produce void-free nanocomposites when the CNT concentration is higher than 1.0 wt %, as the solution viscosity increases manifolds. Therefore, the electrical percolation threshold in CNT-reinforced polymer nanocomposites is generally restricted to levels lower than 5.0 wt % of the CNTs in the polymer matrix.22 Therefore, the low solubility of CNTs in solvents, large agglomeration tendency, high viscosity, and poor dispersion of CNT/polymer composites limit their practical applications. These aforementioned issues can be resolved through incorporation of buckypapers (BPs) into the polymer matrix.

BP can structurally be defined as a free-standing porous mat of entangled CNTs that are cohesively bound by van der Waals interactions.23−26 BPs are promising materials employed to generate new nanostructured composites with outstanding mechanical features for diverse applications in the fields of artificial muscles,27 wearable electronics,28 and e-skin sensors.29,30 The key differences between conventional CNT/polymer composites and polymer composites with infused BPs are associated with the manufacturing process, bundle distribution, and carbon content. The CNT/polymer composites are synthesized via mixing solution,31 in situ polymerization,32 or melt mixing.33 The nanotubes are dispersed randomly throughout the polymer matrix without forming any network and the CNT loading is generally lower than 5.0 wt %. On the other hand, BP-infused composites are fabricated via intercalation,34 hot compression,35 or electro-spinning.36 Also, BPs have a high CNT content of up to 60 wt %37 and CNTs form a dense network with uniform dispersion, which acts as a skeleton for improved transfer of stress, phonons, and electrons across the CNT arrangements, resulting in better mechanical, thermal, and electrical properties of the nanocomposites. As compared to other CNT-derived structures like hydrogels, aerogels, vertically aligned CNTs, and horizontally aligned CNTs, the buckypapers are expected to show multidirectional strain-sensing ability. The sensivities of other structures are more in one direction, showing sensitivity either under compressive strain or under tensile strain, which limits their applications. But, because the BPs are thin films of CNTs, they show multidirectional sensing capabilities under any kind of strain. Recently, researchers have been engaged in developing material with all-directional sensing and BPs allow efficient detection of a variety of forces beyond uniaxial strain, such as compressive, shearing, torsion forces, etc., making them a promising material in wearable devices. These BPs are scalable and can be reproduced on a large scale unlike the other CNT-derived structures.

The mechanical properties of BPs increase manifolds when they are infused in a polymer matrix. Han et al.38 reported that BP/polyurethane composites with 46 vol % of multiwalled carbon nanotube (MWCNT) loading show substantial improvement by 340 and 960% in Young’s modulus and tensile strength, respectively. Pham et al.37 show similar results in that the Young’s modulus and tensile strength are increased by 120 and 200%, respectively, for BP/polycarbonate composites. Several reports are available on conventional CNT composites, but only a few reports are available on the potential of the BP/poly(dimethylsiloxane) (PDMS) composite as a flexible strain sensor designed for monitoring human motions. PDMS was selected as the polymer matrix because of its better mechanical properties as compared to other flexible polymer matrices as investigated in our previous study.39 Also, PDMS is an established biocompatible polymer highly recommended for use in the biomedical industry. It has been used in developing contact lenses to medical devices,40 self-healing scaffold materials,41 therapeutic,42 and sensor applications.43 PDMS is skin friendly and the developed sensor can be directly adhered on skin without any reactivity concerns.

Herein, we report an exclusive study on the BP/PDMS composite for multimodal and human motion monitoring. The epidermal sensor is exotically designed as a buckypaper PDMS sandwich structure, where BP is embedded between the layers of the PDMS matrix. The fast response time, cyclic stability, large stretchability, and high gauge factor of the sensor is one of the best combinations reported for CNT-based wearable sensors (Comparison Table S1, Supporting Information). These properties of the sensor are expected to constitute a promising troop with healthcare-oriented electronic functionalities.

Results and Discussion

The BP/PDMS epidermal sensor was fabricated by sandwiching the as-synthesized buckypaper between layers of the PDMS polymer matrix as schematically illustrated in Figure 1a. The epidermal sensor has been fabricated with the dimensions of ∼1.5 mm height, 1.3 cm width, and 5.7 cm length. The scanning electron microscopy (SEM) image of the buckypaper in Figure 1b displays a densely packed entangled network of randomly oriented MWCNTs without any agglomeration. This orientation gives rise to the multidirectional strain-sensing abilities and isotropic properties of the sensor. The SEM image in Figure 1c clearly reveals the sandwich structure of the epidermal sensor where buckypaper is infused between two layers of the PDMS matrix. Figure 1d is the magnified SEM image of the buckypaper sandwiched section that indicates infiltration of PDMS between the interspaces of the entangled MWCNTs in the buckypaper. Therefore, SEM images were vital to conclude that the buckypaper is infiltrated as well as sandwiched within the PDMS matrix.

Figure 1.

Figure 1

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(a) Schematic representation of the fabrication of the buckypaper and buckypaper PDMS sandwich structure epidermal sensor; (b) SEM image of the as-synthesized buckypaper; (c) SEM image of the buckypaper PDMS sandwiched structure; and (d) SEM image of the buckypaper section that is sandwiched in the PDMS matrix.

Raman studies were also conducted to indicate the sample quality and verify the homogeneity of the composite after embedding BP in between the layers of the PDMS. Both BP and the BP/PDMS composite exhibit the characteristic D-band peak at around 1340 cm–1 that reflects the disorder-induced carbon atoms arising from the defects in CNTs and the G-band peak at around 1570 cm–1 that corresponds to the p-orbital structural intensity of the sp2-hybridized carbon atoms of CNTs, respectively, as shown in Figure S1.

The electromechanical attributes of the sensor were characterized by stretching the sensor and concurrently measuring the resistance across the sensor. The sensor endows tolerable strain up to 56% without any mechanical deformation, after which the sensor starts to develop cracks. The sheet resistance calculated under zero strain was ∼78 Ω. As illustrated in Figure 2a, the resistance increases gradually in the initial loading, and then a sudden rise in resistance is observed. This mechano-electric sensing behavior can be ascribed to the nanotube–polymer interface. Initially, the CNT network is dense and intact, thereby improving the interfacial adhesion between them; thus, CNTs effectively propagate conduction throughout the sensor. As the sensor is further stretched, it causes reduction in the number of contact points between CNTs, which considerably limits the charge transfer and thereby increases the resistance. The proposed electromechanical sensing mechanism of the sensor is depicted in Figure S2.

Figure 2.

Figure 2

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(a) Relative resistance change of the sensor up to 56% strain; (b) gauge factor of the sensor as a function of stretchability; (c) electromechanical stability of the sensor for 5000 sequential tensile cycles at 24% strain; and (d) light-emitting diode (LED) response while stretching the sensor.

As a result, the resistance of the sensor decreases and the gauge factor increases manifolds, i.e., ∼1118 at 24%, ∼2160 at 32%, ∼2882 at 40%, and ∼2963 at 48% strain. Further, in the third sensing stage, while applying the strain up to 56%, the gauge factor is consequently affected by the tunnelling resistance and contact resistance of the CNTs. The rearrangement of the CNT network and consequent deformation of the visco–elastic polymer matrix tremendously increase the bulk resistance of the sensor and lead to huge enhancements in the gauge factor, i.e., ∼9178 at 56% strain. For the gauge factor calculation, see the Supporting Information file. The excellent durability and stability of the sensor was validated for 5000 stretching–releasing cycles over 24% strain as elucidated in Figure 2c. The identical peaks of the relative resistance response thereby signify a long lifetime of the sensor. Then, the sensor was connected to LED and changes in the light intensity pattern were observed as demonstrated in Figure 2d. As the sensor is stretched, the entangled CNTs move apart, thereby increasing the resistance. As the resistance increases, a decrease in the light intensity of LED is observed. On further stretching, the conductive network breaks further, thereby further increasing the resistance and reducing the light intensity of LED. The light intensity is restored when the sensor is released back to the initial length (see Supporting Movie 1: LED).

The connection state of CNT networks within the PDMS can clearly be seen in the SEM image of the BP/PDMS composite in Figure 3a. When the composite is stretched up to ∼24%, the CNT network gets disturbed but CNTs are still entangled, thereby maintaining the conductivity of the composite as seen in Figure 3b. While stretching the sample up to ∼56%, few CNTs still cling to each other and the connections do not break completely, as seen in Figure 3c. However, after further stretching the composite, the conducting network breaks, as seen in Figure 3d. The microcracks of the composite after being stretched at different magnifications are shown in SEM and the optical images are shown in Figure S3.

Figure 3.

Figure 3

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SEM image of the (a) CNT network within PDMS under no strain; (b) CNT network within PDMS when stretched up to ∼24%; (c) CNT network within PDMS when stretched up to ∼56%; (d) CNT network within PDMS when stretched beyond 56%.

The typical current voltage (I–V) characteristics at different strain levels are shown in Figure 4a. The linear I–V curve for all strain ranges is indicative of the ohmic behavior of the strain sensor. Both static and dynamic strain responses were measured to expound the viability of the sensor. For the dynamic strain response the sensor was stretched and released swiftly, whereas for the static strain response the sensor was stretched and held under that strain for a few minutes and then released. The sensor was also tested for stair-type dynamics and for static strains ranging from 8 to 56%. As shown in Figure 4b, the sensor exhibits excellent dynamic stair-type resistance change response, where similar electrical responses are observed under the same strain level over the stretching and releasing processes.

Figure 4.

Figure 4

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(a) I–V curve of the sensor under different strains; (b) dynamic stair-type resistance response of the sensor; (c) relative resistance change of the strain sensor at repeated dynamic strains of 8, 16, 24, 32, 40, 48, and 56%; (d) static stair-type resistance response of the sensor; (e) relative resistance change of the strain sensor at repeated static strains of 8, 16, 24, 32, 40, 48, and 56%; and (f) dynamic response (59 ms) and recovery (88 ms) time of the sensor.

As shown in Figure 4c, the relative resistance change was measured in the large repeated strain regions (8, 16, 24, 32, 40, 48, and 56%). The synchronized electrical responses with identical peaks are worth noting. The sensor also reveals stable performance during the stretching and releasing processes of the static stair-type resistance change response as presented in Figure 4d. Further, we also estimated the static resistance response under repeated strains ranging from 8 to 56% as depicted in Figure 4e. Notably, the intensity and shape of each signal under a particular strain are identifiable. We investigated the response and recovery time of the sensor from the tapping motion as shown in Figure 4f. The sensor displays an instant response time of 59 ms and a rapid recovery time of 88 ms. The response time is the estimated time from the applied stress to the time taken by the sensor to reach the maxima of the electromechanical response, signifying its steady state. The recovery time is the estimated time taken by the sensor to recover by 95%.

This mechanically robust sensor can be conformally adhered on skin and has the perception capability to detect real-time transient human motions and multimodal mechanical stimuli of stretching, bending, tapping, and twisting. This wearable sensor can detect subtle human motions such as phonation, gulping, breathing, pulse monitoring, and vigorous human motions, such as motion of human joints. The sensor accurately detects the regular pulses on the human wrist of a healthy person, as shown in Figure 5a. The corresponding pulse rate calculated for a healthy individual (32 years old, 165 cm height, 64 kg) is 86 beats/min in relaxed mode. The systolic (P1) and diastolic (P2) pressure peaks with different amplitudes are easily discernible, as shown in the inset of Figure 5a. The transit time (ΔT) was calculated as 0.2 s from the time delay between the P1 and P2 peaks. In order to determine the feasibility of the sensor for distinguishing the heart rates of the individual in relaxed state and after exercise, the sensor was affixed on the wrist of the same individual after performing exercise. The pulse waveform can be seen in Figure S4. The pulse throbbing after exercise was measured as 96 beats/min. The results show that the sensor can accurately determine different waveforms before and after exercise along with different amplitudes of the pulse waveforms. Therefore, the sensor is capable of pulmonary and cardiac function analysis. To further monitor small-scale human motions, the wearable sensor was attached to the throat of a healthy person directly to record the gulping and phonation patterns. The phonation of every syllable in vowels can be detected and easily distinguished as represented in Figure 5b. The throat muscles produce different signals when gulping motion is done, as shown in Figure 5c, thereby revealing that the sensor has the proficiency to identify laryngeal (phonation) and esophagus (gulping) muscle movements. Periodic respiration changes were also recorded by fastening the sensor on a mask and consecutively monitoring the change in electrical resistance of the strain sensor as shown in Figure 5d. The waveforms generated while breathing are highly consistent. Additionally, the difference in relative resistance changes for normal and deep/heavy breathing shows a raised resistance response, as is clearly visible in Supporting Movie 2: Breathing. The sensor was also used to monitor the complicated epidermal/muscle movements for vigorous motions.

Figure 5.

Figure 5

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Strain detection of subtle human motions during (a) real-time pulse monitoring; (b) phonation (pronouncing vowels); (c) gulping; and (d) normal and fast breathing.

This strain sensor was fixed firmly on the finger joint to detect the movement of the finger as shown in Figure 6a. An increase in relative resistance change and different response curves are observed when the finger is bent at different angles (Figure 6b). The real-time changes in the signals can be clearly seen in Supporting Movie 3: Finger Bending. The sensor was also well mounted on the wrist joint to monitor the wrist movement and the real-time signals were recorded as illustrated in Figure 6c. The strain sensor was affixed on the heel of a person to detect signals while walking (Figure 6d). The sensor directly measures the force applied by the periodic foot stepping and transduces it into variations in the electrical resistance. The sensor generates distinguishable and reliable signals for all human motions.

Figure 6.

Figure 6

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Strain detection of vigorous human motions. (a) Sensor adhered to the index finger for real-time sensing of finger bending; (b) relative resistance response of finger bending at different angles; (c) sensor mounted on the wrist for real-time sensing of wrist bending; and (d) relative resistance response while walking.

The sensor also shows improved strain-sensing abilities as it can monitor more complicated multidirectional responses like the mechanical stimuli of bending, twisting, rolling, and tapping. The electrical response while manually bending and releasing the sensor is explicitly shown in Figure 7a. The transient resistance features were also observed while rolling (Figure 7b) and twisting the sensor (Figure 7c).

Figure 7.

Figure 7

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Mechanical stimuli strain detection while (a) bending; (b) rolling; (c) twisting; and (d) gently tapping the BP/PDMS sensor.

The sensor can also identify subtle tapping as represented in Figure 7d. The sensor remarkably generates distinct electrical signals while bending, rolling, twisting, and tapping.

Conclusions

A sandwich structure-based buckypaper/PDMS strain sensor has been fabricated through an amenable and cost-effective synthesis route. The sensor possessed a tolerable range of 56% and exhibited high sensitivity with a gauge factor (GF) of 9178. The superior sensitivity was the result of the buckypaper-induced architecture, as proved through both theoretical analysis and experimental tests. The sensor depicted good durability over 5000 stretching–releasing cycles and a rapid response time of 59 ms and recovery time of 88 ms. The smart sensing functionalities with real-time monitoring of human motions and multimodal stimuli response represent the complete physiological sufficiency of the sensor. The different waveforms generated through the different subtle-vigorous muscle movements and mechanical stimuli help in application of the sensor as a real-time health monitor to distinguish humans’ physiological activities. Further, this sensor is capable of generating breakthrough as a versatile sensing module with tremendous scope in the modern-life-oriented trend of wearable electronic applications.

Methods

Materials

The poly(dimethylsiloxane) (PDMS) pre-mixture (Dow SYLGARD 184 silicone elastomer) and curing agent were used in the preparation of composites. MWCNTs used were synthesized according to our previous reported work.44 Ethylene glycol used was AR grade and was purchased from E-Merck.

Synthesis of Buckypaper

Free-standing BP was formulated using the known vacuum filtration-assisted method26,45,46 through a combination of three steps shown schematically in Figure 1. First, a small amount of MWCNTs was homogenized in ethylene glycol. Ethylene glycol was used as a solvent as it has low viscosity and does not affect the properties of MWCNTs. Ultrasonication was used for dispersion of MWCNTs in ethylene glycol. Ultrasonication generates shock waves that promote “peeling-off” of individual nanotubes and prevent the agglomeration of CNT bundles. Second, the homogeneous dispersion of MWCNTs in ethylene glycol was filtered through a Millipore filter paper with 0.22 μm pore size using a vacuum-assisted filtration assembly. In the third step, the MWCNTs that were filtered got deposited on the filter surface in the form of a thin membrane. This thin membrane was dried in an air oven at 60 °C for 24 h, after which free-standing buckypaper was obtained.

Synthesis of the Buckypaper/PDMS Epidermal Sensor

The epidermal sensor was fabricated by sandwiching the as-synthesized buckypaper within PDMS. First, the BP was cut into the desired dimensions with the help of scissors, and the copper tape was fastened on either side of the BP to establish contacts. Second, the PDMS base and curing agent were mixed thoroughly in a specific ratio (10:1) with the help of a cyclo mixer, for homogeneous mixing. Half of the mixture was poured slowly into a Teflon mould and allowed to stand for a few minutes. Thereafter, the BP was carefully kept over it without disturbing the contacts and then the remaining PDMS mixture was poured over it. This not only allowed infiltration of the polymer into the interstitial spaces between the CNT network of the BP, but also gave a firm coating of PDMS on either side of the BP, thereby enhancing the robustness of the sensor.

Characterization

The morphological examination of samples was done with the help of scanning electron microscopy (SEM), LEO435-VF. Raman spectroscopy was performed using a 532 nm laser wavelength over the 400–4000 cm–1 scan range. The continuity of conduction across the CNT network in the BP/PDMS epidermal sensor and the sheet resistance of the sensor assembly were confirmed by the two-point probe method. Electrical resistance across the BP/PDMS sensor was measured using Agilent B2901. The clamps of the multimeter were connected to the copper tape of the sensor assembly, which permitted data acquisition on a computer. All electrical resistance measurements were accomplished at a constant voltage (1 V).

Acknowledgments

The authors thank Director, CSIR-NPL and Director, CSIR-NAL for providing the platform to pursue our practical work at their respective labs. S.J.P. would also like to acknowledge DST for providing the funding under INSPIRE-Fellowship Scheme (IF 160064). I.E. acknowledges SERB for funding the work under the Start-up Research Grant (SRG/2019/001168).

Supporting Information Available

The Supporting Information is available free of charge at https://pubs.acs.org/doi/10.1021/acsomega.2c04563.

  • Comparison of CNT-based wearable sensors reported in literature (Table S1); Raman spectra of PDMS, BP, and BP/PDMS epidermal sensor (Figure S1); mechano-electric sensing mechanism of BP/PDMS epidermal sensor while stretching (Figure S2); SEM images of microcracks of BP/PDMS epidermal sensor at different magnification (a) 100 μm; (b) 50 μm, and (c, d) optical images of microcracks of BP/PDMS epidermal sensor (Figure S3); real-time pulse monitoring after exercise (Figure S4); and gauge factor calculation (PDF)

  • The light intensity is restored when the sensor is released back to the initial length (MOV)

  • The difference in relative resistance changes for normal and deep/heavy breathing show raised resistance response (MP4)

  • The real-time changes in the signals while finger bending (MP4)

Author Contributions

S.J.P. proposed the project, designed the experiment, performed all characterizations, and wrote the manuscript. S.J.P. and I.E. performed the synthesis of buckypaper and its nanocomposites. S.S. performed optical studies. J.S.T. carried out SEM. I.E. and B.K.G. supervised all stages of the research. P.C. encouraged S.J.P. to carry out the work at CSIR-NPL. H.C.B. reviewed the manuscript.

The authors declare no competing financial interest.

Supplementary Material

ao2c04563_si_001.pdf (695.3KB, pdf)
ao2c04563_si_002.mov (6.7MB, mov)
ao2c04563_si_003.mp4 (6.2MB, mp4)
ao2c04563_si_004.mp4 (8.7MB, mp4)

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ao2c04563_si_001.pdf (695.3KB, pdf)
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